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ELECTROCH

EMICA

L DN

A BIOSEN

SORS

ELECTROCHEMICAL DNA

B I O S E N S O R S

Edited by

Mehmet Ozsoz

Ozsoz

Mehmet Sengun Ozsoz is a professor of analytical chemistry in the Faculty of Pharmacy at Ege University and also teaches biosensor technology courses in the Biotechnology Department at Izmir Institute of Technology. Prof. Ozsoz holds a BS in chemical engineering from Middle East Technical University, Ankara, Turkey, and a PhD in analytical chemistry from the Faculty of Pharmacy, Ege University, Izmir, Turkey. He was a postdoctoral fellow with Dr Joseph Wang at New Mexico State University, Las Cruces, between 1989–1991 and 1996–1997. He is a recipient of

the 2008 Scientific and Technological Research Council of Turkey (TUBITAK) science award. Prof. Ozsoz conducts well-recognized international work on electrochemical DNA biosensors.

“The marriage of natural and synthetic nanotechnology in electrochemical DNA sensors is

a fascinating object of research. The reader gets an easy access to the complex matter by

the well-written introductory chapter. This volume builds a bridge from molecular biology

to the applications in medical diagnostics and microbiology.”

Prof. Frieder SchellerUniversität Potsdam, Germany

“This book is a very welcome contribution to the literature of electrochemical DNA

biosensors. It offers extremely useful insights into this exciting and important field.”

Dr. Joseph WangUniversity of California, San Diego, USA

This book focuses on the electrochemical applications of DNA in various areas, from basic

principles to the most recent discoveries. It comprises theoretical and experimental analyses

of various properties of nucleic acids, research methods, and some promising applications.

The topics discussed in the book include electrochemical detection of DNA

hybridization based on latex/gold nanoparticles and nanotubes; nanomaterial-

based electrochemical DNA detection; electrochemical detection of microorganism-

based DNA biosensor; gold nanoparticle-based electrochemical DNA biosensors;

electrochemical detection of the aptamer–target interaction; nanoparticle-induced

catalysis for DNA biosensing; basic terms regarding electrochemical DNA (nucleic

acids) biosensors; screen-printed electrodes for electrochemical DNA detection;

application of field-effect transistors to label-free electrical DNA biosensor arrays;

and electrochemical detection of nucleic acids using branched DNA amplifiers.

ELECTROCHEMICAL DNA

B I O S E N S O R S

This page intentionally left blankThis page intentionally left blank

Edited by

Mehmet Ozsoz

ELECTROCHEMICAL DNA

B I O S E N S O R S

CRC PressTaylor & Francis Group6000 Broken Sound Parkway NW, Suite 300Boca Raton, FL 33487-2742

© 2012 by Taylor & Francis Group, LLCCRC Press is an imprint of Taylor & Francis Group, an Informa business

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March 14, 2012 18:57 PSP Book - 9in x 6in 00-Ozsoz–prelims

Contents

Preface xvii

1 Terminology Related to Electrochemical DNA-BasedBiosensors 1Jan Labuda1.1 Introduction 1

1.2 Detection Features of DNA-Based Biosensors 3

1.3 Detection of Specific DNA Interactions 7

1.3.1 DNA Hybridization Biosensors 7

1.3.2 DNA Damage 9

1.3.3 DNA Association Interactions 13

1.3.3.1 Binding of low molecular mass

compounds 13

1.3.3.2 Binding of proteins 14

1.4 Conclusions 15

2 Electrochemical Aptamer-Based Biosensors 29S. Centi, S. Tombelli, and M. Mascini

2.1 Introduction 29

2.2 Electrochemical Detection Strategies

Based on Labeling 31

2.3 Electrochemical Aptasensors Based on

a Sandwich Assay 32


a Competitive Assay 34

2.5 Electrochemical Aptasensors Based on a Direct

Assay 37

2.6 Electrochemical Metal Nanoparticle-Labeled

Aptasensors 39


vi Contents

2.7 Electrochemical Aptasensors Based on Noncovalent

Redox Species Label 43

2.8 Electrochemical Aptasensors Based on the Aptamer

Conformational Changes 46


Target-Induced Aptamer Displacement 49

2.10 Conclusions 52

3 Carbon-Polymer Bio-Nano-Composite Electrodes forElectrochemical Genosensing 57Marıa Isabel Pividori and Salvador Alegret

3.1 Introduction 57

3.2 Composites Materials: Main Features and

Classification 61

3.3 Carbon Composites 65

3.3.1 Carbon-Based Materials as Conductive

Fillers in Composites 65

3.3.2 Rigid Carbon-Polymer Composite 69

3.3.3 Graphite-Epoxy Composites 71

3.4 Electrochemical Genosensing Based on

Graphite-Epoxy Composite 73

3.4.1 Electrochemical Genosensing Based on DNA

Dry Adsorption on GEC as Electrochemical

Transducer 73

3.4.2 Electrochemical Genosensing Based on DNA

Wet Adsorption on GEC as Electrochemical

Transducer 77

3.4.3 Electrochemical Genosensing Based on

Graphite-Epoxy Biocomposite Modified with

Avidin (Av-GEB) as Electrochemical

Transducer 78


Magnetic Beads and m-GEC Electrochemical

Transducer 81


Graphite-Epoxy Composite Modified with

Gold Nanoparticles (NanoAu-GEC) as

Electrochemical Transducer 87

3.5 Final Remarks 93


Contents vii

4 Gold Nanoparticle-Based Electrochemical DNABiosensors 103Marıa Pedrero, Paloma Yanez-Sedeno,and Jose M. Pingarron

4.1 Introduction 103

4.2 Configurations Used for DNA Immobilization 107

4.2.1 Au-Thiol Binding 108

4.2.2 Gold Nanoparticles: Metallic Oxide

Composites 110

4.2.3 Carbon Nanotube–Gold Nanoparticle

Hybrids 111

4.2.4 Polymer–Gold Nanoparticle Hybrids 111

4.2.5 Avidin–Biotin Affinity Reactions 113

4.3 Signal Transduction and Amplification Strategies 114

4.3.1 Detection Strategies Not Involving Direct

Participation of Au-NPs in the Generation

of the Electrochemical Signal 114

4.3.1.1 Direct detection of redox markers 115

4.3.1.2 Detection based on enzymatic labels 116

4.3.1.3 Detection based on electrochemical

labels intercalated within dsDNA 118

4.3.1.4 Detection involving the use of

Au-NPs as carriers 120

4.3.2 Detection Strategies Involving Direct

Participation of Au-NPs in the Generation of

the Electrochemical Signal 124

4.3.2.1 Detection based on Au-NPs’ acidic or

electrochemical dissolving 124

4.3.2.2 Label-free electrical detection 127

4.3.2.3 Signal enhancement methods 129

4.4 Conclusions and Outlook 136

5 Nanoparticle-Induced Catalysis for ElectrochemicalDNA Biosensors 141Marisa Maltez-da Costa, Alfredo de la Escosura-Muniz, andArben Merkoci


5.2 Catalysis Induced by Gold Nanoparticles 145


viii Contents

5.2.1 Electrocatalytic Activity of Gold Nanoparticle

Labels on Silver Deposition 145

5.2.2 Electrocatalytic Activity of Gold Nanoparticle

Labels on Other Reactions 146

5.2.3 Electrocatalytic Activity of Gold

Nanoparticles Used as Modifiers of

Electrotransducer Surfaces 149

5.3 Catalysis Induced by Platinum and

Palladium Nanoparticles 149

5.3.1 Electrocatalytic Activity of Platinum

Nanoparticle Labels 149

5.3.2 Electrocatalytic Activity of Palladium


5.4 Catalysis Induced by Other Nanoparticles 152

5.4.1 Electrocatalytic Activity of Titanium Dioxide


5.4.2 Electrocatalytic Activity of Osmium Oxide


5.4.3 Electrocatalytic Activity of Other

Nanoparticles 156

5.5 Conclusions 157

6 Application of Field-Effect Transistors to Label-FreeElectrical DNA Biosensor Arrays 163Peng Li, Piero Migliorato, and Pedro Estrela


6.2 Field-Effect Transistors 165

6.2.1 Field-Effect Transistor Technologies 168

6.2.1.1 Single crystalline silicon and CMOS 168

6.2.1.2 Thin-film transistors 170

6.2.2 Field-Effect Transistor Arrays 173

6.3 Field-Effect DNA Sensing 174

6.3.1 Physical Mechanisms of Detection 176

6.3.1.1 Description of the electrochemical

system 177

6.3.1.2 DNA charge fraction 178

6.3.1.3 Quantitation of the field-effect device

signal 180


Contents ix

6.3.1.4 Equivalent electrical circuit model of

functionalized FET 183

6.3.2 Differential OCP Measurement 184

6.4 Electrochemical Impedance Spectroscopy 185

6.4.1 PNA-Based Sensing 188

6.4.2 Modeling of the Signal 189

6.5 Application of FETs on Biosensor Arrays 192

6.5.1 FET-Addressed Biosensor Arrays 192

6.5.2 Specifications of the Biosensor Arrays 194

6.5.3 Development of Biosensor Arrays Based on

FETs 197

6.5.4 Fabrication Technologies and Future

Trends 198

6.6 Conclusions 201

7 Electrochemical Detection of Basepair Mismatches inDNA Films 205Piotr Michal Diakowski, Mohtashim Shamsi,and Heinz-Bernhard Kraatz


7.2 Surface Immobilization 207

7.2.1 Covalent Attachment 208

7.2.2 Adsorption 209

7.2.3 Affinity Binding 210

7.3 Detection Strategies 210

7.3.1 Direct DNA Electrochemistry 211

7.3.2 Charge Transduction Through DNA 212

7.3.3 Hybridization Indicators, Intercalators and

Groove Binders 216

7.3.4 Peptide Nucleic Acids (PNA) 221

7.3.5 Protein Mediated DNA Biosensors 225

7.3.6 DNA Stem-Loops 227

7.3.6.1 Enzyme-mediated sensors 228

7.3.7 Nanoparticle-Based Sensors 231

7.3.8 Metal-Ion Amplified Sensor 233

7.3.9 Miscellaneous Methods 236

7.4 Conclusion 239


x Contents

8 Electrochemical Detection of DNA Hybridization: Useof Latex to Construct Metal-Nanoparticle Labels 245Mithran Somasundrum and Werasak Surareungchai


8.2 Synthesis of Metal Nanoparticles 246

8.3 Use of Metal Nanoparticles as Electrochemical

Labels 249

8.4 Voltammetric Detection of Metal-Nanoparticle

Labels 254

8.4.1 Principles of Analytical Voltammetry 254

8.4.2 Anodic Stripping Voltammetry (ASV) 256

8.4.3 Quantification 258

8.4.3.1 Linear sweep voltammetry 258

8.4.3.2 Differential pulse voltammetry 260

8.4.3.3 Potentiometric stripping analysis 262

8.5 Latex as a Label Support 262

8.5.1 Introduction 262

8.5.2 Latex Synthesis 263

8.5.3 Latex Solution Properties 264

8.5.4 Layer-by-Layer Deposition: Theory 265

8.5.5 Layer-by-Layer Modification of Latex 267

8.5.5.1 Latex surface charge excess 267

8.6 DNA Measurement 278

8.6.1 DNA Immobilization 278

8.6.2 Probe Attachment 280

8.6.3 Detection Sequence 280

8.7 Areas for Further Research 284

9 Screen-Printed Electrodes for Electrochemical DNADetection 291Graciela Martınez-Paredes, Marıa Begona Gonzalez-Garcıa,and Agustın Costa-Garcıa


9.2 Fabrication of Screen-Printed Electrodes 292

9.2.1 Types of Screen-Printed Electrodes 293

9.3 Genosensors on Screen-Printed Electrodes 294

9.3.1 Electrochemical Detection of Hybridization

Reaction 295


Contents xi

9.3.1.1 Direct transduction methods 295

9.3.1.2 Indirect transduction methods 296

9.3.2 Strategies for Immobilization of ssDNA over

SPEs 298

9.3.2.1 Immobilization of ssDNA over

carbon electrodes 300

9.3.2.2 Immobilization of ssDNA over gold

electrodes 302

9.4 Applications 303

9.4.1 Enzymatic Genosensors on

Streptavidin-Modified Screen-Printed Carbon

Electrode 304

9.4.1.1 Genosensor design 305

9.4.1.2 Analytical signal recording 306

9.4.2 Alkaline Phosphatase-Catalyzed Silver

Deposition for Electrochemical Detection 310


9.4.2.2 Results 312

9.4.3 Genosensor for SARS Virus Detection Based

on Gold Nanostructured Screen-Printed

Carbon Electrode 314

9.4.3.1 Gold nanostructuration of

screen-printed carbon electrodes 315


9.4.3.3 Results 315

9.4.4 Simultaneous Detection of Streptococcus and

Mycoplasma Pneumoniae Using

Gold-Modified SPCEs 318


9.4.4.2 Results 320

9.5 Conclusion 321

10 Synthetic Polymers for Electrochemical DNABiosensors 329Adriana Ferancova and Katarına Benıkova10.1 Introduction 329

10.2 Modification of Electrode Surface with Polymers 330

10.2.1 Solvent Casting 330


xii Contents

10.2.2 Spin Coating 330

10.2.3 Electropolymerization 331

10.3 Polymer-Assisted DNA Immobilization 332

10.3.1 Immobilization of DNA onto

Polymer-Modified Electrode Surface 332

10.3.2 Immobilization of DNA Within a Polymeric

Matrix by Electropolymerization 334

10.4 Application of Synthetic Polymers in DNA

Biosensors 334

10.4.1 Electronically (Intrinsically) Conducting

Polymers 334

10.4.1.1 Polypyrroles 335

10.4.1.2 Polyaniline 339

10.4.1.3 Polythiophene and its

derivatives 341

10.4.2 Redox Polymers 342

10.4.2.1 Quinone-containing polymers 342

10.4.2.2 Redox-active polymers

containing organometalic

redox center 343

10.4.3 Nonconducting Polymers 344

10.5 Conclusions 346

11 Electrochemical Transducer for OligonucleotideBiosensor Based on the Elimination and AdsorptiveTransfer Techniques 355Libuse Trnkova, Frantisek Jelen, and Mehmet Ozsoz11.1 Introduction 355

11.2 Theoretical Fundamentals of Elimination

Voltammetry with Linear Scan (EVLS) 356

11.2.1 Elimination Functions 356

11.2.2 EVLS of Adsorbed Species 359

11.2.3 Single and Double Mode of EVLS 360

11.3 EVLS Increasing the Transducer Potential

Range 362

11.4 EVLS in Connection with Adsorptive Stripping

Technique 362

11.4.1 AdS EVLS of Homo- and

Hetero-oligonucleotides 364


Contents xiii

11.4.2 AdS EVLS of Hairpins 366

11.5 EVLS of Nucleobases and Oligonucleotides in the

Presence of Copper Ions 368

11.5.1 Mercury and Mercury-Modified

Electrodes 368

11.5.2 Solid Electrodes 371


12 Electrochemical DNA Biosensors for Detection ofCompound-DNA Interactions 379D. Ozkan-Ariksoysal, P. Kara, and M. Ozsoz12.1 Introduction 380

12.1.1 Aim of Electrochemical DNA Biosensors 380

12.2 The Structure of DNA 380

12.3 Natural Electronalytical Characterictics of DNA 383

12.4 Types of DNA Immobilization Methodologies onto

Sensor Surfaces 385

12.4.1 Adsorption (Wet Adsorption/Electrostatic

Accumulation) 386

12.4.2 Covalent Binding to Activated/

Nonactivated Surfaces 386

12.4.3 DNA Immobilization onto Transducer

Surfaces Via Avidin-Biotin Interaction 387

12.5 DNA-Compound Interactions 387

12.5.1 Types of Molecular Binding to DNA 388

12.5.1.1 Electrostatic interactions 388

12.5.1.2 Groove binding interactions 388

12.5.1.3 Intercalation mode 389

12.5.1.4 Specific binding for

single-stranded DNA 390

12.5.2 Detection Techniques for Compound-DNA

Binding Reactions Using Electrochemical

DNA Biosensors 390

12.5.2.1 Label-free detection based on

intrinsic DNA signals (direct

detection) 390

12.5.2.2 Compound-based detection

(indirect redox indicator-based

detection) 392


xiv Contents

12.6 Calculations About Compound-DNA Interactions 394


13 Electrochemical Nucleic Acid Biosensors Based onHybridization Detection for Clinical Analysis 403P. Kara, D. Ariksoysal, and M. Ozsoz13.1 Introduction 403

13.2 Biosensors 404

13.2.1 Nucleic Acid Hybridization Biosensors 405

13.3 Electrochemical Nucleic Acid Biosensors 407

13.3.1 Label-Based Electrochemical Nucleic Acid

Biosensors 408

13.3.1.1 Electrochemical genosensing by

using hybridization indicator 408

13.3.1.2 Electrochemical genosensing

with labeled signaling probe or

labeled target DNA 414

13.3.2 Label-Free Electrochemical

Genosensing 415

13.4 Conclusion 420

14 Nanomaterial-Based Electrochemical DNA Detection 427Ronen Polsky, Jason C. Harper, and Susan M. Brozik14.1 Introduction 427

14.2 Nanoparticle-Based Electrochemical DNA

Detection 429

14.2.1 Nanoparticle Modification of Electrodes

and Their Use as Supports for DNA

Immobilization 429

14.2.2 Gold Nanoparticle Supports 430

14.2.3 Magnetic Particles 432

14.2.4 Layer-by-Layer Immobilization Techniques 434

14.2.5 Metal Nanoparticle Labels for DNA

Hybridization Detection 435

14.2.5.1 Direct detection of the

nanoparticle label 435

14.2.5.2 Non-stripping-based

nanoparticle electrochemical

DNA detection methods 440


Contents xv

14.3 Nanowires, Nanorods, and Nanofibers 443

14.3.1 Nanorods as Labels 444

14.3.2 Nanowires Interfaced with Electrodes as

an Immobilization Matrix 445

14.3.3 Nanowire Conductance Based DNA

Detection 448

14.3.4 Electrochemical Impedance Spectroscopy

at Nanowires for DNA Detection 451

14.3.5 Dendrimers 452

14.3.6 Apoferritin Nanovehicles 455

14.3.7 Silica Nanoparticles 456

14.3.8 Liposomes 458

14.4 DNA Detection Using Carbon Nanotubes 461

14.4.1 Functionalization of Carbon Nanotubes

with DNA 462

14.4.2 CNTs for Electrochemical DNA Sensing 464

14.4.3 Progress toward CNT-Based Sensors for

DNA Detection 470

14.5 Conclusion 472

15 Electrochemical Genosensor Assay for the Detectionof Bacteria on Screen-Printed Chips 481Chan Yean Yean, Lee Su Yin, and Manickam Ravichandran15.1 Introduction 482

15.2 Methods for the Detection and Identification of

Microorganism Utilizing Enzyme-Based

Genosensors on Screen-Printed Chips 482

15.2.1 Electrochemical Genosensors for the

Detection of Bacteria 482

15.2.2 Principles of Enzyme-Based PCR

Amplicons Target DNA Detection

Methods 486

15.2.2.1 Direct method 486

15.2.2.2 Indirect method 488

15.2.2.3 Rapid method 488

15.2.3 Screen-Printed Transducer Surface 490

15.2.3.1 Screen-printed gold chip

genosensors 490


xvi Contents

15.2.3.2 Screen-printed carbon-chip

genosensors 491

15.3 Advantages of the Enzyme-Based Electrochemical

Genosensors in Detecting Bacteria on

Screen-Printed Carbon Chips 492

15.4 Discussions 493


16 Introduction to Molecular Biology Related toElectrochemical DNA-Based Biosensors 499Yalcin Erzurumlu and Petek Ballar16.1 Introduction 499

16.2 Nucleic Acids 501

16.3 Deoxyribonucleic Acid 506

16.4 DNA in Electrochemical DNA-Based Biosensors 509

16.5 Nucleic Acid Variants Used in Electrochemical

DNA-Based Biosensors 511

16.5.1 Peptide Nucleic Acid (PNA) 511

16.5.2 Locked Nucleic Acid (LNA) 513

Index 517


Preface

The discovery of DNA, the carrier of genetic information in cells,

brought with it many important technological accomplishments

such as the development of various diagnostic tools to unravel the

nature of hereditary diseases, gene expression profiling methods,

and genotyping. Among these, DNA biosensors constitute an

important class of point-of-care diagnostic devices because they

are capable of converting the Watson-Crick base pair recognition

event signal into an interpretable analytical signal in a shorter time

compared with other methods, thereby producing accurate and sen-

sitive results. Moreover, they are also suitable for miniaturization.

The terms “electrochemical DNA biosensor” and “nucleic acid–based

electrochemical biosensor” are used interchangeably.

By definition, biosensors are devices that fall into the subgroup of

biomedical sensors, combine a biological component with a detector

component, and are composed of three parts: (1) the biorecognition

element, such as an antibody, an enzyme, nucleic acids, or cell

lysates, which serves as a mediator; (2) the detector/transducer

element, which converts a biological signal into a readable output;

and (3) the signal processor, which displays a user-friendly version

of the transformed signal. Biosensors are classified according to

either the detector they are equipped with or the biorecognition

element they include. In general, the term “nucleic acid biosensors”

connotes devices that use single-stranded DNA as a biological

element. However, because of the advances in biosensor design, new

nucleic acid/nucleic acid analog interactions have been described

that are also considered to fall in this category, such as aptamer–

nucleic acid, RNA–DNA, peptide nucleic acid (PNA)–DNA, and locked

nucleic acid (LNA)–DNA. For the transduction of biological signals,

various kinds of detectors are available, but they can be categorized


xviii Preface

into three main classes: optical, electrochemical, and piezoelectric.

Because electrochemical DNA biosensors are miniaturizable (i.e.,

reducible in size to nanoscale dimensions), fast, accurate, simple,

and low cost, they have played perhaps the greatest role in the fields

of molecular and medical diagnosis, environmental monitoring,

bioterrorism, food analysis, pharmacogenomics, and drug discovery.

The aim of this book is to cover the full scope of electrochemical

nucleic acid biosensors by emphazing on DNA detection. The

material is presented in 16 chapters. Starting with the terminology

related to electrochemical DNA–based biosensors in Chapter 1,

the researchers active in the fields of biosensor design, molecular

biology, and genetics describe types of detection used for analysis

(chapters 6, 9, 11, and 13), types of materials used for biosensor

design (chapters 3, 4, 5, 8, 10, and 14), and types of nucleic acid

interactions detected (chapters 2, 7, 12, and 15).

I hope that this state-of-the-art book will continue to inform and

inspire all levels of scientists for many years. I wish to express my

gratitude to the researchers throughout the world who contributed

to the book by sharing their valuable studies in the field of

biosensors. In their honor, I quote the amazing scientist Albert

Einstein: “Imagination is more important than knowledge.”

I would also like to thank my wife, Ayse, for her love and patience

as well as the editorial group of Pan Stanford Publishing for their

assistance and support.

Mehmet OzsozIzmir, Turkey

March 19, 2012 18:56 PSP Book - 9in x 6in 01-Ozsoz-c01

Chapter 1

Terminology Related to ElectrochemicalDNA-Based Biosensors

Jan LabudaInstitute of Analytical Chemistry, Slovak University of Technology in Bratislava,81237 Bratislava, [email protected]

1.1 Introduction

With respect to low costs and high detection/information effec-

tiveness, physical and chemical sensors help us today widely to

check and control more and more processes everywhere around

us. Biosensors were introduced to chemical sensors about 50 years

ago with the aim of utilizing the recognition ability of biological

components such as enzymes, antibodies, etc., for the detection

of species of interest. Among them, biosensors with electrical and

electrochemical transducers are most popular in development and

application due to general advantages of electroanalytical methods

such as rather simple sensor fabrication, low costs of equipment and

analysis, possibility of miniaturization, and automation in chemical

analysis. Techniques and terms of electroanalytical chemistry have

been reviewed in technical reports of the Union for Pure and Applied

Electrochemical DNA BiosensorsEdited by Mehmet OzsozCopyright c© 2012 Pan Stanford Publishing Pte. Ltd.ISBN 978-981-4241-77-9 (Hardcover), 978-981-4303-98-9 (eBook)www.panstanford.com


2 Terminology Related to Electrochemical DNA-Based Biosensors

Chemistry (IUPAC) titled “Classification and Nomenclature of Elec-

troanalytical” Techniques” [1], “Recommended Terms, Symbols, and

Definitions for Electroanalytical Chemistry” [2], and “Recommended

Terms, Symbols, and Definitions for Electroanalytical Chemistry

(Recommendations 1985)” [3] and in Compendium of AnalyticalNomenclature: The Orange Book [4]. Some special articles charac-

terize electrochemical sensors [5]. A special IUPAC technical report,

“Electrochemical Biosensors: Recommended Definitions and Clas-

sification” [6], deals with techniques and terms of electrochemical

biosensors.

Since the 1990s [7] deoxyribonucleic acid (DNA) has been, and

today a rather large scale of nucleic acids (NA) is being, utilized

as the biorecognition element at a new group of biosensors–so-

called DNA or generally nucleic acid biosensors (more exactly DNA-

based biosensors). Very recently, a new technical report of the IUPAC

under the title “Electrochemical Nucleic Acid-Based Biosensors:

Concepts, Terms and Methodology” has been prepared [8]. It

represents a critical classification of terms and techniques used in

this dynamically developing field. With respect to construction and

utilization of DNA-based biosensors, specific terminology is used

(often not uniformly) in literature. The aim of this chapter is to

present the terminology of electrochemical DNA-based biosensors

and frequently used terms in a glossary format.

The electrochemical DNA-based biosensor can be characterized

as a device that integrates DNA (generally a nucleic acid) as a

biological recognition element and an electrode as a physicochem-

ical transducer. It is often presented as an electrode chemically

modified by nucleic acid. The pioneering concept of an electrode

modified with the DNA layer has allowed a significant decrease in

the amount of DNA tested/determined [9]. Following the definition

of a chemically modified electrode [10, 11], this is true for thin

(<100 μm) DNA layer coverage. Depending on the way of biosensor

fabrication, thicker films of DNA occur on the electrode’s surface,

which is sometimes even not considered and reported.

The choice of electrode material is connected, on one hand,

with the electrochemical process of interest. DNA immobilization

at the electrode surface is an initial step that plays a major

role in the overall biosensor performance. Methods used vary


Detection Features of DNA-Based Biosensors 3

depending on the kind of transducer and biosensor application,

and detailed experimental conditions have to be optimized for each

special application. The role of transducers (working electrodes)

is fulfilled by bulk electrodes – typically mercury-based (mercury

film, mercury amalgam), carbon-based (glassy carbon, carbon paste,

graphite, graphite-epoxy composite), and some other (gold, indium

tin-oxide) electrodes, or by various thin- and thick-film electrodes

(e.g., screen-printed carbon and gold electrodes). DNA array sensors

utilize transducers realized with interdigitated electrode [12]. There

are also a variety of techniques used for DNA immobilization [13–

15]. Surface and also “bulk” phase of the electrodes have been

modified by DNA [16]. Measurements with electrochemical DNA

biosensors are mostly performed in voltammetric and chronopo-

tentiometric detection modes [17]. With general improvement in

impedimetric biosensors, electrochemical impedance spectroscopy

(EIS) has become popular as the measurement technique for DNA-

based biosensors [18, 19].

Electrochemical DNA-based biosensors and electrochemical

sensing (assay) without use of the true biosensor are sometimes

confused in the literature [8]. While in the electrochemical DNA

biosensor the DNA layer has to be in an intimate contact with the

electrode prior to and during the NA interaction with an analyte,

in electrochemical sensing the DNA itself or product of any DNA

interaction, which was performed in solution or even at another

solid surface (magnetic beads, etc.), is detected electrochemically,

usually after preconcentration by an accumulation on the electrode

surface.

1.2 Detection Features of DNA-Based Biosensors

DNA-based biosensors possess specificity of response, which is

typical for biosensors taking advantage of the bioaffinity properties

of DNA. Compared with enzyme sensors and immunosensors,

DNA biosensors are mostly used for the investigation of DNA

interactions rather than for conventional determination of the

concentration of an analyte. They exhibit typical biosensor selec-

tivity/specificity to the analyte (e.g., nucleotide bases sequence,



protein) or class selectivity to DNA as the recognition element itself

(e.g., damage to DNA) [8]. With respect to this characteristic, DNA-

based biosensors represent irreplaceable testing (bio)analytical

devices.

Working procedures with these biosensors utilize special detec-

tion principles. In the first place, label-free techniques utilizing

electrochemical and/or surface activity of DNA have to be men-

tioned [17]. Electrochemical activity of DNA is based on the

presence of redox changes in nucleobases and sugar residues.

All common nucleobases are known to undergo electrochemical

oxidation at carbon electrodes. At neutral and weakly acidic pH,

adenine, cytosine, and guanine residues in DNA produce reduction

signals at mercury-based electrodes at highly negative potentials,

while guanine residues yield anodic signals due to oxidation of

their reduction product back to guanine. Protonation of base

residues is involved in the electrode process. Mercury electrodes are

particularly sensitive to minor conformational changes in DNA such

as those induced by nucleases and chemical and physical agents,

including ionizing radiation [17].

Nucleic acids are usually strongly adsorbed on electrodes,

particularly on mercury and carbon ones. For mercury electrodes,

the adsorption/desorption behavior of DNA strongly depends on

the structure of the DNA molecules. DNA electrochemical surface

activity depends on what DNA components take part in adsorption

at the electrode surface. The height of the tensammetric peak

increases with the chain length. Adsorption of DNA on mercury

electrodes proceeds only in one layer, and the formation of further

layers does not influence the intensity of electrochemical signals.

The polyanionic nature of nucleic acids leads to characteristic

adsorption/desorption (reorientation) processes at mercury-based

electrodes upon application of negative electrode potentials due

to interplay between electrostatic repulsion and relatively strong

adsorption via hydrophobic parts of the polynucleotide chains

(particularly bases) [13, 17]. Electrochemical analysis of the DNA

can, thus, in principle, be performed without introducing labels

into the DNA recognition element (label-free techniques) and even

without introducing any additional reagent into the measuring

system (reagent-less techniques).


Detection Features of DNA-Based Biosensors 5

As only guanine moieties in the close vicinity of the electrode

surface can undergo direct electrooxidation, soluble redox media-

tors such as rhodium or ruthenium complexes are sometimes used

to shuttle electrons from guanine residues in distant parts of DNA

chains to the electrode [20]. In such a case, we cannot speak more

about the reagent-less technique. Nevertheless, the electrochemical

reduction and oxidation of nucleobases are irreversible and thus do

not allow reusability of biosensors.

An alternative approach to the intrinsic DNA electrochemical

activity utilizes electroactive species as redox indicators of the

presence of immobilized DNA as well as its interaction events such

as hybridization, damage, and association with another substance

[14]. This mode was also used in a pioneering work on the DNA

biosensor used for sequence detection [7]. In this case, it is still a

label-free method in the sense that DNA probes or targets are not

chemically modified by a special label; however, as the indicator has

to be added to a test system as an additional reagent, we cannot

speak more about the reagent-less technique. Redox indicators

typically possess electrochemical responses at a “safe” electrode

potential and often reversibly. The terms redox probe and redoxmarker are sometimes used in the literature to mean the redox

indicator, which is confusable with the DNA capture probe used as

a recognition element at hybridization and with markers used in

medical diagnostics [8].

DNA redox indicators bind to DNA or are present in the

solution phase. Some of them interact with DNA on the basis of

electrostatic forces [21]. Cationic indicators such as metal complex

cations can be attracted to the DNA by the negative charge of

the DNA backbone. On the other hand, anionic indicators, for

instance hexacyanoferrate (III/II) [Fe(CN)6]3–/4–, work on the

principle of repulsion by the negatively charged DNA backbone. As

a consequence, its voltammetric current response is lower than and

anodic to the cathodic peak potential separation, higher than that

observed at bare electrodes without DNA. Electrostatic indicators

can also respond to differences in negative charge density between

ssDNA and dsDNA [14].

Other DNA redox indicators intercalate into the dsDNA structure

(e.g., daunomycin, phenoxazines, metal complexes with condensed



aromatic heterocyclic ligands) or bind to dsDNA grooves (e.g.,

Hoechst 33258). All together, cationic indicators, intercalators, and

groove binders accumulate at the immobilized dsDNA layer (e.g.,

after hybridization or prior to damage to the DNA duplex), thus

increasing their measured voltammetric response. The biosensor

can be used repeatedly after its renewal using the sequence

of steps: indicator accumulation, voltammetric measurement and

chemical removal or desorption of the accumulated indicator

from the DNA layer. Then, for one and the same biosensor a

mean indicator response and its standard deviation are calculated

[21].

Indicators associating preferentially with ssDNA have been

advantageously used with electrochemical DNA hybridization sen-

sors. For instance, the phenothiazine dye methylene blue (MB) asso-

ciates with unpaired guanine moieties. In dsDNA this interaction

is hampered, which results in decrease in the current response

due to MB reduction [22]. On the other hand, there are also ds-

specific electroactive indicators such as the intercalator ferrocenyl

naphthalene diimide, which results in a detection limit of 10 zmol at

the differential pulse voltammetric mode [23, 24].

Finally, electrochemically active DNA labels (tracers), which are

covalently bound to DNA, can be used for detection. The DNA labels

considerably improve analytical selectivity/specificity, for instance,

at DNA hybridization as the labeled DNA can be distinguished

from the unlabeled one [17, 25]. Among such labels, ferrocene,

daunomycin, anthraquinone, thionine, bipyridine complexes of Ru

and Os, nitrophenyl, and aminophenyl groups have to be mentioned.

Osmium tetroxide complexes with nitrogen ligands (OsVIII,L) [26, 27]

or analogous osmate complexes (OsVI,L) [28] represent examples

of electroactive tags. Nanoparticles or nanocrystals of gold, indium,

zinc, cadmium, or lead chalcogenides and other materials have

been used as labels covalently (often via thiol linkage) attached

to DNA probes applied in amplifying the response. By combining

various nanoparticles such as ZnS, CdS, and PbS, electrochemical

“multicolor” DNA coding has been attained [29]. Carbon nanotubes

as DNA tags can also be loaded with multiple nanoparticles or

enzyme molecules, thus offering considerable signal enhancement

[22, 29, 30].


Detection of Specific DNA Interactions 7

1.3 Detection of Specific DNA Interactions

Among specific DNA interactions tested using DNA-based biosen-

sors, DNA hybridization, DNA association with low molecular mass

compounds (drugs, chemicals), and DNA damage are typically

considered.

1.3.1 DNA Hybridization Biosensors

DNA hybridization is a chemical interaction of DNA based on the

ability of ssDNA to form a helix, dsDNA with ssDNA counterpart

exhibiting nucleotide sequence complementarity. In DNA hybridiza-

tion biosensors, a specifically designed ssDNA probe (capture

probe [CP]) with a defined (known) nucleotide sequence is usually

immobilized on the electrode surface and allowed to interact as

a recognition element with target DNA (tDNA) in test solution.

By varying experimental conditions such as the pH, temperature,

and ionic strength, hybridization efficiency can be controlled, thus

allowing detection of single- or multi-base mismatches [15, 31].

Experimental arrangement for electrochemical DNA hybridiza-

tion biosensors includes the following:

1. Label-free and indicator (reagent)-less detection of target DNA

typically based on guanine residues response.

2. Noncovalent redox indicators that allow distinguishing between

the ssCP and dsDNA hybrid at the electrode surface (successful

hybridization) [22, 23].

3. Sandwich hybridization assay that employs a covalently labeled

reporter or signaling probe (RP) and involves two tDNA recogni-

tion steps (CP-tDNA and tDNA-RP) [32]. The RPs are designed to

hybridize with the tDNA at a site next to the sequence recognized

by the capture probe to confer efficient electronic communication

between the label and the electrode.

4. Peptide nucleic acid (PNA) probes as a DNA analogue that possess

an uncharged pseudopeptide backbone instead of the charged

phosphate-sugar backbone of natural DNA and, consequently,

greater affinity to complementary DNA and better distinction

between closely related sequences [33].



Electrochemical biosensors of single nucleotide polymorphisms

(SNP, point mutations) can use

(i) different stabilities of duplexes displaying full complemen-

tarity between the probe and tDNA (homoduplexes between

wild-type probe and wild-type tDNA or mutant probe and

mutant target) and those involving mismatched nucleotides

(heteroduplexes between wild-type probe and mutant target,

or vice versa) [13]. Discrimination of perfectly matched and

mismatched duplexes can be achieved by performing DNA

hybridization at stringent conditions achieved by elevated

temperature and decreased ionic strength or via applying

a peptide nucleic acid probe instead of DNA. Under opti-

mum conditions, the homoduplex gives positive hybridization

response, while the heteroduplex is not stable, thus giving a

signal-off response to the mutation in one of the hybridizing

strands.

(ii) primer extension incorporation of a labeled nucleotide within

the SNP site [30]. The target template is annealed with a primer

complementary to the target segment “upstream” (relative to

DNA polymerase catalyzed elongation of the primer that always

proceeds in the 5’→3’ direction) to the position of interest, and

a labeled dNTP (e.g., with biotin to attach an enzyme in the

following step, or with a redox marker) is added to the reaction

mixture. Under proper conditions, the labeled nucleotide is

attached to the primer only when it is complementary to

the base at the first “free” position. Using different labels for

different nucleotides, all four possible bases within the SNP site

can be probed in a single reaction.

(iii) electronic properties of the duplex DNA and perturbations

in the DNA electronic properties in the presence of single

base mismatches [34]. Disruption of π -stacks within the DNA

double helix due to presence of the mismatch has been shown

to prevent DNA-mediated charge transfer between electrode

and an intercalator bound at the opposite (relative to the

electrode surface) end of the double helix, which was efficient

in the perfectly matched (and perfectly base-pair-stacked)

homoduplex.



(iv) Electrochemical determination of the length of guanine-

containing triplet repeats was achieved by the mediator-based

guanine electrocatalytic oxidation technique. Other approaches

applied for this purpose involve multiple hybridization of a

labeled RP spanning several triplet units with the expanded

triplet repeat [25]. The number of RP molecules hybridized

(or labels collected) per tDNA strand is proportional to the

length of the repetitive sequence, which is – after proper

normalization to the number of target strands – reflected by

intensity of the measured signal.

1.3.2 DNA Damage

As DNA belongs to main body substrates that undergo serious

structural changes such as oxidation of the DNA bases and sugar

moieties and/or release of the bases as well as DNA strand breaks

caused by chemical systems generating so-called reactive oxygen

(ROS), nitrogen (RNS), or sulfur (RSS) species [35, 36] and by other

classes of genotoxic substances [37], the second main application

area of DNA-based biosensors is detection of damage to DNA. ROS

are produced endogenously, during normal aerobic metabolism

and under various pathological conditions, and exogenously, such

as upon exposure to UV light, ionizing radiation, environmental

mutagens, and carcinogens. About 104 to 106 DNA damage events

occur to a cell per day [37]. Accumulation of oxidative DNA lesions

is associated with aging and with a variety of human diseases,

including cancer and neurodegeneration. The terms DNA damage(see below) and mutation should not be intermingled. While

mutation refers to a change in DNA sequence, in damaged DNA

the chemical nature of individual nucleotides is changed, which can

result in mutation.

Altered chemical, physicochemical, and structural properties of

damaged DNA are reflected in its redox behavior, which is utilized

in numerous techniques of DNA damage detection. Electrochemical

DNA-based biosensors have been used not only to detect but also

to induce and control DNA damage at the electrode surface via

electrochemical generation of the damaging (usually radical) species

[13]. This way, chemicals and drugs such as niclosamide, adriamycin,



benznidazole, thiophene-S-oxide, and nitroderivatives of polycyclic

aromatic compounds have been investigated [38–40].

Experimental arrangement for electrochemical DNA hybridiza-

tion biosensors includes the following:

(a) Label-free detection of strand breaks with mercury-based DNA

biosensors. These biosensors are based on strong dependence

of accessibility of DNA bases to the transducer surface (which

is lower at intact DNA compared with damaged DNA) and DNA

conformation (and/or local perturbations). Hence, mercury-

based DNA biosensors are able to discriminate between

DNA molecules containing (e.g., ssDNA) and lacking (e.g.,

sc plasmid DNA) free chain ends when free ends produce

specific electrochemical responses under certain conditions.

Nicking of supercoiled (sc) plasmid DNA with enzymes (such

as DNase I) as well as reactive radical species that destroy the

deoxyribose moieties, some types of nucleobase lesions after

their conversion to strand breaks by specific enzymes, and

repair of the strand breaks by action of the DNA ligases were

detected as well [13].

Detection of the sb at the hanging mercury drop electrode

(HMDE) is highly sensitive. By using alternating current (AC)

voltammetry, one sb was detected among more than 2 × 105

nucleotides [41]. Although conventional HMDE possesses such

unique features, successful attempts have been made to replace

it by other electrodes in which the liquid mercury content would

be minimized or eliminated. Both redox and tensammetric DNA

signals have been measured at a mercury-film-coated solid

glassy carbon electrode (MF/GCE) and at different variants of

silver solid amalgam electrodes (AgSAE). MF/GCE [42], as well

as AgSAE and MF-AgSCE [43] modified with scDNA, was applied

to sb formation.

(b) Detection of DNA degradation at carbon-based biosensors

using redox indicators. Deep degradation of DNA during the

step of biosensor incubation for a given time (minutes to

hours) in a cleavage medium under investigation, after the

medium exchange for the follow-up electrochemical measure-

ment, results into diminution of the voltammetric response



of the metal complex indicator that binds to DNA (such

as [Co(phen)3]3+ [14, 44–46]) or into enhancement of the

voltammetric response of the negatively charged metal complex

like [Fe(CN)6]3–/4–, which is repulsed by the negatively charged

DNA layer depending on the degree of DNA damage [47, 48].

Change in the indicator electrochemical response depends on

the portion of DNA damaged in the cleavage reaction. Similarly,

a decrease in the charge transfer resistance at an impedimetric

biosensor with hexacyanoferrate as the redox indicator in

solution was used [47, 48].

These types of DNA detection can also be applied to studies of

antioxidative properties of various natural substances preserv-

ing DNA from damage [49, 50]. The detection scheme exploits

quantification of the DNA portion that survives previous incu-

bation of the biosensor in a mixture of the DNA cleavage agent

and antioxidant/mixture of antioxidants under investigation.

Using this approach, yeast polysaccharides, phenolic acids such

as rosmarinic and caffeic acids, selected flavonoids, as well as

aqueous plant extracts and tea extracts were studied [51].

(c) Guanine residues’ redox responses [13, 14]. Among DNA base

residues, those of guanine not only possess electrochemical

response but are also the most frequent target for a range

of genotoxic agents. Consequently, the guanine residues’ redox

responses represent the most frequently used approach for

DNA damage detection. Decrease in the guanine peak current

relative to that yielded by undamaged DNA represents the

response to damage to the nucleobase and/or its release from

the polynucleotide chains, which is an event often following

modifications within the guanine imidazole ring. Since natural

DNA contains many guanine residues, partial decrease in the

guanine peaks is usually observed, depending on the extent of

DNA damage.

In contrast to analysis with HMDE, MFE, or AgSAE, measure-

ments of the guanine oxidation signal at carbon electrodes

(GCE, CPE, SPCE) cannot provide information about formation of

individual sb due to a lack of differences in the signal intensity

of sc and dsDNA (both oc and lin DNA that possess free ends)

but can be used for monitoring deep DNA degradation, involving



damage to the guanine base and/or disintegration of DNA

molecules into small fragments [52]. With various biosensor

arrangements, effects of agents such as antitumor platinum

complexes [53] and various aromatic hydrocarbons derivatives

[54] on arsenic oxide [55] have been investigated using the

guanine response at the mercury-based electrodes. Besides

low specificity of this type of response, the general problem

of relatively low sensitivity is connected with the signal-off

approach.

(d) Detection of electroactive products of DNA damage. Some

products of DNA damage exhibit characteristic electrochemical

activity possessing a new signal. For example, 8-oxoguanine

(8-OG) is electrochemically oxidized at carbon electrodes at

a potential significantly less positive than the parent guanine

base [14, 38–40]. Compared with the previous one (described

under c), this approach exhibits much better sensitivity and

specificity. New species can be detected also using a redox

mediator. The complexes of osmium (such as [Os(bipy)3]3+) and

ruthenium with different redox potentials have been shown as

electrocatalysts for 8-OG and guanine, respectively [17, 56].

(e) Layered assemblies for genotoxicity screening. Multilayer

assemblies of cationic redox-active polymer films, DNA, and

heme proteins at carbon electrodes were designed for testing

the genotoxic activity of various chemicals [57]. In these devices,

layers of enzymatically active hemoproteins mimic metabolic

carcinogen activation processes (e.g., styrene is enzymatically

converted to styrene oxide). The activated species diffuse into

the DNA layer and attack guanine residues, and the damaged

DNA double helix is indicated by using guanine oxidation

mediated by a cationic polymeric film.

(f) A molecular beacon-like sensor for the evaluation of nuclease

and ligase activities. An electrochemical biosensor using a

hairpin DNA with an oxidizable ferrocene label was published

for the detection of activities of enzymes such as nucleases

(generating single-strand breaks) and DNA ligases (sealing the

break) [58]. At a single-strand break in the duplex part of the

hairpin structure, the ferrocene-labeled segment was removed

under conditions of danaturation with diminution of the current



signal. In the presence of ligase activity, the break was joined,

preventing removal of the ferrocene-labeled segment.

1.3.3 DNA Association Interactions

1.3.3.1 Binding of low molecular mass compounds

DNA association interactions are of interest for chemistry, molecular

biology, and medicine, particularly for drug discovery and envi-

ronmental/medical processes [59, 60]. They concern association

with both inorganic and organic compounds as well as various

types of assisted interactions such as metal and metal complex–DNA

chemistry [61]. DNA-based biosensors serve as effective screening

tools for in vitro tests of this large group of DNA interactions.

Due to the preconcentration effect within the DNA structure, the

detection/concentration determination of a trace low molecular

mass analyte or group of analytes could also be a result of the study.

These noncovalent host–guest interactions are represented

mainly by [14]

(a) intercalation between the stacked base pairs of dsDNA,

(b) binding at major or minor grooves of the DNA double helix, and

(c) electrostatic interactions.

The intercalation as an insertion of guest molecules between

the stacked base pairs of the double helix structure leads to a

change in the dsDNA chain, which must lengthen and unwind

slightly. The intercalation can also have an influence on the

electrochemical activity of the intercalator. For instance, doxorubicin

and complexes of transient metals with 1,10-phenanthroline or

ferrocene naphthalene diimide retain their redox response after the

intercalation, but some others, e.g., phenothiazines, do not show

significant current signals after the intercalation. Sometimes the

intercalation can result in secondary interactions that can be used

for the detection, e.g., electron transfer from the guanine residues

(using, say, the [Ru(bpy)2]2+ complex), or generation of ROS able

to initiate oxidative cleavage of ribose cycles in the primary DNA

sequence.



In contrast to intercalation, electrostatic interactions are formed

between positively charged guest molecules and the negatively

charged DNA sugar-phosphate backbone. However, depending on

the experimental conditions, these interaction modes can also be

combined [21]. For instance, the dsDNA interaction with positively

charged metal complex compounds with aromatic ligands is

predominantly electrostatic at low ionic strength and predominantly

intercalative at high ionic strength. The character of the binding

interaction of the components of electrically charged redox couples

(e.g., metal complexes) can be estimated from a net negative or

positive formal potential shift when the first one indicates the

stabilization of the component in a higher oxidation state over that

in a lower oxidation state, i.e., the electrostatic interaction, and the

second one can be ascribed to the intercalation [21, 62].

There are also compounds, particularly from the drug family (e.g.,

mitomycin C), that form covalent bonds with DNA bases and create

adducts yielding specific electrochemical responses [13].

The voltammetric response of association interaction relates

to an electrochemically active analyte, to an electrochemically

active species competing with analyte binding, or to guanine and

8-oxoguanine. Using an impedimetric DNA biosensor, distortion

of the surface-attached DNA can also be specified by appropriate

changes in the resistance of the charge transfer and capacity of the

surface layer. Impedimetric measurements provide also the possi-

bility of detecting electrochemically inactive analytes, which do not

bring about remarkable changes in the guanine oxidation current

[18, 19]. Recently, impedimetry performed in the presence of inter-

calators has been reported to specify the type of DNA interaction

[63].

1.3.3.2 Binding of proteins

Using DNA biosensors, two types of DNA–protein interactions can be

investigated: first, detection of catalytic activity of DNA-processing

enzymes such as nucleases, ligases, and polymerases; and second,

affinity interactions of DNA with proteins that can but need not be

enzymes. The detection techniques used can be the same as those

mentioned above for DNA hybridization sensors. Electroactivity of


Conclusions 15

amino acid residues in proteins allows for direct electrochemical

measurement without any labeling [64].

In specific cases, disturbance of the base pair stacking via flipping

out a nucleobase or via bending the duplex was found to affect the

dsDNA-mediated charge transfer at a gold electrode [65].

1.4 Conclusions

In this chapter, DNA-based biosensors were presented as special

analytical devices capable of selective or class-selective detec-

tion/recognition of chemical interactions of the surface-confined

DNA with substances of interest such as oligonucleotides, low

molecular mass compounds, and species leading to DNA damage

and preservation of DNA structure, together with related, rather

special terminology. As was stated for electrochemical biosensors

generally [6], definitions, terminology and classification cannot

unambiguously address every detail, nuance and contingency of this

diverse subject. This is also fully true for the rapidly developing

field of DNA-based biosensors with new forms of nucleic acids used;

new ways of sensor fabrication, measurement arrangement and

procedures; and finally new practical utilization. Nevertheless, the

terminology and classification presented here rather systematically

and documented by numerous examples could help build up

communication and understanding between experts and students in

this field.

We believe that it will also stimulate progress in the systematic

development of DNA biosensors and their application as screening

tools for drug investigation, as warning systems in rapid chemical

toxicity tests, as testing devices in food and water analysis, in

the evaluation of effects of antioxidants, and in the investigation

of interactions of nucleic acids with other biomacromolecules as

proteins.

Glossary

AC voltammetry/polarography An analysis of the current response

to a small-amplitude sinusoidal voltage perturbation superimposed



on a DC (ramp or constant) potential [66]. A plot of the AC current

vs. sweep potential produces a derivative-type polarographic curve

[4].

Antioxidants Substances that at low concentrations than those

of an oxidizable biochemical substrate markedly delay or prevent

oxidation of this substrate [67]. Their behavior could be ascribed

to scavenging reactive radicals and chelation of redox-active metals,

particularly iron and copper. The most active and evaluated dietary

antioxidants belong to the family of phenolic and polyphenolic

compounds.

Antioxidative activity Complex parameter based on the (bio)

chemical reactivity of antioxidants. The antioxidative activity

belongs to characteristics typically defined operationally regarding

the procedure used. This applies to the utilization of DNA-based

biosensor as well.

Array electrodes Replacement of a single electrode (with dimen-

sions in the micrometer or centimeter range) by an array of

(ultra)microelectrodes [66].

Biological recognition system/biological receptor An element that

translates information from the biochemical domain, usually an

analyte concentration, into a chemical or physical output signal with

a defined sensitivity. The main purpose of the recognition system is

to provide the sensor with a high degree of selectivity for the analyte

to be detected [6].

Bases of nucleic acids Nitrogenous bases (purines such as adenine

and guanine or pyrimidines such as cytosine, thymine, and uracil).

Adenine, guanine, and cytosine are found in both deoxynucleotides

and ribonucleotides, whereas uracil is found primarily in ribonu-

cleotides, and thymine in deoxynucleotides.

Biosensor An integrated device incorporating a biological/

biomimetic recognition system either integrated within or inti-

mately associated with a physicochemical transducer [68]. Biosen-

sors are chemical sensors in which the recognition system utilizes a

biochemical mechanism [6].


Glossary 17

Capture probe (CP) A specifically designed ssDNA with a defined

(known) nucleotide sequence usually immobilized on a transducer

or other surface. The CP is utilized as a recognition element to test

nucleotide sequence of target DNA (tDNA) in the sample solution by

using hybridization.

Chemical sensor A device that converts chemical information

such as the presence/concentration of specific sample components

into a measurablel signal [66]. Chemical sensors contain two basic

functional units connected in a series: a chemical (molecular)

recognition system (receptor) and a physicochemical transducer

[6]. It is capable of continuously recognizing the presence and/or

concentration of a chemical constituent in a liquid or gas and

converting this information in real time to an electrical or optical

signal.

Chemically modified electrode An electrode made of a conduct-

ing or semiconducting material that is coated with a selected

monomolecular, multimolecular, ionic, or polymeric film of a

chemical modifier and that by means of faradaic (charge transfer)

reactions or interfacial potential differences (no net charge transfer)

exhibits chemical, electrochemical, and/or optical properties of

the film [10, 11]. The chemically altered bare (working) electrode

exhibits new qualities concerning selectivity and sensitivity as well

as against fouling and interferences.

Circular DNA A structure of DNA when its double-helical segment

is closed to a circle by joining its two ends.

DNA (deoxyribonucleic acid) A polyanionic biopolymer consisting

of a chain of nucleotides linked with phosphates bridge at the 3’ and

5’ positions of neighboring sugar (2-deoxyribose) units (ssDNA).

Complementary base pairing results in the specific association of

two polynucleotide chains that wind around a common helical axis

to form a double helix (dsDNA).

DNA-based biosensor A biosensor that uses DNA as the biorecog-

nition element.



DNA biosensor In general, a biosensor used for detection of DNA

and/or its specific interactions. It is mostly represented by a DNA-

based biosensor.

DNA damage Alteration in the DNA chemical structure resulting

from interactions with physical or chemical agents occurring in

the environment, generated in the organisms as by-products of

metabolism or used as therapeutics [13]. The main types of DNA

damage include interruptions of the sugar-phosphate backbone

(strand breaks), release of bases due to hydrolysis of N-glycosidic

bonds (resulting in abasic sites), and a variety of nucleobase lesions

(adducts) resulting from reactions of DNA with a broad range of

oxidants, alkylating agents, and others.

DNA hybridization Chemical interaction of DNA based on the

ability of ssDNA to form a helix, dsDNA with a counterpart exhibiting

nucleotide sequence complementarity. A process of the formation of

dsDNA from ss polynucleotide chains based on complementary base

pairing.

DNA label (tracer) Species covalently bound to DNA and used in its

electrochemical detection.

Electrochemical biosensor A self-contained integrated device that

is capable of providing specific quantitative or semiquantitative

analytical information using a biological recognition element (bio-

chemical receptor), which is retained in direct spatial contact with

an electrochemical transduction element [6]. A biosensor with an

electrochemical transducer may represent a chemically modified

electrode.

Electrochemical DNA-based biosensor A biosensor that integrates

DNA (generally a nucleic acid) as the biological recognition element

and an electrode as the physicochemical transducer.

Electrochemical cell/voltammetric cell A cell where electrochemi-

cal/voltammetric measurements are performed. It incorporates an

ionic conductor (electrolyte, sample solution) and typically three

electrodes: a working electrode (a microelectrode), a current-


Glossary 19

conducting electrode (auxiliary or counterelectrode), and a refer-

ence electrode.

Electrochemical impedance spectroscopy A technique based on

evaluation of the interfacial impedance, which is obtained upon

application of a small AC voltage overlaid on a DC bias potential to

the sensing (working) electrode and measurement of the AC current

obtained in the steady state.

Electrode/working electrode In general, an electrode that serves

as a transducer responding to the excitation signal and the

concentration of the substance of interest in the solution being

investigated, and that permits the flow of current sufficiently large to

effect appreciable changes of bulk composition within the ordinary

duration of a measurement [2–4]. In electrochemical analysis, differ-

ent working electrodes are used, e.g., dropping mercury electrode

(DME) (typically in polarography), static mercury drop electrode

(SMDE), or solid electrodes (in voltammetry and other electroana-

lytical techniques). In electrochemical sensors/biosensors, suitable

working electrodes are used as physicochemical transducers that

convert a biological recognition event into a measurable signal.

Groove binding Binding of a guest molecule, typically of a moon-

shaped and flat in structure, into the exterior of the DNA helix.

Impedimetric DNA biosensor A DNA biosensor based on electro-

chemical impedance spectroscopy (EIS) detection. It is a device

that transduces changes in interfacial properties between the

electrode (with the DNA film) and the electrolyte induced by

DNA hybridization, conformational changes, or DNA damages to an

electrical signal [19].

Immobilization A method that can immobilize a biological receptor

with high biological activity in a thin layer at the transducer surface.

It is a step in biosensor fabrication.

Intercalation Insertion of a guest molecule between the base pairs

of the DNA helix.



Intercalator A compound that undergoes intercalation, typically a

molecule with a planar structure containing three or four aromatic

rings.

Label-free detection technique Procedure that utilizes electro-

chemical and/or surface activity of DNA (reduction and tensammet-

ric responses of DNA at mercury and some amalgam electrodes,

guanine oxidation at carbon electrodes, detection by using nonco-

valent DNA redox indicators, etc.). The label-free technique uses no

chemical modification of a DNA probe or target or another analyte

interacting with NA.

Microelectrode/ultramicroelectrode An electrode with a char-

acteristic dimension ranging from 25 μm to 1 mm [66]. An

ultramicroelectrode has a characteristic dimension less than 25 μm.

This characteristic dimension refers to the diameter of a disk, a

sphere, a hemisphere, and a cylinder, and the width of a band

ultramicroelectrode.

Nucleic acid aptamers Single-stranded oligonucleotides (mainly

DNA or RNA) originating from in vitro selection that, starting

from random sequence libraries, optimize the nucleic acids for

high-affinity binding to a given target [69, 70]. Aptamers, upon

association with their target, fold into complex three-dimensional

shapes in which the target becomes an intrinsic part of the nucleic

acid structure.

Nucleobase lesion A chemical modification of nucleobase, e.g., its

oxidative change.

Nucleotide A molecule composed of a nitrogenous base (purine or

pyrimidine) linked to a sugar (deoxyribose or ribose) to which at

least one phosphate group is attached.

Nucleoside A molecule composed of a nitrogenous base (purine or

pyrimidine) linked to a sugar (deoxyribose or ribose).

8-oxoguanine (8-OG) The oxidation product of guanine, which can

be electrochemically oxidized at carbon electrodes at a potential

significantly less positive than the parent guanine base.


Glossary 21

Physicochemical electrochemical transducers See Electrode/workingelectrode

Reagent-less detection technique A procedure that uses no addi-

tional chemical reagents (indicator, redox mediator, enzyme sub-

strate) to generate an analytical signal of the DNA biosensor.

Redox reaction A chemical reaction in which the reactants

exchange electrons between each other. As a consequence, the

oxidation states of the elements prior to and following the redox

reaction are altered [66].

Electrode reaction An interfacial reaction that necessarily involves

a charge transfer step [66] between a chemical reactant (depolar-

izer) and the electrode (an electrochemical reaction). The electrode

reaction involves all processes (chemical reaction, structural reorga-

nization, adsorption) accompanying the charge transfer step.

Redox mediator A chemical compound that can shuttle electrons

between two other chemical compounds in solution or between an

electrode and a chemical species in solution [66].

Screen-printed electrode An electrode prepared by forced screen

printing of a powder-based ink through a screen stencil typically

on a plastic sheet or foil, or ceramic plate, as a single or set of film

electrodes [66].

Selectivity of the DNA-based biosensor It can be truely considered

as an analytical parameter regarding the analyte detected such as

a specific ssDNA base sequence or protein interacting with nucleic

acid aptamer. Generally, class selectivity to DNA as the recognition

element itself can be considered (e.g., at damage to DNA).

Signal-on/signal-off measurement technique A procedure based

on appearance/diminution of analytical response resulting from

molecular interaction at the biosensor.

Single-/multi-base mismatch A defect in the double-stranded DNA

structure that distinguishes DNA hybrid containing a mismatched

base pair or pairs from that with fully matched bases.



Single nucleotide polymorphisms (SNPs, point mutations) A

variant of DNA sequence in which the purine or pyrimidine base (as

cytosine) of a single nucleotide is replaced by another such base (as

thymine). It is the most common type of change in DNA. SNPs occur

normally throughout a person’s DNA once in every 300 nucleotides

on average, which means there are roughly 10 million SNPs in the

human genome. They can act as biological markers.

Strand break An interruption of the sugar-phosphate backbone of

the nucleotide.

Supercoiled DNA A contortion of circular DNA into the shape

of the simple figure eight. DNA supercoiling is important for DNA

packaging within all cells.

Tensammetry Measurement of the interfacial capacitance as a

function of potential. It is used especially in the analysis of surface-

active substances that are not electroactive [66].

Transducer Part of the sensor/biosensor that converts a detected

physical or chemical change into a measurable (usually electronic)

signal. Working electrodes are used as transducers in electrochemi-

cal biosensors.

Voltammetry/polarography Measurement of current as a function

of a controlled electrode potential and time, which results in a

current–voltage (or current–time or current–voltage–time) display,

commonly referred to as the “voltammogram” [66]. The working

electrode is situated typically in the voltammetric cell and is a

dropping mercury electrode in the case of polarography.

List of abbreviations

AC alternating current

AgSAE silver solid amalgam electrode

CNTs carbon nanotubes

CP capture probe

CPE carbon paste electrode

DC direct current


References 23

DME dropping mercury electrode

DNase deoxyribonuclease

dNTP deoxynucleotide triphosphate

ds double stranded

dsDNA double-stranded DNA

EIS electrochemical impedance spectroscopy

GCE glassy carbon electrode

HMDE hanging mercury drop electrode

L ligand

lin linear

MB methylene blue

MFE mercury film electrode

oc open circular

8-OG 8-oxoguanine

ODN oligodeoxyribonucleotide

PCR polymerase chain reaction

PNA peptide nucleic acid

RNS reactive nitrogen species

ROS reactive oxygen species

RSS reactive sulfur species

RP reporter probe

SMDE static mercury drop electrode

SNP single-nucleotide polymorphisms

SPE screen-printed electrode

SPCE screen printed carbon electrode

sb strand break

sc supercoiled

ss single stranded

ssDNA single-stranded DNA

ssb single-strand break

tDNA target DNA

UV ultraviolet

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Chapter 2

Electrochemical Aptamer-BasedBiosensors

S. Centi, S. Tombelli, and M. MasciniDipartimento di Chimica, Universita degli Studi di Firenze,Via della Lastruccia 3, 50019 Sesto Fiorentino, andIstituto Nazionale Biostrutture e Biosistemi (INBB),Viale Medaglie d’Oro 305, 00136 Roma, [email protected]

2.1 Introduction

Aptamers are oligonucleotides (DNA or RNA molecules) which are

able to bind selectively to low-molecular-weight molecules or to

macromolecules such as proteins. The interest in aptamers as lig-

ands is related to their ease of preparation by an evolutionary selec-

tion procedure that eliminates the need for structural design of the

ligand sites. Selection of the aptamers for the specific target is based

on the SELEX (systematic evolution of ligands by exponential enrich-

ment) procedure.

In recent years, great progress has been made toward the devel-

opment of aptamer-based assays. These assays can be set up in a

wide variety of formats (direct, sandwich, or competitive assays),

which are summarized in Fig. 2.1.



30 Electrochemical Aptamer-Based Biosensors

Figure 2.1. Schematic representation of aptamer-based assays configu-

ration. (A) Direct assay (B) Sandwich assay format with two aptamers (C)

Sandwich assay format with an aptamer used as the primary ligand and an

antibody as the secondary ligand (D) The opposite configuration to case C

(E) Competitive assay.

In a direct assay (Fig. 2.1A) the aptamer is immobilized on the

solid support and the binding of analyte is monitored; a sandwich

assay can be carried out using two aptamers as ligands or combin-

ing an aptamer with an antibody. In this assay format, a capturing

aptamer or antibody is first immobilized on the solid support and

then analyte is added so that the capturing ligand could bind it.

At this point, a detection aptamer or antibody is added and binds

with another site of the target analyte (Fig. 2.1B–D). In addition to

detecting macromolecules, such as proteins, small ligands can also

be bound by aptamers. For this purpose, a competitive assay can

be performed by immobilizing the analyte on the solid support and

then adding to it a solution containing the target analyte and a fixed

and optimized concentration of aptamer.

The main differences between the different formats are the

immobilized species (aptamer, antibody, or target analyte), the num-

ber of experimental steps involved, and in which order the differ-

ent reagents are exposed to the surface. The choice of the format

depends on the molecular size of the analyte, the availability of

reagents, and the cost. The main advantages involved in the use of

a sandwich format are the selectivity and sensitivity of the assay.

When it is possible to perform different assay formats for the detec-

tion of the same target analyte, it is useful to compare the analytical


Electrochemical Detection Strategies Based on Labeling 31

performances of each, in order to choose the approach that is the

best compromise in terms of sensitivity, specificity, analysis time,

and costs.

The high sensitivities requested by the aptamer-based assays for

the detection of the target analytes cannot be reached by a “direct

format,” since the affinities of aptamers for their targets are not high

enough, ranging from the micro to the nanomolar level. For this pur-

pose, several strategies have been used as signal amplification tools,

such as metallic and magnetic nanoparticles (NPs), enzymatic labels,

and quantum dots.

The potential use of aptamers as receptors in biosensors and

bioassays has been extensively reviewed [1–7] and also several

books have appeared in the last years [8, 9]. In this chapter, the cur-

rent status of research in electrochemical aptasensors is considered.

Attention is focused on label-free and labeled aptasensors, and the

analytical capabilities of these devices are discussed.

2.2 Electrochemical Detection StrategiesBased on Labeling

Labels such as enzymes, NPs, and redox species, such as fer-

rocene (Fc) or methylene blue (MB), are often used for recognition

processes. Electrochemical aptasensors with a label have received

and yet continue to receive considerable attention because they

combine the specificity of the aptamer–analyte recognition to the

advantages of an amplified signal. These strategies are generally

highly sensitive due to the analytical characteristics of the label used.

Labels are commonly covalently linked to terminal groups of

aptamers. The labeling position has to be carefully chosen so as not

to interfere with the folding of the aptamer and, thus, not to lose the

bioactivity or stability.

Among the most used labels are enzymes, such as peroxidise

(HRP), glucose oxidase (Gox), and alkaline phosphatase (AP), that

generate an electroactive product close to the transducer surface;

the formation of a relatively high local concentration of the enzyme

product leads to a significant signal amplification.



Electrochemical aptasensor based on the use of a label has found

various applications. In the following sections, some examples are

discussed.

2.3 Electrochemical Aptasensors Based ona Sandwich Assay

The use of a sandwich format allows detecting the target ana-

lyte with very high sensitivity and selectivity. Two conditions are

required: (1) the analyte possesses two epitopes which are so dif-

ferent that both ligands can bind to the analyte without the binding

of one affecting the binding of the other, and (2) two aptamers are

selected against such analyte. The disadvantage related to this for-

mat consists of several incubation steps that make the assay time

consuming.

This format is widely used in the case of large molecules such

as proteins and hormones; in particular it has been applied to

the detection of thrombin, which has been mostly used as model

system. Thrombin is an important serine protease in the blood

coagulation cascade. It contains a heparin-binding exosite and

fibrinogen-recognition exosite. In 1992, the first DNA thrombin

aptamer was isolated by Bock and coworkers [10] and the most

active strand was a 15-mer oligonucleotide with a K d around

100 nM. This aptamer interacts with the fibrinogen-recognition

exosite. The other thrombin-binding aptamer selected by Tasset and

coworkers [11] is a 29-mer single-stranded DNA with a K d around

0.5 nM. This aptamer binds to the heparin-binding exosite of throm-

bin. RNA aptamers for thrombin have also been selected.

Ikebukuro et al. [12] first reported an electrochemical aptasen-

sor for the detection of thrombin based on a sandwich-based assay.

Two different aptamers specific for thrombin were used: the 29-

mer thiolated aptamer and the 15-mer aptamer labeled with glu-

cose dehydrogenase (GDH). The thiolated aptamer was immobilized

onto gold electrodes; thrombin at different concentrations and then

the enzyme-labeled aptamer was added to the aptamer-modified

electrodes. The electric current generated by the addition of glu-

cose was measured at 0.1 V vs. Ag/AgCl in a buffer containing


Electrochemical Aptasensors Based on a Sandwich Assay 33

Streptavidin-coated magnetic bead

Streptavidin-alkaline phosphatase conjugate

5’ biotinylated aptamer

5’ biotinylated secondary aptamerThrombin

Working electrode

Magnetic bar

Figure 2.2. Schematic representation of the electrochemical sandwich

assay performed for the detection of thrombin.

1-methoxyphenazine methosulfate (m-PMS). Using this approach,

10 nM of thrombin was detected.

Centi et al. [13] developed an electrochemical aptamer-based

sandwich assay for analysis of thrombin in complex matrices, using

a simple-target capturing step by aptamer functionalized magnetic

beads. The assay was carried out by immobilizing the 15-mer

biotinylated aptamer on streptavidin-coated magnetic beads and

then incubating the modified beads with the target analyte and with

the 29-mer biotinylated aptamer (Fig. 2.2). At this point, a solution

of the conjugate streptavidin-alkaline phosphatase was added to the

beads and, after streptavidin-biotin recognition the enzymatic sub-

strate (1-naphthyl phosphate) solution was added: the enzymatic

substrate was converted by AP into 1-naphthol, which was oxidized

at the working electrode surface. The amount of oxidized naphthol

was quantified by differential pulse voltammetry. The assay was

applied to the analysis of thrombin in buffer [detection limit (DL)

found was 0.45 nM], spiked serum, and plasma with similar ana-

lytical performances. Moreover, thrombin was generated in situ in

plasma by the conversion of its precursor prothrombin, and the for-

mation of thrombin was followed at different times.



An example of electrochemical sandwich assay for detection of

immunoglobulin E (IgE) is reported [14]. The assay was performed

by coupling an antibody against IgE with an aptamer anti-IgE.

Antibody molecules were covalently immobilized as capture probe

on gold electrodes via a self-assembled monolayer of cysteamine

through the glutaraldehyde-based bifunctional linking. After that the

target was captured, the biotinylated anti-IgE aptamer was added

because it could interact specifically with the analyte, followed by

the addition of a streptavidin-alkaline phosphatase (streptavidin-

AP) solution. Once biotin-streptavidin recognition occurred, the

signal amplification was achieved based on enzymatic silver depo-

sition. Ascorbic acid 2-phosphate was converted by AP into ascor-

bic acid, a strong reducing agent. It could reduce the silver ions in

the solution to silver metal that preferentially deposited on the gold

surface of electrodes. The amount of deposited silver, which is pro-

portional to the amount of IgE target bound on the electrode surface,

was quantified by linear sweep voltammetry.

The detection limit calculated using this approach was 0.02 nM of

thrombin. A similar approach was reported by Degefa and cowork-

ers [15] as well.

2.4 Electrochemical Aptasensors Based ona Competitive Assay

In literature no many examples of electrochemical enzyme-labeled

aptasensors based on a competitive assay are present. The advan-

tages of a competitive format (direct and indirect) are mainly related

to the fact that only one aptamer is required (considering that two

or more aptamers are not selected for many target analytes) and the

time necessary for the assay is faster.

A disposable electrochemical competitive assay for detection

of IgE was proposed by Papamichael et al. [16]. In this work the

IgE antigen was immobilized on the surface of screen-printed elec-

trodes, then a competition step between IgE bound to the electrode

surface and IgE in solution for the biotinylated aptamer was left to

occur. At this point the streptavidin-alkaline phosphatase conjugate


Electrochemical Aptasensors Based on a Competitive Assay 35

was added to the electrode. p-Aminophenyl phosphate (p-APP) was

used as enzymatic substrate and differential pulse voltammetry as

electrochemical technique. The results obtained with the electro-

chemical aptasensor were compared with those based on an ELISA-

type assay (ELONA). The aptasensor showed high specificity and

selectivity toward IgE with a detection limit of 23 ng/mL, which is

a concentration sufficient for detection of IgE in blood considering

that the IgE concentration in blood samples of healthy subjects is in

the range 240 to 290 ng/mL. Moreover, in this work the stability of

the assay performed using the aptamer against IgE was compared

with that of the assay carried out with a monoclonal antibody spe-

cific for IgE. Authors reported that the assay performed using the

aptamer was more stable than that with antibody, considering that

aptamers can be easily regenerated by also using harsh conditions

and are thermo-stable, because the aptamer folding is not affected

by the temperature.

Among the several strategies reported by Mir et al. [17] for the

detection of thrombin, an electrochemical aptasensor based on a

competitive assay resulted to be the most sensitive. Thrombin was

immobilized on gold-mercaptoethanol–treated electrodes by pas-

sive adsorption and then the modified electrodes were incubated

with a biotinylated aptamer anti-thrombin. The sensor was subse-

quently incubated with streptavidin-horseradish peroxidase conju-

gate, which bound to the biotin on the aptamer. The aptamer was

quantified by the electrochemical detection of the reaction catalyzed

by the peroxidase. Hydrogen peroxide was used as oxidizing agent

and [Os(bpy)2(pyr-CH2–NH2)]Cl as mediator. In this case the limit

of detection of thrombin was 3.5 nM. Thrombin was immobilized by

direct adsorption also on bare gold electrodes and on polystyrene

surfaces but it was not detectable on these unmodified surfaces.

It was supposed that in these surfaces the adsorption position of

thrombin created steric impediments preventing the subsequent

binding with the aptamer; alternatively, the binding of thrombin to

the surfaces may have denatured the protein.

Centi et al. [18] described various approaches for the devel-

opment of electrochemical aptasensors for the detection of

thrombin using magnetic beads as solid support and carbon



screen-printed electrodes as electrochemical transducers; among

the developed assay formats, a direct and an indirect competitive

assay was reported. The experimental work was also supported by

the use of the surface plasmon resonance (SPR) device Biacore XTM,

through which different information on the tested assay formats was

obtained. For the direct competitive assay the biotinylated 15-mer

aptamer was immobilized on streptavidin-coated magnetic beads,

whereas in the case of the indirect competitive assay, thrombin and

biotinylated thrombin was immobilized on the tosyl-activated and

streptavidin magnetic beads, respectively. Using a direct competitive

format, a detection limit of 430 nM of thrombin was achieved and a

good specificity of the assay was found using human serum albumin

as an interfering molecule.

Impedance spectroscopy, in this case faradic impedance spec-

troscopy (FIS), was used as transduction technique for a competi-

tive aptamer-based assay for the detection of neomycin B [19]. The

interesting feature of this work is the possibility of easily detecting a

small molecule like neomycin B with an electrochemical aptamer-

based assay. Actually, the detection of small molecules, such as

aminoglycoside antibiotics, is particularly challenging since only

time-consuming label-based immunoassays or HPLC methods have

been developed. On the contrary, in the present work an aptamer

specific for neomycin B [20] was used in a competitive/displacement

assay format. In particular, neomycin B was immobilized onto gold

electrodes and this modified surface was saturated with the specific

aptamer by affinity binding. By exposing the modified system to dif-

ferent concentrations of neomycin B, a displacement of the bound

aptamer was observed resulting in a drop of the electron-transfer

resistance consistent with the reduction of the negative charge of the

electrode surface. The competitive assay resulted very fast (equilib-

rium in 5 minutes) with a linear range covering two orders of mag-

nitude (0.75–500 μM) and a submicromolar limit of detection.

Very high specificity toward neomycin B was observed with

respect to other very similar antibiotics. Application of the method

to the analysis of real samples was also demonstrated by testing

neomycin spiked whole milk with a recovery of 102% and 109%,

respectively, for two different neomycin concentrations.


Electrochemical Aptasensors Based on a Direct Assay 37

2.5 Electrochemical Aptasensors Based on a Direct Assay

Some papers have used the catalytic activity of thrombin for the

determination of this protein. α-Human thrombin is a highly spe-

cific serine protease that catalyses the hydrolysis of the throm-

bin chromogenic substrate, β-Ala-Gly-Arg- p-nitroaniline producing

p-nitroaniline. The rate of yellow colored p-nitroaniline formation

can be followed by its UV absorption at 405 nm, or electrochemi-

cally by the reduction of its nitro group. The electrochemical detec-

tion offers benefits in terms of sensitivity and speed. When saturated

by enzyme substrate the formation rate of p-nitroaniline is propor-

tional to the enzyme concentration.

Mir et al. [17] first reported the detection of thrombin bound

to an aptamer selective for thrombin by the quantification of

p-nitroaniline produced by the enzymatic reaction catalyzed by

thrombin.

A mixed self-assembled monolayer was used for the aptamer

immobilization on the gold electrode. The aptamer-modified elec-

trodes were then incubated for 1 h at 37◦C with thrombin

(18 μg/mL). Electrochemical measurements were recorded in the

thin-layer cell configured to contain a total volume of 20 μL.

Thrombin chromogenic substrate (β-Ala-Gly-Arg- p-nitroaniline)

was injected into the cell and differential pulse voltammetry (DPV)

measurements between −0.2 and −1 V with a pulse height of −0.05

V and pulse duration of 70 ms were carried out. The DPV measure-

ments showed that β-Ala-Gly-Arg- p-nitroaniline substrate and the

p-nitroaniline product have different redox potentials. Moreover, the

DPV experiments showed a current peak at −0.45 V in the pres-

ence of the thrombin substrate. After 5 min, the peak at −0.45 V

decreased and a new peak was detected at −0.70 V, indicating the

formation of p-nitroaniline. The same measurements carried out on

a control electrode in order to test the specificity of the assay: in this

experiment bovine serum albumin (BSA) substituted thrombin and

in this case only the peak at 0.45 V was measured.

The authors demonstrated a huge reduction in the assay time

considering that the optical detection of p-nitroaniline needed 3 h

against 5 min necessary for the electrochemical measurement.



Centi et al. [18] compared the performances of several assay for-

mats based on the coupling of magnetic beads with electrochemi-

cal transduction always for the detection of thrombin. Among the

developed assays, one of the used strategies was based on the

direct measurement of the enzymatic product of thrombin captured

by the immobilized aptamer. The main differences between this

work and the work by Mir et al. [17] involve the use of magnetic

beads as solid support on which the aptamer-based assay is per-

formed. Streptavidin-coated magnetic beads modified by immobi-

lization of the biotinylated thrombin aptamer were incubated with

different concentrations of thrombin in the range 100 to 600 nM

for 30 min. Bound thrombin was detected by re-suspending the

beads in the thrombin substrate, β-Ala-Gly-Arg- p-nitroaniline, for

30 min at 37◦C. The solution containing the thrombin reaction prod-

uct was deposited onto the surface of the working screen-printed

graphite electrode, without any stirring. The aptamer-bound throm-

bin was detected by quantification of p-nitroaniline produced from

the thrombin catalyzed reaction. The DPV measurements showed a

decrease of the peak at −730 mV vs. Ag/AgCl pseudo-reference elec-

trode related to the β-Ala-Gly-Arg- p-nitroaniline substrate and the

appearance of a new peak at −870 mV vs. Ag/AgCl pseudo-reference

electrode, indicating the formation of p-nitroaniline (Fig. 2.3). The

same measurements were carried out in absence of thrombin, and

only a reproducible peak at −730 mV was observed (16.1±0.4 μA).

A linear increase of p-nitroaniline peak current was observed in the

studied concentration range of thrombin. On the contrary, a linear

decrease in thrombin substrate was observed increasing the throm-

bin concentration. The detection limit (DL) found for thrombin using

this approach was 175 nM.

In another direct approach, non-faradic electrochemical

impedance spectroscopy (NIS) was used for the direct detection of

platelet-derived growth factor-BB (PDGF-BB) [21]. Binding of PDGF

to its aptamer immobilized on a silicon electrode surface leads to

a decrease in capacitance measured by electrochemical impedance

spectroscopy (NIS). Because of the high sensitivity and specificity

(DL 40 nM) and the absence of reagent to be used when performing

the test, this biosensor design could be promising for in vivo moni-

toring.


Electrochemical Metal Nanoparticle-Labeled Aptasensors 39

Figure 2.3. DPV scans of the thrombin substrate (β-Ala-Gly-Arg- p-

nitroaniline) solutions incubated with aptamer-thrombin modified beads.

Different concentrations of thrombin in the concentration range 100 to

600 nM were incubated with the aptamer-modified beads, while a fixed con-

centration of thrombin substrate was used (200 μM). The thrombin sub-

strate and the p-nitroaniline released during hydrolysis showed different

redox potentials (the DPV peak potential of β-Ala-Gly-Arg- p-nitroaniline

was −730 mV vs. Ag/AgCl pseudo-reference electrode, whereas the

released p-nitroaniline peak potential was −870mV vs. Ag/AgCl pseudo-

reference electrode).

2.6 Electrochemical Metal Nanoparticle-LabeledAptasensors

The use of NP labels is a strategy relatively new in the develop-

ment of electrochemical aptasensors. The labels used are essentially

metallic NPs or inorganic (semiconductor) nanocrystals [22–25].

They allow developing elegant strategies for interfacing aptamer-

target analyte recognition events with electrochemical transduction

amplifying the resulting electrical response and thus giving rise to an

improvement of the assay sensitivity. In particular, the redox prop-

erties of gold NPs have made possible their widespread use as elec-

trochemical labels in aptasensor development [24]. Most of these

strategies involve a stripping measurement of the metal tag: metal

NPs can be oxidized to form the corresponding metal ions that can

be then detected electrochemically.



Figure 2.4. Scheme of the analytical procedure based on the use of Pt-NPs

for the analysis of thrombin.

The high sensitivity of such measurements is based on the pre-

concentration step, during which the metals are electrodeposited

onto the working electrode [26]. It is important for the minimiza-

tion of the non-specific adsorption and for the corresponding back-

ground signal. For this reason, surface blocking steps should be

employed to avoid the amplification of the background signal. More-

over, the control of the coverage allows to ensure a good accessibility

and stability of the surface bound probe.

The first electrochemical aptasensor using NPs was reported by

Polsky [27] for the detection of thrombin. A sandwich configura-

tion was designed and for this purpose thiolated aptamer molecules

were immobilized on gold-covered slide, then the aptamer-modified

surface was incubated first with thrombin and at the end with

aptamer-modified Pt-NPs (aptamer-Pt-NPs) (Fig. 2.4). The Pt-NP

labels associated with the thrombin were then used as sites for the

electrocatalytic reduction of H2O2 that was added to the working

medium before analysis and linear sweep voltammetry was used as

electrochemical technique. The reduction of hydrogen peroxide gave

rise to a cathodic current which directly related to the concentration

of thrombin. The detection limit found using this method was 1 nM

of thrombin.

Another example of electrochemical aptasensor based on Au-

NPs as labels for the detection of thrombin is reported by Zheng

et al. [28]. The assay was based on a sandwich format, in which the

aptamerI (15-mer DNA aptamer with an amino group at its 5’ end)

was immobilized onto carboxyl functionalized magnetic beads. Such

aptamer-coated magnetic beads were used for capturing and sepa-

ration. Thrombin and Au-NP–labeled aptamerII were then added to


Electrochemical Metal Nanoparticle-Labeled Aptasensors 41

the modified magnetic beads and after an incubation step, the excess

of reagents was removed by magnetic separation. In addition, for the

signal amplification, thiocyanuric acid (3.5·10 − 6μM) was added

before incubation and the magnetic bead–aptamerI/thrombin/Au-

NP–aptamerII conjugates were re-suspended in HCl 0.1 M solution.

A scheme of the sandwich assay is shown in Fig. 2.4.

A signal amplification was obtained by forming a network of thio-

cyanuric acid/Au-NPs. The electrochemical oxidation of Au-NPs was

performed at +1.25 V for 120 s on a glassy carbon electrode. Imme-

diately after the electrochemical oxidation, differential pulse voltam-

metry was performed resulting in an analytical signal due to the

reduction of AuCl4−, which relates to the amount of the Au-NPs for

the sandwich format.

A detection limit of 8 aM was achieved. To demonstrate the feasi-

bility of this approach, the aptasensor was applied to the detection

of thrombin in some plasma samples.

Hansen and coworkers [29] reported an electrochemical

aptasensor involving nanocrystal tracers for the detection of throm-

bin. The aptasensor was based on a displacement assay (Fig. 2.5).

Thiolated-aptamers specific for thrombin and lysozyme were immo-

bilized on a gold electrode. Thrombin and lysozyme were modified

with CdS and PbS quantum dots (QDs), respectively and these were

bound to the respective aptamers immobilized on the surface. In the

presence of the target protein, the QD-tagged protein was displaced

and the number of QDs left on the surface decreased. After dissolv-

ing the remaining QDs on the surface using 0.1 M HNO3, the metal

ions (Cd2+ and Pb2+) were identified and their concentration at mer-

cury coated glassy carbon electrode was detected by electrochemi-

cal stripping. The concentration of the metallic ions was correlated

with the concentration of the target proteins in solutions. Owing to

the amplification effect originated by dissolving QDs and by the high

sensitivity correlated to the electrochemical stripping detection, a

detection limit of 0.5 pM was achieved for thrombin. It is impor-

tant to underline that, using this approach, different aptamers could

be immobilized on the same gold substrate, since different protein

targets can be labelled with QDs with different cation compositions

(CdS, ZnS, CuS and PbS). As demonstrated by authors, thrombin and

lysozyme were labeled with CdS and PbS and both proteins were

simultaneously detected on the same gold substrate.



(C)

(B)

(A)

Thrombin Lysozyme

Wash

0.1M HNO3Cd Pb

Cd2+ & Pb2+

Figure 2.5. Schematic representation of the displacement assay. (A) Mixed

monolayer of thiolated aptamers on the gold substrate with the bound

protein-QD conjugates (B) Sample addition and displacement of the tagged

proteins (C) Dissolution of the remaining captured nanocrystals followed

by their electrochemical-stripping detection at a coated glassy carbon elec-

trode.

Another aptasensor for the detection of thrombin has been

recently reported [30]. It is based on a sandwich assay and on the use

of NP labels. In this work an interesting aspect concerns the coupling

of signal amplification due to NPs with the preparation of a nanogold

electrode by electrochemical deposition of Au-NPs on a gold elec-

trode. The surface of the gold electrode was pretreated as follows:

heated in a piranha solution for about 5 min, polished with alu-

mina slurries, washed ultrasonically with water, dried with nitrogen

gas, and cycled in 0.5 M H2SO4 aqueous solution scanning between

0.3 and 1.5 V until a stable gold oxide formation/reduction cyclic


Electrochemical Aptasensors Based on Noncovalent Redox Species Label 43

voltammogram was obtained. The electrochemical deposition of Au-

NP was carried out in the HAuCl4 solution containing 0.1 M KNO3

as electrolyte at −400 mV. The freshly prepared nanogold electrode

was incubated with the thiolated 15-mer aptamer anti-thrombin for

about 16 h to produce an aptamer attached electrode. Then the

modified electrode was immersed in a solution of 6-mercapto-1-

hexanol for 1 h to block the uncovered gold surface. At this point,

the aptamer-modified electrode was interacted with different con-

centrations of thrombin and then with a solution of Au-NP probe.

It consisted of Au-NPs conjugated to the thiolated 15-mer aptamer

anti-thrombin and to CdS-NPs linked with a single-stranded DNA

sequence. The resulting sandwich complex was treated with 1.0 M

of HNO3 solution for 5 min to dissolve the CdS-NPs and then with

acetate buffer containing Hg2+. The DPV measurements of the dis-

solved Cd2+ were performed using an in situ prepared mercury film

on a glassy carbon electrode with a deposition time of 300 s and

deposition potential of −1.1 V. An anodic stripping peak current at

−0.67 V was taken as the analytical response.

A detection limit of 0.55 fM of thrombin was calculated. Authors

attribute the significant improvement of the sensitivity of such

aptasensors with respect to others present in literature to the use

of a nanoelectrode, formed by immobilization of Au-NPs on the sur-

face of a gold electrode, to the use of NPs as labels, and to the use of

DPV technique for the detection of the dissolved Cd2+ in the solution.

Moreover, the electrochemical aptasensor was successfully tested in

some serum samples.

2.7 Electrochemical Aptasensors Based on NoncovalentRedox Species Label

These aptasensors are based on the use of a redox probe such

as methylene blue (MB) that undergoes an oxidation and reduc-

tion due to the electron transfer from an electrode surface to a

probe. These redox probes are noncovalently bound to aptamers and

intercalate or interact with aptamers mainly by electrostatic inter-

actions. For example, MB, positively charged, interacts with nega-

tively charged proteins or other negatively charged analytes. When



the analyte is bound to the aptamer molecules immobilized on the

electrode surface and interacts also with MB, an increased redox

current is recorded. Hianik et al. [31] first reported an electrochem-

ical aptasensor for the detection of thrombin based on the inter-

action of MB with the aptamer-thrombin complex. A biotinylated

DNA aptamer was immobilized on a gold electrode via streptavidin-

biotin interactions. When thrombin was bound to the immobilized

aptamer and MB interacted with thrombin, measurable changes of

charge transfer measured by differential pulse voltammetry (DPV)

were obtained. However, since MB can also non-specifically bind to

the DNA aptamer and streptavidin, the background signal and sig-

nal changes were high and the detection limit of thrombin obtained

using this approach was relatively low (10 nM).

Recently, an aptasensor based on a redox probe ([Ru(NH3)5Cl]2+)

was developed for the detection of platelet-derived growth fac-

tor (PDGF) [32]. A sandwich assay format was carried out, since

PDGF has two aptamer-binding sites, which made it possible for

one PDGF molecule to connect with two aptamers simultaneously.

Gold electrodes were modified by immobilization of a thiolated

aptamer against PDGF; then the aptamer-modified electrodes were

incubated first with different concentrations of PDGF and then with

aptamer-loaded Au-NPs. [Ru(NH3)5Cl]2+ molecules, which were fur-

ther immobilized onto the surface of the above “sandwich” structure

(Fig. 2.6), were used as redox probes and a suitable concentration of

the redox probe was optimized. Cyclic voltammetry measurements

were performed. The authors reported that the sandwich format and

the use of Au-NPs allowed to amplify the signal of the redox probe

allowing to obtain a very low detection limit (1 × 10−14 M for puri-

fied samples). The aptasensor was successfully applied to the analy-

sis of PDGF in serum samples.

Another commonly used redox probe is Fe(CN)3−/4−6 which has

been coupled to different electrochemical techniques as summa-

rized in Table 2.1.

A very recent example of the use of this redox probe in an

aptamer-based biosensor was published by Kim et al. [33]. An

electrochemical biosensor for oxytetracycline detection was devel-

oped using ssDNA aptamer immobilized on gold interdigitated array

(IDA) electrode chip (Fig. 2.7). Cyclic voltammetry and square wave

voltammetry were used to measure the current at the electrode chip


Electrochemical Aptasensors Based on Noncovalent Redox Species Label 45

Figure 2.6. Schematic representation of the electrochemical aptasensor

based on a sandwich assay and on the use of [Ru(NH3)5Cl]2+ as redox probe.

Table 2.1. Examples of Aptamer-Based Electrochemical Biosensors Based

on the Use of Fe(CN)3−/4−6 as Redox Probe

Target Electrochemical Technique Analytical Characteristics References

Oxytetracycline Cyclic voltammetry Square DL 5 nM Range 1–100 nM Kim et al. (2009)

wave voltammetry

17b-estradiol Cyclic voltammetry Square Linear range 0.01–1 nM Kim et al. (2007)

wave voltammetry

Thrombin Impedance spectroscopy Range 0.5–500 nM Lee et al. (2008)

Cancer cells Impedance spectroscopy DL 6 × 103 cells/mL Pan et al. (2009)

Thrombin Impedance spectroscopy DL 0.01 nM Range 1–50 nM Zhang et al. (2009)

Adenosine Cyclic voltammetry DL 1 nM Range 0.1–100 nM Zheng et al. (2008)

Adenosine Impedance spectroscopy DL 0.1 nM LI et al. (2007)

Cocaine Impedance spectroscopy DL 5 nM Elbaz et al. (2008)

AMP Impedance spectroscopy DL 10 nM Elbaz et al. (2008)

due to the presence of [Fe(CN)3−6] in solution. A decrease in current

was evident after the binding of oxytetracycline to the aptamer: this

was probably due to the changes in the conformation of the aptamer

which caused changes in permeability and in charges on the elec-

trode. The biosensor could detect oxytetracycline in the range 1

to 100 nM with high specificity since negligible interference was

present when analyzing structurally similar antibiotics such as doxy-

cycline and tetracycline.



Figure 2.7. The electrochemical detection system for oxyteracycline

(OTC) using aptamer-immobilized interdigitated array (IDA) gold electrode

chip. Left: An IDA gold electrode chip and the IDA gold electrode. Right: A

typical electrochemical reaction occurring after aptamer binds to its target

molecole.

A very interesting biosensor was developed based on a similar

principle for the detection of cancer cells [34]. The aptamer selected

for acute leukaemia cells was fixed onto a gold electrode and elec-

trochemical impedance spectroscopy (EIS) technique was used to

characterize the surface with [Fe(CN)6]3−/4− as a redox probe. Upon

binding of the aptamer-modified electrode with leukaemia cells,

the electron-transfer resistance of [Fe(CN)6]3−/4− on the sensor

surface increased substantially. A linear relationship was observed

between the electron-transfer resistance and the concentration of

the leukaemia cells in a range 1 × 104 to 1 × 107 cells/mL with a

detection limit of 6 × 103 cells/mL and high selectivity.

2.8 Electrochemical Aptasensors Based on the AptamerConformational Changes

Aptamers bind their targets through adaptive recognition; in solu-

tion aptamers are unstructured, but fold upon associating with their

molecular targets into molecular architectures in which the ligand


Electrochemical Aptasensors Based on the Aptamer Conformational Changes 47

Table 2.2. Aptamer-based biosensors based on the “switch-on” or

“switch-off” approach

Target Signal Label Analytical Characteristics References

Cocaine ON MB Baker et al. (2006)

Thrombin ON Ferrocene DL 0.5 nM Radi et al. (2006)

Thrombin OFF MB DL 3 nM Xiao et al. (2005)

Adenosine OFF Ferrocene DL 0.02 μM Wu et al. (2007)

Lysozyme OFF [Ru(NH3)6]3+ 0.5 μg/mL Cheng et al. (2007)

Theophylline ON MB DL 2 μM Feropontova et al. (2009b)

Thrombin ON Ferrocene Picomolar range Huang et al. (2008)

Thrombin ON Ferrocene DL 30 fM Mir et al. (2008)

Cocaine ON MB Low micromolar Swensen et al. (2009)

Thrombin ON Glucose/Glucose DL 2.5 nM Tan et al. (2009)

oxidase

Thrombin OFF Ferrocene DL 3.9 nM Tan et al. (2009)

Adenosine OFF MB DL 0.01 μM Wang et al. (2009)

Botulinum OFF Fluorescein/anti- DL 40 pg/mL Wei et al. (2009)

neurotoxin fluorescein-HRP

PDGF ON MB DL 50 pM Rodriguez et al. (2005)

becomes an intrinsic part of the nucleic acid structure [35]. This

feature represents an almost unique mechanism that can be

exploited in the design of new electrochemical biosensors [2]. In

this approach the interaction of a labeled aptamer with its tar-

get can modulate the distance of the electroactive labels from the

sensor electrode, thereby altering the redox current. Various

aptasensors (Table 2.2), based on this approach, are used for the

detection of different targets such as theophylline [36, 37], lysozyme

[38], botulinum neuorotoxin [39], adenosine [40, 41], cocaine [42],

or thrombin [43–45].

In the two studies using a “signal-off” approach by Xiao et al.[44, 45], thrombin was detected by monitoring the decrease in the

amperometric response of a redox label present at one end of the

thrombin aptamer as a result of the association of thrombin with

the aptamer. The interaction of the labeled aptamer with its target

modulates the distance of the electroactive labels from the sensor

electrode, thereby altering the redox current. In the absence of the



target the aptamer is in an unfolded conformation, allowing rapid

interaction between the MB redox label and the electrode. Upon

target binding the aptamer forms a stable structure, reducing the

distance between the label and the electrode, therefore, reducing the

electron transfer between MB and the electrode.

An alternative strategy has been developed, based on a “signal-

on” configuration [46]. A thiolated thrombin aptamer modified at

the non-thiolated end with a ferrocene group was immobilized

onto a polycrystalline gold electrode. The long and flexible aptamer

chain prevented contact between the redox label and the electrode,

inhibiting the generation of the electrochemical signal. The binding

of thrombin to the aptamer caused the formation of the characteris-

tic G-quadruplex aptamer structure, orientating the ferrocene units

toward the electrode and leading to a positive amperometric signal.

The differential pulse voltammetry measurements demonstrated a

thrombin detection limit of 0.5 nM and a detection range between

5 and 35 nM. A similar “signal on” approach [47] utilizing a cocaine

aptamer and MB as electrochemical label was used for the detection

of cocaine.

More recently, a detailed study was conducted on an electro-

chemical biosensor based on a “signal- on” approach for the detec-

tion of theophylline [36, 37]. The RNA aptamer for theophylline

was first labeled with ferrocene and anchored to a gold electrode:

its conformation switching upon binding of theophylline caused

the formation of a folded structure with an increased electron

transfer between ferrocene and the electrode. In this approach

theophylline could be detected at the micromolar range, but the

biosensor response was inhibited in serum. The biosensor was then

optimized by substituting ferrocene with MB, shifting the redox

potential from positive to negative potential (−0.25 V vs. Ag/AgCl).

The modified biosensor could detect theophylline in the relevant

range 2 to 100 μM and the biosensor response in serum was sim-

ilar to the response in buffer.

A very interesting biosensor was developed for the detection

of cocaine by using a microfluidic electrochemical device [42]

(Fig. 2.8).

Cocaine at the micromolar range was detectable in continuous,

real-time, and in undiluted and untreated blood samples.


Electrochemical Aptasensors Based on Target-Induced Aptamer Displacement 49

Figure 2.8. (a) The microfluidic electrochemical aptamer-based sensor

chip. (b) The syringe pumps connected to a four-input, single-output mul-

tiplexed valve. (c) The detection mechanism of the switch on sensor.

2.9 Electrochemical Aptasensors Based onTarget-Induced Aptamer Displacement

This type of biosensor took advantage of the strong affinity of the

aptamer for its specific analyte and used a competition scheme as

the detection methodology [6].

In the target-induced strand displacement strategy, the aptasen-

sor is usually assembled by fixing a complementary DNA–aptamer

duplex on an electrode (Fig. 2.9). Upon binding to their target mole-

cules, the aptamers or complementary DNA are displaced from the

electrode, resulting in a significant change in the electrochemical sig-

nal. This strategy is particularly interesting since it is easy to gener-

alize for any aptamer without prior knowledge of its secondary or

tertiary structure, and it is well suited for the development of elec-

trochemical aptasensors.

This approach has been adopted for the development of biosen-

sors for the detection of different targets and some examples will be

presented here.

Faradic impedance spectroscopy (FIS) was the electrochemi-

cal technique used in an aptasensor for the detection of lysozyme

[48]. The duplex formed by the lysozyme aptamer and a partial



Figure 2.9. General scheme of an aptasensor based on target-induced

strand displacement.

complementary single-stranded DNA was fixed onto a gold elec-

trode. [Fe(CN)6]3−/4− was then used as a redox couple to monitor

the change in electron transfer at the electrode upon binding of the

target molecule, lysozyme. In the presence of lysozyme, the aptamer

was displaced from the duplex and the electron-transfer resistance

was decreased. This decrease could be monitored by FIS in a con-

centration range between 0.2 and 4 nM and with a detection limit of

0.07 nM.

In another work [49], the thrombin aptamer was hybridized with

a ferrocene-labeled DNA oligonucleotide and immobilized onto a

gold electrode. The binding of thrombin to the aptamer causes the

displacement of the complementary oligonucleotide resulting in a

decrease of current recorded at the electrode by differential pulse

voltammetry (DPV). A linear range for the detection of thrombin

between 0 and 10 nM was obtained.

Another aptasensor based on the displacement of a complemen-

tary strand from an aptamer was developed for the detection of ATP

by coupling this approach to signal amplification by Au-NPs [50]. In

this work the hybrid was formed by a reporter DNA labeled with

Au-NPs, a thiol-modified DNA anchored to an electrode and a target-

responsive DNA (the aptamer) (Fig. 2.10). Moreover, [Ru(NH3)6]3+


Electrochemical Aptasensors Based on Target-Induced Aptamer Displacement 51

Figure 2.10. The representation of the aptasensor based on NPs’ amplifi-

cation for ATP detection.

which binds to surface-confined DNA via electrostatic interaction

was used for signal generation, interrogated by chronocoulometry.

In presence of the target molecule, ATP, the binding to the

aptamer causes its conformational change leading to the release

of the reporter DNA labeled with Au-NPs. In this way numerous

molecules of [Ru(NH3)6]3+ are released in solution generating a sig-

nal amplification. A wide linear range for the detection of ATP was

obtained between 1 nM and 10 μM with a detection limit of 0.2 nM.

A similar approach was conducted by using quantum dots (QDs)

for signal amplification [51]. In this case the hybrid formed by a

thiol-labeled oligonucleotide and the thrombin aptamer was immo-

bilized onto a gold electrode. When binding to thrombin the aptamer

adopts its G-quartet structure and only the single-stranded probe

remained onto the electrode, which is now available for hybridiza-

tion with a QD-labeled complementary oligonucleotide (Fig. 2.11).

The CdS-QD were then dissolved and CD2+ was detected on a

mercury-film electrode: this technique led to a detection limit of

0.43 fM for thrombin with a linear range between 2.3 nM and 2.3 fM.

Other works based on this kind of approach without or with sig-

nal amplification were recently published for the detection of adeno-

sine [52, 53] and lysozyme [48].



Figure 2.11. Scheme of the aptasensor based on the formation of a hybrid

containing the thrombin aptamer, strand displacement, and signal amplifi-

cation via QD.

2.10 Conclusions

In this chapter several applications of aptamers in the development

of electrochemical biosensors have been reported. Different electro-

chemical methods based on aptamers have been considered both

for the detection of proteins (PDGF, VEGF, lysozyme, or thrombin)

or small molecules.

Aptamer-based assays opened new scenarios in the field of ana-

lytical chemistry. New combinations of aptamer-based biosensors

with innovative ideas in molecular biology, electrochemistry, and

nanotechnologies are encouraged and expected with aim of devel-

oping easy, sensitive, selective, and fast analytical methods.

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Chapter 3

Carbon-Polymer Bio-Nano-CompositeElectrodes for ElectrochemicalGenosensing

Marıa Isabel Pividori and Salvador AlegretGrup de Sensors i Biosensors, Departament de Quımica,Universitat Autonoma de Barcelona, Barcelona, [email protected]

This chapter reports the main features of rigid carbon–polymer

composite materials for electrochemical DNA biosensing. Novel

approaches based on composites modified with biomolecules (bio-

composites) and nanostructured materials (nanocomposites) for

the improved biosensing of DNA are also discussed.

3.1 Introduction

The use of nucleic acids as biorecognition elements represents an

exciting interdisciplinary area of research in converging technolo-

gies. The oriented and improved immobilization of single-stranded

DNA to solid substrates, followed by hybridization and detection

of this event, has gained importance over the past decade, due

to the growing demand for genetic information in an increasingly



58 Carbon-Polymer Bio-Nano-Composite Electrodes for Electrochemical Genosensing

broad range of disciplines. The Human Genome Project (HGP) [1]

has stimulated the development of analytical methods that yield

genetic information quickly and reliably. Examples of this develop-

ment are the DNA chips [2–4], lab-on-chips based on microfluidic

techniques [5, 6], and self-assembled molecular electronic circuits

[7]. The development of new methodologies possessing the conve-

nience of solid-phase reaction, along with the advantages of rapid

response, sensitivity, and ease of multiplexing, is now a challenge in

the development of new bioanalytical diagnostic tools. Electrochem-

ical DNA biosensors can meet these demands, offering consider-

able promise for obtaining sequence-specific information in a faster,

simpler, and cheaper manner compared to traditional hybridization

assays. Such devices possess great potential for applications, rang-

ing from decentralized clinical testing, to environmental monitoring,

food safety, and forensic investigations.

The development of new transducing materials for DNA analy-

sis is a key issue in the current research efforts of electrochemical-

based DNA analytical devices. While DNA immobilization and

detection of the hybridization event are important features, the

choice of a suitable electrochemical substrate is also of great impor-

tance in determining the overall performance of the analytical

electrochemical-based device, especially regarding the immobiliza-

tion efficiency of DNA.

Carbonaceous materials such as carbon paste [8], glassy carbon

[9], and pyrolitic graphite [10] are the most popular choices of elec-

trodes used in biosensing devices. However, the use of platinum

[11], gold [12], indium-tin oxide [13], solid copper amalgam [14],

mercury [15], and other continuous conducting metal substrates has

been reported [16]. Conducting polymers—such as polypyrrole and

polyaniline—[17] and conducting composites—based on the com-

bination of non-conducting polymers with conductive fillers—[18]

have also been continuously studied during the past few decades.

Finally, nanostructured materials such as carbon nanotubes (CNT)

[19] and metal nanoparticles (NPs) [20] have also been reported as

a base material or fillers for conducting composites or as surface-

modifiers of many types of electrochemical transducers in order

to improve their electrochemical properties. Other nanostructured

materials including gold NPs have been intensively investigated as


Introduction 59

a component of electrochemical transducers [21]. Nanocompositescan be fabricated not only with nanostructured materials, but also

with biomolecules and redox polymers to achieve unique hybrid

and synergistic properties. It is expected that the combination of

nanoengineered “smart” polymers with novel biocompatible nanos-

tructured fillers—like NPs and CNT—may generate composites with

new and interesting properties, such as higher sensitivity and sta-

bility of the immobilized molecules, thus constituting the basis for

improved electrochemical biosensors.

The immobilization of the oligo probe—which specifically recog-

nizes the DNA target—onto the transducer is also a key issue in the

construction of biosensing devices. The choice of the immobilization

method depends mainly on the biomolecule to be immobilized,

the nature of the solid surface, and the transducing mechanism

[22]. Besides the sensitivity, the ability of the electrochemical trans-

ducer to provide a stable immobilization environment while retain-

ing the bioactivity must also be considered: a current problem

regarding the immobilized biomolecules is the lack of stability and

activity in the solid transducer, which is usually overwhelmed by

mimicking in vivo-like environment or the use of spacer arms.

The most successful immobilization methods involve (i) multisiteattachment, either electrochemical—by the application of a poten-

tial to the solid support—or physical adsorption, or (ii) single-pointattachment—mainly covalent immobilization, affinity linkage such

as strept(avidin)/biotin binding [23]) and chemisorption based on

self-assembled monolayers (SAMs) [16].

Among the different immobilization strategies, multisite adsorp-tion is the simplest and most easily automated procedure, avoiding

the use of pretreatment procedures based on previous acti-

vation/modification of the surface transducer and subsequent

immobilization. Such pretreatment steps are known to be tedious,

expensive, and time-consuming. Furthermore, the adsorption prop-

erties of DNA on various supports (e.g., nylon, nitrocellulose) have

been known for a long time [24]. The binding forces involving physi-cal adsorption include hydrogen bonds, electrostatic interaction, van

der Walls forces, and hydrophobic interactions if water molecules

are excluded by dryness [25]. Wet adsorption originates a weak

binding that causes easy desorption of the biomolecule from the



surface, eventually leached to the sample solution during measure-

ments. However, dry adsorption also promotes hydrophobic bonds

and more stable adsorbed layers on solid surfaces [25]. Classical

strategies such as physical entrapment in membranes and crosslink-

ing by bifunctional reagents—such as glutaraldehyde—also repre-

sent multisite attachment methods for retaining the biorreceptor in

close contact of the transducer.

Single-point attachment is beneficial for the kinetics of the bio-

logical reaction, especially if a spacer arm is used. Single-point cova-lent immobilization can be performed on different surface-modified

electrochemical transducers, such as glassy carbon [26, 27], carbon

paste [28], gold [29], or platinum [11], or, lately, carbon nanotubes

[30] through the linkage of a –COOH with a –NH2 group by the use

of the carbodiimide chemistry. Single-point affinity linkage also pro-

vides an interesting strategy for the oriented and stable immobi-

lization of biotinilated biomolecules to solid transducers throughout

biotin/strept(avidin) binding [23]. Finally, chemisorption based on

SAMs has also been extensively used for single-point attachment on

gold-based transducers [31, 32].

Electrochemical detection of the DNA hybridization event should

also be considered, involving the transduction of the hybridization

event into a useful and easy-to-amplify electrical signal. The DNA

recognition event for electrochemical transducing can be detected

mostly by means of external electrochemical markers such as elec-

troactive indicators [33, 34] or enzymes. Enzyme labeling has been

transferred from non-isotopic DNA classical methods to electro-

chemical genosensing. The enzyme labeling relies on the reaction

between a small tag (usually biotin or digoxigenin modified DNA

probe) with the streptavidin [35] or anti-digoxigenin [36] enzyme

conjugates, respectively. Although a second incubation step is usu-

ally required for labeling, higher sensitivity and specificity have been

reported for the enzyme labeling method compared with the other

reported methods [37]. The use of metal NPs—especially gold NPs

[39–42]—as labels for biosensing devices are also gaining impor-

tance. The direct electrochemical detection of DNA was initially pro-

posed by Palecek [43, 44] who recognized the capability of both

DNA and RNA to yield reduction and oxidation signals after being

adsorbed. The oxidation of DNA was shown to be strongly depen-


Composites Materials 61

dent on both the DNA adsorption conditions as well as the substrate

on which DNA is being adsorbed, thus requiring a meticulous con-

trol of the DNA adsorbed layer. Although it is very simple, this strat-

egy requires multisite attachment—such as adsorption—as immo-

bilization technique. Among the different kinds of electrochemical

transducers, it is expected that composites will have the greatest

impact of nanotechnology for improved electrochemical biosensors.

Next section is focused on the main features of composites as elec-

trochemical transducers.

3.2 Composites Materials: Main Features andClassification

When different materials are combined, the properties of the result-

ing composite material depend on the properties of the constituent

materials, the length scale, as well as chemical and morphological

details of the dispersion. Each individual component maintains its

original characteristics while giving the composite distinctive chem-

ical, mechanical, physical, or biological qualities [45]. These global

features are different and synergist from those shown by the indi-

vidual elements of the composite [18, 46].

The first classification of composite can be made in terms of the

nature, in biological and engineering composites. Two simple, bio-

logical polymer-based examples of composites are wood, made up

of fibrous chains of cellulose in a matrix of lignin and bone, com-

posed of hard inorganic crystals (hydroxyapatite) embedded in a

tough organic matrix (collagen). Composite materials that consist

of a matrix (metal, polymer, ceramic) with embedded reinforcement

(filament, whiskers, particles) comprise of many high performance

engineering materials.

A nanostructured composite or nanocomposite results when the

characteristic length scales of at least one of the components is in

the nanometer range. Nanometer-sized filler materials, with their

inherently large surface-area-volume ratios, are particularly inter-

esting as they facilitate increasing efficiency of a given property.

Moreover, due to the small size of the filler, certain properties may be

modified while not affecting others [47]. Being in the nanotechnol-



ogy era, novel nanostructured composite materials are expected to

be designed showing improved properties due to nanostructuration.

The composite materials can also be classified according to the

physical properties in soft or rigid composites.

A conducting composite results if at least one of the phases is an

electrical conductor. The overall electrical properties of the conduct-

ing composite will be determined by the nature, the relative quan-

tities, and the distribution of each phase. Recent developments in

the field of conducting composites applied to electrochemistry have

opened a new range of possibilities for the construction of electro-

chemical sensors and biosensors. The main features of these mate-

rials have been described elsewhere in detail [48, 49].

A polymer composite results if at least one of the components is a

polymeric matrix, which can be a conducting or nonconducting poly-mer. As such, a conducting polymer composite can be obtained with

a conducting polymer matrix, or, instead, by using a non-conductingpolymer matrix but a conducting filler (such as platinum, gold, car-

bon, CNT, metal NPs, etc.).

Conducting polymers are basically organic conjugated poly-

mers, and their unusual electrochemical characteristics (e.g., low

ionization potential, high electrical conductivity, and high electronic

affinity) are due to the conjugated π -electron backbones in their

chemical composition. This is the reason why these conducting poly-

mers are often called “synthetic metals.” Their organic chains with

single- and double-bonded sp2 hybridized atoms generate a wide

charge delocalization and therefore are responsible for the metal-

like semiconductive properties of conducting polymers [17]. The

electrical and optical properties of conducting polymers are simi-

lar to those of metals and inorganic semiconductors. Moreover, bio-

molecules can be immobilized onto conducting polymers without

any loss of activity. Their mechanical and electronic properties can

be properly tailored by chemical modeling and synthesis [50]. The

attractive feature of biosensor applications results essentially from

the rapid electron transfer that they provide in electrode surfaces as

a consequence of the biological event. Conducting polymers can be

synthesized either by chemical or electrochemical oxidation. Elec-

trochemical method is based on the oxidation of monomers leading

to the formation of cation radicals that repeatedly bind to the grow-


Composites Materials 63

ing polymer [51]. Polypyrrole and polyaniline are considered nowa-

days the most promising conducting polymers for the development

of biosensor devices owing to their good biocompatibility, conduc-

tivity, and stability. The combination of nanoengineered “smart” con-

ducting polymers with biomolecules and nanostructures, like metal

NPs and carbon nanotubes (CNTs), may generate conducting com-

posites with new and interesting properties, providing higher sensi-

tivity and stability of the immobilized biomolecules.

Nonconducting polymers are polymeric binders (epoxy,

methacrylate, silicone, araldite) which confer to the conducting com-

posite a certain physical, chemical, or biological stability, while the

electrical conductivity is provided by the conducting filler (micro or

nanoparticles of platinum, gold, graphite, carbon nanotubes, etc.).

Conducting composites based on nonconducting polymers are

classified by the nature of the conducting material and the arrange-

ment of its particles (i.e., whether the conducting particles are dis-

persed in the polymer matrix or if they are grouped randomly in

clearly defined conducting and insulating zones).

The inherent electrical properties of the conducting compositedepend on the nature of each of the components, their relative quan-

tities, and their distribution. Micro and nanostructurated conduct-

ing particles are usually used as fillers in conducting composites:

the electrical resistance is determined by the connectivity of these

conducting micro or nanoparticles inside the nonconducting matrix;

therefore, the relative amount of each component has to be assessed

to achieve optimal composition. A percolation curve [52], as shown

in Fig. 3.1, is a representation of the logarithmic variation of the elec-

trical resistance of a composite as a function of its conducting phase

content. By constructing a percolation curve, it is possible to deter-

mine the minimum conductor content required to achieve certain

conductivity. This point is known as the percolation threshold.

The composite acquires particular electrochemical features from

the nature of the conductive filler in the bulk.

The extensive range of unique properties inherent to metal

NPs, including electrical conduction, makes them very attractive

candidates for integration into polymers as NP-polymer compos-

ites. Embedding NPs into host polymers provides a means for

introducing a variety of properties to the polymer-based com-



0 100Volume fraction (%)

Conductive filler

f

Log

resi

stiv

ity

12

3

Figure 3.1. Percolation curve of a conducting composite based on noncon-

ducting polymeric binder with conductive filler. Theoretical dependence of

composite resistivity on conductive filler content. In the zone 1, the electri-

cal resistance of the composite is similar to that of the polymer. In zone 2,

the percolation fraction f represents a critical conductive filler content that

permits the formation of the first conducting filament consisting of particle-

to-particle contacts. In zone 3, electrical resistance of the composite is simi-

lar to that of pure conductive filler.

posite materials, including conductivity in the case of gold NPs,

magnetic properties when using cobalt or iron oxide NPs, and

mechanical properties using NP fillers such as clay [53]. Whatever

the particular NP’s composition or shape, its blending with most

polymers tends toward phase separation that results in particle clus-

tering or aggregation within the host polymer. This problematic

issue can be addressed very effectively by appropriate NP surface

modification. A number of studies involves gold NP–polymer com-

posites introducing thiol-terminated polymers at some stage of the

NP growth process [54] or subsequent to the initial NP synthesis

[55, 56]. Covering metal surfaces with electronically active polymers


Carbon Composites 65

provides a close proximity of the two materials that facilitate their

electronic communication.

Besides metal NPs, carbon micro and nanoparticles as well as

carbon nanotubes are usually used as fillers in composites for elec-

trochemical transducers. The following section focuses on the prop-

erties and main issues of carbon-polymer conducting composites

based on nonconducting binders and polymers.

3.3 Carbon Composites

3.3.1 Carbon-Based Materials as ConductiveFillers in Composites

Carbon is an ideal choice as composite filler due to its high chem-

ical inertness, wide range of working potentials, low electrical

resistance, and low residual currents. The extraordinary ability of

carbon to combine with itself and other chemical elements in dif-

ferent ways is the basis of organic chemistry. As a consequence,

there is a rich diversity of structural forms of solid carbon because

it can exist as any of the several allotropes. It is found abundantly

in nature as coal, a natural graphite, and also in much less abundant

form as diamond. Engineered carbons [57] are the product of the

carbonization process of a carbon-containing material, conducted

in an oxygen-free atmosphere. Depending on the starting precur-

sor material (hydrocarbon gases, petroleum-derived products, coals,

polymers, biomass), the product of a carbonization process will have

different properties, including the adsorption capability. Traditional

engineered carbons can take many forms, such as coke, graphite, car-

bon and graphite fiber, carbon monoliths, glassy carbon (GC), car-

bon black, carbon film, and diamond-like film [57]. The discovery

of nanostructured carbon-based materials added a new dimension

to the knowledge of carbon science. The first TEM evidence for the

tubular nature of some nanosized carbon filaments, that is, of carbon

nanotubes (MWCNT, multiwalled carbon nanotubes) was reported

in 1952 by Radushkevich and Lukyanovich [58]. The subsequent dis-

covery in the “nano” era of “fullereness” [59] has also impacted the

carbon science. Finally, the growing of SWCNTs was first reported in



1993 by two papers submitted independently, one by Iijima and Ichi-

hashi [60], and the other by Bethune et al. [61]. CNTs can be grown

by using the arc-discharge method, the laser vaporization method,

and the chemical vapor deposition (CVD) [62]. Compared with

arc-discharge and laser methods, CVD is a simple and economic

technique for synthesizing CNTs at low temperature and ambient

pressure, at the cost of crystallinity. It is versatile in that it harnesses

a variety of hydrocarbons in any state (solid, liquid, or gas), enables

the use of various substrates, and allows CNT growth in a variety

of forms, such as powder, thin or thick films, aligned or entangled,

straight or coiled, or even a desired architecture of nanotubes at pre-

defined sites on a patterned substrate. It also offers better control

over growth parameters.

Carbon-based materials have found intensive use as adsorbents

because of their porous and highly developed internal surface areas

as well as their complex chemical structures. The porous struc-

ture and the chemical nature of the carbon surface are signifi-

cantly related to its crystalline constitution. The crystal structure

of graphite consists of parallel layers of condensed, regular hexag-

onal rings. The in-plane C–––C distance is intermediate between the

Csp3–––Csp3 and the Csp2===Csp2 bond lengths. Graphene is the hypo-

thetical infinite aromatic sheet of sp2-bonded carbon that is the 2-D

counterpart of naturally occurring 3-D graphite. It is found in the

π -stacked hexagonal structure of graphite with an interlayer spac-

ing of 3.34 A, which is the van der Waals distance for sp2-bonded

carbon [63].

The pore structure and surface area of carbon-based materials

determine their physical characteristics, while the surface chemi-

cal structure affects interactions with polar and nonpolar molecules

due to the presence of chemically reactive functional groups. Active

sites—edges, dislocations, and discontinuities—determine the reac-

tivity of the carbon surface. Graphitic materials have at least two dis-

tinct types of surface sites, namely, the basal-plane and edge-plane

sites [64]. It is generally considered that the active sites for electro-

chemical reactions are associated with the edge-plane sites, while

the basal plane is mostly inactive. Heteroatoms (usually oxygen)

play an important role in the chemical nature of the carbon “active”

surface [57]. The adsorption process is thus strongly dependent on



the type, quantity, and bonding of these functional groups in the

structure. Heteroatoms distributed randomly in the core of the car-

bon matrix may be nonreactive due to their inaccessibility. How-

ever, the heteroatoms can also be concentrated at the exposed

surface of carbons or presented as an “active” dislocation of the

microcrystalline structure. Much of the research being carried out

is focused not only on the identification and characterization of

oxygen-containing functional groups in oxidized carbon surfaces,

such as carboxyl, phenolic, quinonic, and lactones, but also in the

changes that take place in the carbon surface under different oxida-

tion treatments.

The electrochemical oxidation pretreatment was found to

improve the electrochemical behavior by introducing more active

edge sites on the treated carbon surface. The effect of oxidation

on the chemical composition is related to the increased concentra-

tion of strong and weak acidic groups found upon electrochemi-

cal oxidation of the graphite surface [65]. The acidity of carboxylic

groups on the oxidized carbon surface could be stronger than that

of a carboxylic resin. The weight increase after electrochemical pre-

treatment was attributed to the formation of the oxidized graphite

and the intercalation of solvent molecules and anions into graphitic

material.

Regarding the CNTs, the growing methods provide not only

the CNT product, but also different contaminants (mainly amor-

phous carbon and catalyst metallic particles) which are commonly

removed by treatment with oxidizing acids, for example, HNO3,

which results in ends largely decorated with carboxyl groups [66].

However, defects in the sidewalls can also be introduced under

such drastic conditions. Functionalization or modification of CNTs

has become a major activity within the interdisciplinary fields of

nanoscience, nanotechnology, bioengineering, and bionanotechnol-

ogy, as it promises to be the best approach for improving the sol-

ubility and compatibility of CNTs. Defects in SWNTs are important

in the covalent chemistry of the tubes because they can serve as

anchor groups for further functionalization, and therefore a promis-

ing starting point for the development of the covalent chemistry of

SWNTs [67].



CNTs present a larger surface area and outstanding charge-

transport characteristics and might therefore greatly promote

electron-transfer reactions which can dramatically improve elec-

trochemical performance compared to that of other carbonaceous

materials [68]. The open end of a CNT is expected to show a fast

electron-transfer rate similar to the graphite edge plane while the

sidewall is inert like the graphite basal-plane. Fast electron-transfer

rate is demonstrated along the tube axis [69].

CNTs are expected to present a wide electrochemical window,

flexible surface chemistry, and biocompatibility, similar to other

widely used carbon materials.

Among the different classes of carbon allotropes, carbon-based

composites, such as carbon paste (CP), are usually made of polycrys-

talline graphite. A key property of polycrystalline graphite is poros-

ity. Most polycrystalline graphite—such as powdered carbon—is

made by heat treatment of high-molecular-weight petroleum frac-

tions at high temperatures to perform graphitization. The term

“graphite” is used to designate materials that have been subjected

to high temperatures, and thus have aligned the sp2 planes parallel

to each other.

Commercially available microcrystalline graphite exists as

extremely hydrophobic 1 to 20 μm particles that aggregate into

thin films on contact with solvent. When treated under strongly

oxidizing acidic conditions, graphite oxide is formed. Structurally,

graphite oxide is an epoxidized form of the sp2-bonded carbon net-

work together with acidic functional groups at the edges with the

oxidants intercalated in the interlaminar space [63].

While the electrical conductivity is provided by the conducting

carbonaceous filler, in order to prepare a carbon composite, a binder

is also needed. The binder will confer the conducting composite a

certain physical, chemical, or biological stability. One of the sim-

plest carbon composite approaches for electrochemical biosensor

is based on soft carbon pastes [70]. These pastes are built by mix-

ing an inert conductor (e.g., graphite powder) with a nonconducting

liquid (e.g., paraffin oil, silicone, Nujol). This insulating liquid has a

specific viscosity and the paste has a certain consistency. The result-

ing devices are easy to prepare and inexpensive. However, these

pastes have limited mechanical and physical stability, especially in



flow systems. Additionally, the pastes are dissolved by some nonpo-

lar electrolytic solvents, leading to a deterioration of the signal. The

general degradation of these devices occurs quickly and has limited

their use to the research laboratory [18]. Unlike soft carbon paste,

rigid carbon polymer composites allow the design of different con-

figurations, and these materials are compatible with nonaqueous

solvents. Next section is focused in the preparation and properties

of rigid carbon polymer composites.

3.3.2 Rigid Carbon-Polymer Composite

Rigid carbon-polymer composites are obtained by mixing a car-

bon filler (such as graphite or CNT) with nonconducting polymeric

binders (epoxy, methacrylate, silicone, araldite), obtaining a soft

paste that becomes rigid after a curing step.

Rigid carbon-polymer composites are interesting alternatives for

the construction of electrochemical (bio)sensors. The capability of

integrating various materials (including nanostructured particles

and biomolecules) is one of their main advantages. Some materi-

als which are incorporated within the composite result in enhanced

sensitivity and selectivity. The best composite components will give

the resulting material improved chemical, physical, and mechani-

cal properties. As such, it is possible to choose between different

binders and polymeric matrices and conductive fillers in order to

obtain a better signal-to-noise ratio, a lower non-specific adsorption,

and improved electrochemical properties (electron transfer rate and

electrocatalytic behavior). This incorporation is possible to be per-

formed either through a previous modification of one of the com-

ponent of the composite before its preparation or through physical

incorporation into the composite matrix.

The electrical resistance is determined by the connectivity of

the conducting particles inside the nonconducting matrix; there-

fore, the relative amount of each composite component has to

be assessed to achieve optimal composition. Figure 3.2 shows

scanning electron micrographs of different carbon-based materi-

als based on the same conductive filler (graphite) but different

polymeric binders: (A) Araldite-M–graphite (73.2%), (B) Araldite-

CW2215–graphite (45.8%), (C) silicone–graphite (61.0%), (D)



Figure 3.2. Scanning electron micrographs of mechanically polished

rigid conducting composite electrodes based on (A) Araldite-M–graphite

(73.2%), (B) Araldite-CW2215–graphite (45.8%), (C) silicone–graphite

(61.0%), (D) epoxy-H77-graphite (20.0%) (Adapted with permission from

Analyst 2002, 127, 1512–1519. Copyright 2002, The Royal Society of

Chemistry).



epoxy-H77–graphite (20%). The optimal composition in each case

was obtained from the percolation curve [52]. As shown in the scan-

ning electron micrographs, completely different materials have been

obtained with different polymeric binders. Moreover, the polymeric

binders accept different “optimal” amount of graphite.

Being in the nanotechnology era, novel nanostructured compos-

ite materials are expected to be designed showing improved proper-

ties due to nanostructuration. Moreover, composite electrodes from

expensive metals (gold, platinum, etc.) can be prepared using the

NPs as conductive fillers, with enhanced properties but at lower

prices compared to their pure conductor counterparts.

Rigid carbon-polymer composite electrodes offer many potential

advantages compared to more traditional electrodes consisting of

a surface-modified continuous conducting material. Rigid carbon-

polymer composite electrodes can often be fabricated with great

flexibility in size and shape, allowing the construction of different

electrode configurations. Rigid composite surfaces can be smoothed

or polished to provide fresh active material ready to be used in a

new assay. Each new surface yields reproducible results because

all individual compounds are homogeneously dispersed in the bulk

of the composite. Moreover, rigid carbon-polymer composites show

improved electrochemical performances, similar to an array of car-

bon fibers separated by an insulating matrix and connected in paral-

lel. The signal produced by this macroelectrode formed by a carbon

fiber ensemble is the sum of the signals of the individual microelec-

trodes. Composite electrodes thus showed a higher signal-to-noise

(S/N) ratio than the corresponding pure conductors, accompanied

by an improved electrochemical behavior [46].

3.3.3 Graphite-Epoxy Composites

Rigid conducting graphite-epoxy composites (GEC) [25, 35, 36] and

biocomposites (GEB) [18, 23] have been extensively used in our

laboratories for electrochemical (bio)sensing due to their unique

physical and electrochemical properties. In particular, we have used

GEC (graphite-epoxy composite) made by mixing the nonconduct-

ing epoxy resin (Epo-Tek, Epoxy Technology, Billerica, MA, USA) with

graphite microparticles (particle size below 50 μm).



This paste can be easily prepared by mixing graphite powder

with epoxy resin in a 1:4 (w/w) ratio. The soft paste is thoroughly

hand mixed to ensure the uniform dispersion of the graphite pow-

der throughout the polymer. The moldable soft paste is put on the

body of the electrodes and cured at 100◦C for 2 days to obtain

the rigid graphite-epoxy composite (GEC) [25]. A magneto graphite-

epoxy composite (m-GEC) electrode is prepared in the same way as

the GEC transducer, but in this case, a small magnet (3 mm i.d.) is

placed in the center of this electrode after the addition of a thin layer

of GEC paste in order to avoid the direct contact between the mag-

net and the electrical connector. After filling the electrode body gap

completely with the soft paste, the electrode is tightly packed and

then cured at the same temperature. This magneto electrode can be

easily coupled with magnetic particles [71, 72].

Biocomposites can also be easily prepared by adding the biore-

ceptor (an enzyme [18] and antibody [23], or an affinity receptor

such as Protein A [73] or avidin [74, 75]).

As an example, in the case of avidin graphite-epoxy biocomposite

(Av-GEB), graphite powder and epoxy resin are also hand mixed in

a ratio of 1:4 (w/w). In this case, for every gram of graphite/epoxy

mixture, an additional 20 mg of avidin is added—resulting in a 2%

(w/w) avidin-graphite-epoxy biocomposite. This mixture is thor-

oughly hand mixed to ensure the uniform dispersion of the avidin

and carbon throughout the polymer. The moldable soft paste is put

on the body of the electrodes and cured at 40◦C for 1 week to obtain

the rigid avidin-graphite-epoxy biocomposite (Av-GEB).

Finally, gold nanocomposites are prepared by hand-mixing the

following ratios of gold-NPs, graphite powder, and epoxy resin:

0.075/0.925/4 (w/w) for nanoAu(7.5%)-GEC. The resulting soft

paste is placed in the gap of electrode and cured at 80◦C for 1 week to

obtain the rigid gold NPs graphite-epoxy composite (nanoAu-GEC).

The GEC-based transducers present numerous advantages over

more traditional carbon-based materials: higher sensitivity, robust-

ness, and rigidity. Additionally, the surface of GEC can be regenerated

by a simple polishing procedure.

An ideal material for electrochemical genosensing should allow

an effective immobilization of the probe on its surface, a robust

hybridization of the target with the probe, a negligible non-specific


Electrochemical Genosensing Based on Graphite-Epoxy Composite 73

adsorption of the label and a sensitive detection of the hybridiza-

tion event. Graphite-epoxy composites (GEC) fulfill all these require-

ments.

In the present work, graphite-epoxy composite, biocomposite,

and nanocomposite materials for the development of electrochem-

ical genosensors are reviewed. Different graphite-epoxy platforms

for electrochemical genosensing, as well as strategies for detect-

ing DNA hybridization are presented. The advantages of these new

graphite-epoxy platforms for electrochemical genosensing are dis-

cussed and compared with the current state of the art in DNA sens-

ing techniques.

3.4 Electrochemical Genosensing Based onGraphite-Epoxy Composite

3.4.1 Electrochemical Genosensing Based on DNADry Adsorption on GEC as ElectrochemicalTransducer

Adsorption is an easy way to attach nucleic acids to solid surfaces,

since no reagents or modified-DNA are required, as shown in Fig. 3.3.

These features have promoted extensive use of adsorption as immo-

bilization methodology in genetic analysis. The mainly claimed dis-

advantages of adsorption with respect to covalent immobilization

are (i) nucleic acids may be readily desorbed from the substrate and

(ii) base moieties may be unavailable for hybridization if they are

bonded to the substrate in multiple sites [76]. However, the electro-

chemical detection strategy based on the intrinsic oxidation of DNA

requires the DNA to be adsorbed in close contact with the electro-

chemical substrate by multisite attachment, as schematically shown

in Fig. 3.4. This multisite attachment of DNA can be thus detrimen-

tal for its hybridization but is crucial for the detection based on its

oxidation signals. The common method for the multisite physical

adsorption of DNA on carbonaceous-based materials can be classi-

fied into dry or wet adsorptions.

Dry adsorption relies on leaving DNA to dry on the carbonaceous

surface. It can be assisted by light treatment (except UV that is able



Figure 3.3. Different strategies for ssDNA probe immobilization in

genosensing devices based on GEC composites, biocomposites, and

nanocomposites. (A) Dry or wet multisite adsorption on GEC; (B) avidin-

biotin linkage on Av-GEB; (C) (strept)avidin-biotin linkage on magnetic

beads captured on m-GEC; (D) chemisorption on nanoAu-GEC. See also

Color Insert.

to induce changes in DNA molecule) or heated until 100◦C. DNA can

adopt a variety of conformations depending on the degree of hydra-

tion. As an example, the most familiar double helix DNA—called

“B-DNA”—can become into the “A-DNA” form if it is strongly dehy-

drated. A structural alteration occurs due to a greater electrostatic

interaction between the phosphate groups, leading to A-DNA. When

the DNA solution is evaporated to dryness, the bases of DNA which

have been dehydrated are exposed, thus the hydrophobic bases are

strongly adsorbed flat on the electrode surfaces. Once it is adsorbed,

DNA is difficult to re-hydrate. Hence, DNA is not desorbed, no mat-

ter how long the adsorbed-DNA is soaked in water, characteristic of

irreversible adsorption.

The “irreversible” behavior of the dry adsorbed DNA layer has

been previously reported on glassy carbon electrodes [77]. DNA can

be tightly and irreversibly immobilized on GEC by both dry and wet

adsorption procedures under static conditions [78]. The dual nature



Figure 3.4. Different detection strategies in electrochemical genosensing

based on GEC transducers. See also Color Insert.

of GEC composed of islands of conducting material within the non-

conducting and hydrophobic epoxy resin could play an important

role in stabilizing the dehydrated A-form of DNA adsorbed on GEC.

Once immobilized on GEC, DNA preserves its unique hybridization

properties, which can be revealed using different strategies based

on both enzymatic labeling and the intrinsic signal coming from the

DNA oxidation, as schematically shown in Fig. 3.4.

The DNA immobilization on GEC surface by dry adsorption was

performed by covering the GEC surface with a small drop of DNA

in 10 X SSC, and allowing the electrode to dry at 80◦C for 45 min

in upright position [25]. The DNA electrochemical detection was

then achieved by an enzymatic labeling step [36]. Briefly, the pro-

cedure consists of the following steps: (i) DNA target immobiliza-

tion; (ii) hybridization with the complementary probe modified

with either biotin or digoxigenin; (iii) enzyme labeling of the DNA

duplex using streptavidin-HRP or anti-DIG-HRP; and (iv) ampero-

metric determination based on the enzyme activity by adding H2O2



and using hydroquinone as a mediator. If the PCR product—or any

other double-stranded DNA—is directly adsorbed on GEC trans-

ducer, a denaturing alkaline procedure after the DNA dry adsorption

is mandatory to break the hydrogen bonds linking the comple-

mentary DNA strands in order to ensure proper hybridization

[79].

Although a compact thick ssDNA layer can be achieved by

dry adsorption, DNA preserves its unique hybridization properties,

which can be monitored using different strategies, suggesting that

the DNA bases are not fully committed in the adsorption mechanism.

DNA bases are mostly available for hybridization, taking into account

the differences in signal compared with the non-specific adsorption

[79, 80]. This strategy was able to electrochemically detect the PCR

amplicon coming from Salmonella spp. in a very simple and cheap

way [79].

Besides this strategy in which the DNA target can be easily

attached and detected by its complementary DNA signaling probe,

a sandwich assay in which the DNA target is in solution can be

easily performed by a double hybridization with a capture and a

signaling probe [25]. This strategy was demonstrated to be use-

ful for the detection of a novel determinant of β-lactamase resis-

tance in S. aureus using one- and two-step capture format. Accord-

ing to the results, the one-step capture format is more convenient,

as a higher sensitivity was achieved [25]. When compared with

other reported genosensor designs using a similar capture format

[81] and the same labeling system, the genosensor design based on

dry adsorption is simpler and cheaper, showing detection limits of

the same order of magnitude. The procedures based on previous

activation/modification of the surface transducer and subsequent

immobilization, as well as some blocking and washing steps that are

tedious, expensive, and time-consuming, were avoided using GEC as

electrochemical platform. These are the principal advantages of GEC

platform with respect to other reported devices [37, 81, 82].

Moreover, by easily controlling the concentration of the DNA

solution being dried on GEC platform, a thick or a thin layer of DNA

can be formed on the GEC surface by dry adsorption [36]. Depend-

ing on the application of the DNA-modified substrate, a thick or thin

DNA layer would be necessary. If a stringency control of non-specific



DNA adsorption issues is required, a thick DNA layer is more con-

venient. However, the yield in hybridization is better on a thin DNA

layer [36].

Furthermore, GEC has shown unique and selective adsorption

behavior. While DNA is firmly adsorbed under dry conditions, the

wet adsorption of non-specific DNA, proteins, enzymes, or other

biomolecules is negligible under stirring or convection conditions

in solution [25, 79, 80]. The DNA-modified GEC surface does not

require blocking steps to minimize the non-specific adsorption on

the free sites of the surface [36] since the non-specific adsorp-

tion is very low and similar to the instrumental background noise.

Moreover, no blocking reagents are required during hybridization

to reduce the non-specific adsorption. It was previously demon-

strated that the hybridization signals (as well as the non-specific

adsorption signals) were essentially the same when performing the

hybridization without blocking reagents and using different block-

ing commercial solution, such as (i) 5XSSC, 1XDenhardt’s, 100 μg/ml

chloroform extracted salmon testes DNA, 0.5% (w/v), SDS and 50%

(v/v) formamide; (ii) 5XSSC, 1XDenhardt’s, 0.5% (w/v) SDS and

50% (v/v) formamide; and (iii) 5XSSC, 0.5% (w/v) SDS and 50%

(v/v) formamide. Comparable hybridization signals as well as non-

specific adsorption are achieved by using those three different solu-

tions for hybridization [25].

3.4.2 Electrochemical Genosensing Based on DNAWet Adsorption on GEC as ElectrochemicalTransducer

Wet adsorption relies on leaving DNA to interact with the carbona-

ceous surface through physical forces in the presence of water. Dur-

ing wet adsorption, the stabilization of B-DNA is expected to occur

on the carbonaceous surface, by keeping the hydration water of the

DNA molecule. In this case, the hydrated B-DNA form is stabilized

over the GEC surface by weaker forces: as the water is kept on the

DNA adsorbed molecule, it can be easily desorbed from the GEC sur-

face if soaked in aqueous solutions.

DNA can be easily immobilized on GEC by simple wet adsorption

onto GEC surface. A small drop of DNA probe in acetate saline



solution pH 4.8 [25] is put onto the surface of a GEC electrode in

an upright position. The immobilization of the probe was allowed

to proceed for 15 min without applying any potential under static

conditions.

After the inosine-modified DNA probe immobilization, the DNA

target was detected by the intrinsic DNA oxidation signal coming

from the guanine moieties, as schematically outlined in Fig. 3.4.

Briefly, the procedure consists of the following four steps: [83] (i)

electrochemical pre-treatment of the GEC transducer; (ii) inosine-

substituted probe immobilization by wet adsorption on GEC trans-

ducer; (iii) hybridization with the target; and (iv) electrochemical

determination based on differential pulse voltammetry (DPV), in

which the oxidation signal of guanine (or adenine) was measured

by scanning from +0.30 to +1.20V at a pulse amplitude of 100 mV

and a scan rate of 15 mV/s. This procedure was demonstrated to be

useful for the detection of IS200 element specific for Salmonella spp.

[83].

Although a thick or a thin layer of DNA can be attached on the

surface during dry adsorption by controlling the concentration of

the DNA solution being dried, the wet adsorption normally yields

a thin DNA layer. Less compact DNA layers with wider gaps exposing

free-GEC surface are normally obtained during wet adsorption. As

a consequence, the thin-layer DNA/GEC surface required blocking

treatment to avoid non-specific adsorption. During wet adsorption,

the substrate is progressively modified with negative charges com-

ing from the DNA being adsorbed; thus, rejecting the successive DNA

molecules that are approaching the substrate. Wet adsorption thus

leads to a “self-control” surface coverage and is less stringent than

dry adsorption [25].

3.4.3 Electrochemical Genosensing Based onGraphite-Epoxy Biocomposite Modified with Avidin(Av-GEB) as Electrochemical Transducer

A rigid and renewable transducing material for electrochemical

biosensing, based on avidin bulk-modified graphite–epoxy biocom-

posite (Av-GEB), can be easily prepared by adding a 2% avidin

(or streptavidin) in the formulation of the composite and using



dry chemistry techniques, avoiding tedious, expensive, and time-

consuming immobilization procedures. The rigid conducting bio-

composite acts not only as a transducer but also as a reservoir for

avidin. After use, the electrode surface can be renewed by a simple

polishing procedure for further uses, highlighting a clear advantage

of this new material with respect to surface-modified approaches

such as classical biosensors and other common biological assays.

DNA probe can be easily immobilized on the surface of the avidin-

modified transducer through the avidin–biotin reaction, since both

nucleic acids as well as short oligonucleotides can be readily linked

to biotin without serious effects on their biological, chemical, or

physical properties (Fig. 3.3B). The knowledge about the avidin–

biotin interaction has advanced significantly and offers an extremely

versatile tool. Moreover, this interaction presents a variety of specific

advantages over other single-point immobilization techniques. In

particular, the extremely specific and high affinity reaction between

biotin and the glycoprotein avidin (association constant Ka = 1015)

leads to strong associations similar to the formation of a cova-

lent bonding. This interaction is highly resistant to a wide range of

chemicals (detergents, protein denaturants), pH range variations,

and high temperatures [84]. In addition, much progress has been

done in the modification of biomolecules with biotin. Moreover, the

strept(avidin) could be considered as a universal affinity biomole-

cule because it is able to link not only biotinylated DNA or ODNs but

also enzymes or antibodies [23].

Biotinylated DNA can be firmly single-point attached in Av-

GEB (Fig. 3.3B). In this case, a capture format was used in which

the immobilization of the biotinylated probe together with the

hybridization was performed in a one-step procedure [74]. Briefly,

the three-step experimental procedure consists of (i) one-step

immobilization/hybridization procedure in which the biotin-labeled

capture probe is immobilized onto the electrode surface through

a biotin–avidin interaction, while the hybridization with the target

and with a second complementary probe—in this case labeled with

digoxigenin—is occurring at the same time; (ii) enzymatic labeling

using as enzyme label the antibody anti-DIG-HRP; and (iii) ampero-

metric determination based on the enzyme activity by adding H2O2

and using hydroquinone as a mediator.



The utility of Av-GEB platform was demonstrated for the deter-

mination of the mecA DNA sequence related with methicillin-

resistant S. aureus (MRSA) [85] in a simpler and specific manner

with respect to previous DNA biosensing devices [25].

The genosensor design based on Av-GEB not only is able to

successfully immobilize onto the electrode surface with the mecA

biotin-labeled capture probe, while the hybridization with the mecA

target and the mecA digoxigenin-labeled probe is occurring at the

same time, but is also capable of distinguishing SNPs.

Compared to genosensors based on GEC, the novelty of this

approach is in part attributed to the simplicity of its design, com-

bining the hybridization and the immobilization of DNA in one ana-

lytical step.

The optimum time for the one-step immobilization/

hybridization procedure was found to be 60 min [74]. The pro-

posed DNA biosensor design has proven to be successful in using

a simple bulk modification step; hence, overcoming the complicated

pre-treatment steps associated with other DNA biosensor designs.

Additionally, the use of a one-step immobilization and hybridiza-

tion procedure reduces the experimental time. Stability studies con-

ducted demonstrate the capability of the same electrode to be used

for a 12-week period [74].

The rapid electrochemical verification of the amplicon com-

ing from the Escherichia coli O157:H7 genome was performed by

double-labeling the amplicon during PCR with a set of two labeled

PCR primers—one of them with biotin and the other one with digox-

igenin. During PCR, not only the amplification of the E. coli was

achieved but also the double-labeling of the amplicon ends with (i)

the biotinylated capture primer to achieve the immobilization on a

biosensor based on a bulk-modified avidin biocomposite (Av-GEB)

and (ii) the digoxigenin signaling primer to achieve the electrochem-

ical detection. The procedure consisted briefly of the following steps:

(i) DNA amplification and double-labeling of the eaeA gene, related

with the pathogenic activity of Escherichia coli O157:H7; (ii) immobi-

lization of the doubly labeled amplicon in which the biotin extreme

of the dsDNA amplicon was immobilized on the Av-GEB biosensor;

(iii) enzymatic labeling with anti-DIG-HRP capable of bonding with

the other labeled extreme of the dsDNA amplicon; and (iv) ampero-

metric determination [75].



3.4.4 Electrochemical Genosensing Based on MagneticBeads and m-GEC Electrochemical Transducer

One of the most promising materials, which have been developed, is

biologically modified magnetic beads [86] based on the concept of

magnetic bioseparations. Magnetic beads offer some new attractive

possibilities in biomedicine and bioanalysis since their size is com-

parable to those of cells. Moreover, they can be coated with biological

molecules and they can also be manipulated by an external mag-

netic field gradient. As such, the biomaterial, specific cells, proteins,

or DNA, can be selectively bound to the magnetic beads and then

separated from its biological matrix by using an external magnetic

field. Moreover, magnetic beads of a variety of materials and sizes,

and modified with a wide variety of surface functional groups, are

now commercially available. They have brought novel capabilities

to electrochemical immunosensing [87–89]. The magnetic beads

have also been used in novel electrochemical genosensing protocols

[90]. These approaches using magnetic beads for detection of DNA

hybridization have been combined with different strategies for the

electrochemical detection, such as label-free genosensing [91–93]

or different external labels, such as enzymes [90], electrochemical

indicators [94], or metal tags, for example, gold or silver NPs [95],

and using different electrochemical techniques, such as DPV, poten-

tiometric stripping analysis (PSA), or square wave voltammetry

(SWV).

Instead of the direct modification of the electrode surface, the

biological reactions (as immobilization, hybridization, enzymatic

labeling, or affinity reactions) and the washing steps can be suc-

cessfully performed on magnetic beads. After the modifications, the

magnetic beads can be easily captured by magnetic forces onto the

surface of GEC electrodes holding a small magnet inside (m-GEC).

Once immobilized on m-GEC, the hybridization performed on the

magnetic beads can be electrochemically revealed using different

strategies based on both enzymatic labeling and the intrinsic sig-

nal coming from the DNA oxidation [96] (as shown in Fig. 3.4).

In the case of using magnetic beads, a single-point immobilization

procedure based on streptavidin–biotin interaction is performed

(Fig. 3.3C). Biotinylated DNA can be firmly attached on streptavidin-

modified magnetic beads in that way.



When the electrochemical detection is based on enzymatic

activity determination, a capture format was used in which

the immobilization of the biotinylated probe together with the

hybridization was performed in a one-step procedure. The proce-

dure consists briefly of the following steps, as schematically outlined

in Fig. 3.5: (i) one-step immobilization/hybridization procedure

Figure 3.5. Schematic representation of the electrochemical strategy for

the detection of Salmonella spp. (A1) One-step procedure based on immo-

bilization of the biotinylated probe onto magnetic beads and hybridization

with the ssDNA target; (A2) rapid verification of PCR amplification based

on the doubly labeled PCR product detection; and (A3) real-time PCR reac-

tor based on PCR amplification with magnetic primers on streptavidin-

modified magnetic beads. Enzymatic labeling (B), magnetic capture of

the modified magnetic beads by the magneto electrode (m-GEC) (C), and

chronoamperometric determination (D) are common steps for all of these

strategies (A–C). (Reprinted with permission from Biosens. Bioelectron. 22,

2010–2017 (2007). Copyright 2006, Elsevier B.V.). See also Color Insert.



in which the biotin-labeled capture probe is immobilized on the

streptavidin magnetic beads, while the hybridization with the tar-

get and with a second complementary probe—in this case labeled

with digoxigenin—is occurring at the same time (30 min at 42◦C)

(Fig. 3.5, A1); (ii) enzymatic labeling using as enzyme label the

antibody anti-DIG-HRP (Fig. 3.5, B); (iii) magnetic capture of the

modified magnetic particles (Fig. 3.5, C); and (iv) amperometric

determination based on the enzyme activity by adding H2O2 and

using hydroquinone as a mediator (Fig. 3.5, D) [97]. When the

DNA target immobilized on the magnetic beads was detected by

the intrinsic DNA oxidation signal coming from the guanine moi-

eties, the procedure consists of the following steps, as previously

described in detail [96]: (i) electrochemical pre-treatment of the m-

GEC transducer; (ii) biotinylated inosine-substituted probe immobi-

lization on streptavidin magnetic beads; (iii) hybridization with the

target; (iv) magnetic capture of the modified magnetic particles, fol-

lowed by dry adsorption, was performed during 45 min at 80◦C; and

(v) electrochemical determination based on DPV.

The procedure for electrochemical DNA biosensing based on

magnetic beads was also used for the detection of IS200 element

specific for Salmonella spp.

This new electrochemical genomagnetic strategy using magneto

electrodes in connection with magnetic particles offers many poten-

tial advantages compared to more traditional strategies for detecting

DNA.

This new strategy takes advantages of working with magnetic

particles, such as improved and more effective biological reactions,

washing steps, and magnetic separation after each step. This elec-

trochemical genomagnetic assay provides much sensitive, rapid, and

cheaper detection than other assays previously reported. This sen-

sitivity of the GEC with respect to other electrochemical transducer

and selectivity conferred by the magnetic separation were also used

for the detection of PCR amplicons coming from real samples.

The rapid electrochemical verification of the amplicon com-

ing from the Salmonella IS200 element [97] as well as the eaeAgene, related with Escherichia coli O157:H7 [75] was performed

by double-labeling the amplicon during PCR with a set of two

labeled PCR primers—one of them with biotin and the other one



with digoxigenin. During PCR, not only the amplification of the

bacteria genome is achieved but also the double-labeling of the

amplicon ends with (i) the biotinylated capture primer to achieve

the immobilization on the streptavidin-modified magnetic bead,

and (ii) the digoxigenin signaling primer to achieve the electro-

chemical detection. Besides this double-labeling PCR strategy, a

single-labeling PCR strategy with a further confirmation of the

amplicon by its hybridization was achieved by performing the PCR

with the biotin primer and a further hybridization step with a digoxi-

genin probe. The procedure consists briefly of the following steps, as

schematically outlined in Fig. 3.5: (i) DNA amplification and double-

labeling of Salmonella IS200 insertion sequence; (ii) immobiliza-

tion of the doubly labeled amplicon in which the biotin extreme of

the dsDNA amplicon was immobilized on the streptavidin magnetic

beads (Fig. 3.5, A2); (iii) enzymatic labeling using as enzyme label

the antibody anti-DIG-HRP capable of bonding the other labeled

extreme of the dsDNA amplicon (Fig. 3.5, B); (iv) magnetic capture of

the modified magnetic particles (Fig. 3.5, C); and (v) amperometric

determination (Fig. 3.5, D) [97].

Moreover, a PCR reactor for real-time electrochemical detec-

tion was also designed (Fig. 3.5, A3). In this case, the amplifica-

tion and double-labeling is performed directly on the streptavidin

magnetic beads by using “magnetic bead primers” [97]. The proce-

dure consists briefly of the following steps: (i) in situ DNA amplifi-

cation and double-labeling of Salmonella IS200 insertion sequence

on streptavidin-modified magnetic beads by using a magnetic bead

primer (Fig. 3.5, A3); (ii) enzymatic labeling using as enzyme label

the antibody anti-DIG-HRP capable of bonding the other labeled

extreme of the dsDNA amplicon (Fig. 3.5, B); (iii) magnetic capture of

the modified magnetic particles (Fig. 3.5, C); and (iv) amperometric

determination (Fig. 3.5, D).

The rapid and sensitive verification of the PCR amplicon related

with Salmonella can be achieved with 2.8 fmol of amplified product

[97]. In the case of E. coli the assay showed to be very sensitive, being

able to detect 0.45 ng μl−1 of the original bacterial genome after only

10 cycles of PCR amplification [75]. Moreover, the electrochemical

strategies for the detection of the amplicon showed to be more sen-

sitive compared with Q-PCR strategies based on fluorescent labels

such as TaqMan probes.



This strategy can be used for the electrochemical real-time quan-

tification of amplicon since a linear relationship with the amount of

amplified product was obtained [75]. Moreover, this strategy is use-

ful only when a unique and specific band is observed by gel elec-

trophoresis, because of the high specificity of the set of primer being

used in the PCR for the amplification of the bacteria genome. On

the contrary, if the set of primers amplifies not only the sequence of

interest but also other non-specific fragments, it is necessary to con-

firm the internal sequence of the amplicon by a second hybridiza-

tion with a digoxigenin signaling probe. In this case, a single

labeling with biotin during PCR was performed followed by a further

selective hybridization with a digoxigenin signaling probe. More-

over, magnetic bead primers were used for in situ amplification on

magnetic beads of the Salmonella genome and for further electro-

chemical detection of the amplified product. The DNA amplification

on magnetic beads by using the magnetic bead primer with electro-

chemical detection of the amplified product demonstrated to be an

alternative strategy to the classic detection systems. This strategy

was also easily adapted to an immunoseparation step of the bacteria

to improve the LOD for detecting pathogenic bacteria [98].

The procedure consisted briefly of the following steps, as

schematically outlined in Fig. 3.6: (i) immunomagnetic separation of

the bacteria from food samples (Fig. 3.6A); (ii) lysis of the bacteria

and DNA separation (Fig. 3.6B); (iii) DNA amplification and double-

labeling of Salmonella IS200 insertion sequence (Fig. 3.6C); (iv)

immobilization of the doubly labeled amplicon in which the biotin

extreme of the dsDNA amplicon was immobilized on the strepta-

vidin magnetic beads (Fig. 3.6D); (v) enzymatic labeling using as

enzyme label the antibody anti-DIG-HRP capable of bonding the

other labeled extreme of the dsDNA amplicon; (vi) magnetic capture

of the modified magnetic particles; and (vii) amperometric determi-

nation [98].

In this approach, the bacteria are captured and preconcentrated

from food samples with magnetic beads by immunological reac-

tion with the specific antibody against Salmonella. After the lysis

of the captured bacteria, further amplification of the genetic mate-

rial by PCR with a double-tagging set of primers is performed to

confirm the identity of the bacteria. Both steps are rapid alter-

natives to the time-consuming classical selective enrichment and



Figure 3.6. Schematic representation of the IMS/double-tagging PCR/m-

GEC electrochemical genosensing approach. (Reprinted with permission

from Anal. Chem. 81, 5812–5820 (2009). Copyright 2009, American Chemi-

cal Society.) See also Color Insert.

biochemical/serological tests. The double-tagged amplicon is then

detected by electrochemical magneto genosensing using m-GEC

electrodes. The “IMS/double-tagging PCR/m-GEC electrochemical

genosensing” approach was used for the first time for the sensitive

detection of Salmonella artificially inoculated into skim milk sam-

ples. A limit of detection of 1 CFU mL−1 was obtained in 3.5 h with-

out any pretreatment, in LB broth and in milk diluted 1/10 in LB.

When the skim milk was pre-enriched for 6 h, the method was able

to feasibly detect as low as 0.04 CFU mL−1 (1 CFU in 25 g of milk)

with a signal-to-background ratio of 20 [98].



3.4.5 Electrochemical Genosensing Based onGraphite-Epoxy Composite Modified with GoldNanoparticles (NanoAu-GEC) as ElectrochemicalTransducer

Chemisorption based on self-assembled monolayers (SAMs) is

a single-point immobilization strategy that allows the oriented

attachment of a wide range of biomolecules on gold-based trans-

ducer surfaces. One of the main drawbacks of using SAMs for the

immobilization of biorreceptors in electrochemical biosensing is

that a compact layer is achieved which produces a dramatically

reduction of the diffusion of electroactive species toward the sur-

face of the transducer. Moreover, the tightly packed layer may also

produce steric hindrance and, as a consequence, lower rate of reac-

tion between the probe and the target. As such, a stringent control

of the surface coverage of the gold-based transducer is an important

factor, which can be performed by using auxiliary reagents such as

lateral spacer thiols and mixed monolayers to obtain bioactive gaps.

In order to avoid the stringent control of surface coverage para-

meters during immobilization of thiolated oligos on continuous

gold surface films, the use of gold NPs in a graphite-epoxy com-

posite (nano-AuGEC) has been proposed (Fig. 3.3D) [99]. In this

novel transductor, islands of chemisorbing material (gold NPs) sur-

rounded by rigid, non-chemisorbing, conducting, graphite-epoxy

composite are obtained, as shown in Fig. 3.7.

With this arrangement in the electrochemical transducer, the

resulting less-packed surface provides improved hybridization fea-

tures with a complementary probe minimizing steric and electro-

static repulsion. The spatial resolution of the immobilized thiolated

DNA was easily controlled by merely varying the percentage of gold

NPs in the composition of the composite (Fig. 3.8).

For GEC electrodes, graphite powder (particle size below 50 μm)

and epoxy resin (Epo-Tek, Epoxy Technology, Billerica, MA, USA) in

a 1:4 (w/w) ratio were thoroughly hand mixed to ensure the uni-

form dispersion of the graphite powder throughout the polymer. For

nanoAu-GEC electrodes, the following ratios of gold NPs/graphite

powder/epoxy resin were prepared: 0.075:0.925:4 (w/w) for

nanoAu(7.5%)-GEC; 0.250:0.750:4 (w/w) for nanoAu(25%)-GEC;



Figure 3.7. Schematic representation of (A) nanoAu-GEC material show-

ing isolated gold nanoparticles able to produce “bioactive chemisorbing

islands” instead of SAMs on a continuous layer of gold. (B1) Hybridization

assay on the nanoAu-GEC electrode. (B2) Rapid electrochemical verification

of thiolated and double-tagged amplicons on the nanoAu-GEC electrode.

Parts C to E are common steps (electrode modification, enzymatic labeling,

and amperometric determination) for both parts B1 and B2 (Reprinted with

permission from Anal. Chem. 2009, 81,1332–1339. Copyright 2009, Ameri-

can Chemical Society.) See also Color Insert.

0.500:0.500:4 (w/w) for nanoAu(50%)-GEC, and finally, 1:0:4

(w/w) for nanoAu(100%)-EC. The soft paste became rigid after a

curing step of 80◦C during 1 week.

Thereby, the designation of the electrodes is based on the ratio of

gold NPs toward graphite particles (the conductive filler).

The location and spatial pattern of the gold NPs on the surface

of the sensor was observed with scanning electron microscopy with

an EDX detector. Figure 3.8, first column, shows, as bright spots,



Figure 3.8. Microscopic characterization of nanoAu-GEC electrodes.

Microphotographs showing the distribution of gold nanoparticles on the

surface of nanoAu-GEC electrodes while increasing the amount of gold

nanoparticles from 0 to 100% of the conductive phase. First column, low-

resolution (100 μm) SEM with an EDX detector to identify gold element.

Acceleration voltage, 20 kV. Second column, fluorescence stereomicroscopy

at low resolution showing the fluorescence pattern of the different nanoAu-

GEC electrodes after the immobilization of 200 pmol of double tagged oligo

with thiol and fluorescein. Third column, stereomicroscopy without the flu-

orescence filter (Fig. 3.8, third column). See also Color Insert.



Figure 3.9. High-resolution (1−μm) SEM microphotographs showing the

isolated gold nanoparticles on the surface of nanoAu-GEC sensors, acceler-

ation voltage, 20 kV.

the aggregates of gold NPs for nanoAu(7.5%)-GEC, which appear to

be in increased frequency when increasing the percentage of gold

NPs until nanoAu(100%)-EC. However, high-resolution SEM micro-

graphs for nanoAu(7.5%)-GEC (Fig. 3.9) show clearly isolated gold

NPs of about 100 nm within the composite, demonstrated with

the EDX detector providing the characteristic gold x-ray spectrum.

Moreover, the availability of gold NPs in the composite for the immo-

bilization of thiolated oligos was also studied with fluorescence

stereomicroscopy. In this case, 200 pmol of double-tagged oligo with

both a thiolated 5′ end and the fluorescein 3′ end was immobi-

lized on the electrodes with different composition. As can be seen

in Fig. 3.8, an increasing amount of fluorescence was obtained with a

higher amount of gold NPs in the composite. The fluorescence shows

a discontinuous pattern as fluorescence dots of chemisorbing mate-

rial surrounded by nonreactive graphite-epoxy composite, except in

the case of nanoAu(100%)-EC, in which a continuous fluorescence

pattern is clearly observed. Moreover, it should be pointed out that

the fluorescence can be related with the isolated gold NP pattern

because it is not located in the aggregate zones, when comparing

with the same photos taken with the stereomicroscope without the

fluorescence filter (Fig. 3.8, third column). Thereby, the nanome-

ter scale of gold NPs seems also to play a role in the chemisorbing



ability of the gold nanocomposite material, especially in nanoAu

(7.5%)-GEC, since the fluorescence pattern is related to the isolated

gold NPs instead of the gold aggregates when increasing the amount

of goldNPs from 7.5 to 100%.

From the electrochemical evaluation of the electrodes, it can be

concluded that gold is more available for the electrochemical oxi-

dation in the nanoAu(7.5%)-GEC electrode and is present mostly as

NPs, instead of aggregates, showing the characteristic anodic peak

current near +1.1 V (vs. Ag/AgCl) [100]. The voltammetry for a vari-

ety of redox molecules at DNA-modified electrodes can provide addi-

tional qualitative information about the system’s organization on

the surface. The voltammetric reversibility of highly charged redox

ions is markedly influenced by the attractive/repulsive interactions

with the polyanionic DNA layer that the ions must penetrate to

reach the electrode surface, in the case of highly packed DNA mono-

layers [101]. The voltammetry for the ferrocyanide (3−/4−) redox

markers at the DNA-modified nanoAu-GEC is slightly affected by the

electrostatic interactions with the polyanionic layer, in contrast to

previous reported results for SAMs of DNA in continuous gold elec-

trodes [101], confirming the laxity of the DNA layer created on the

nanoAu(7.5%)-GEC electrode. These data suggest an architecture

that is made up of a disperse layer of oligonucleotide immobilized on

the isolated gold NPs and confirms the microscopic pattern achieved

by SEM and fluorescence microscopy [99].

Instead of SAMs on continuous layers of gold, isolated gold NPs

are able to produce “bioactive chemisorbing islands” for the immobi-

lization of thiolated biomolecules, avoiding stringent conditions for

surface coverage as well as the use of auxiliary reagents such as lat-

eral spacer thiols. Less compact layers are thus achieved favoring the

biological reaction on biosensing devices. Hybridization efficiency is

expected to be higher on the edging of the gold NPs surrounded by

nonreactive graphite-epoxy composite.

Briefly, the procedure consists of the following steps, as schemat-

ically outlined in Fig. 3.7: (i) thiolated probe immobilization by

chemisorption (Fig. 3.7A); (ii) hybridization with the complemen-

tary probe modified with either biotin or digoxigenin (Fig. 3.7B1);

(iii) enzyme labeling of the DNA duplex using streptavidin-HRP

or anti-DIG-HRP (Fig. 3.7C); and (iv) amperometric determination



based on the enzyme activity by adding H2O2 and using hydro-

quinone as a mediator (Fig. 3.7D).

The chemisorbing ability of gold NPs in the nano-AuGEC was

demonstrated with an excellent LOD (9 fmol/60 pM of ssDNA) in

hybridization studies. Regarding other electrochemical transducers

previously reported, such as an avidin graphite-epoxy biocompos-

ite (Av-GEB) or protein A graphite-epoxy biocomposite (ProtA-GEB)

[23], the main advantages of the inorganic nanoAu-GEC electrode

compared to the biocomposite is the lack of loss of activity and that

the latter requires the temperature to be kept at 4◦C due to the bio-

logical nature of the modifier, the protein avidin.

Moreover, and for the first time, a double-tagging PCR strat-

egy was performed with a thiolated primer for the detection of

Salmonella sp. The rapid electrochemical verification of the ampli-

con coming from the pathogenic genome of Salmonella performed

by PCR with a set of two labeled primers was demonstrated to be

an easy way for the thiolation of the PCR product. The thiolated end

allowed the immobilization of the amplicon on the nano-AuGEC elec-

trode in an easy way.

The procedure consists briefly of the following steps, as schemat-

ically outlined in Fig. 3.7: (i) DNA amplification and double-labeling

of Salmonella IS200 insertion sequence; (ii) immobilization of

the doubly labeled amplicon in which the SH end of the dsDNA

amplicon was immobilized on the nanoAu-GEC nanocomposites

by chemisorption (Fig. 3.7B2); (iii) enzymatic labeling using as

enzyme label the antibody anti-DIG-HRP capable of bonding the

other labeled extreme of the dsDNA amplicon (Fig. 3.7C); and (iv)

amperometric determination (Fig. 3.7D). With this strategy, as low

as 200 fmol can be easily detected, with an electrochemical signal of

almost 3 μA. This double tagging PCR strategy opens new routes not

only for immobilization purposes, but also act as an easy strategy for

labeling with gold or quantum dots during PCR.

The nanoAu-GEC material shows interesting properties for elec-

trochemical genosensing in hybridization experiments and very

promising features for electrochemical biosensing of a wide range of

biomolecules, such as dsDNA, PCR products, affinity proteins, anti-

bodies, or enzymes.


Final Remarks 93

3.5 Final Remarks

GEC-based platforms are useful and versatile transducer materials

for electrochemical DNA genosensing.

DNA can be attached directly onto a GEC surface by simple

adsorption (the simplest immobilization method and the easiest to

automate), avoiding the use of procedures based on previous activa-

tion/modification of the surface transducer and subsequent immo-

bilization, which are tedious, expensive, and time-consuming. This

procedure implies multisite attachment. Although DNA has been

widely attached onto carbonaceous materials, the underlying mech-

anism of adsorption has not been fully clarified. Adsorption is a

complex interplay between the chemical properties, structure, and

porosity of the substrate surface with the molecule being adsorbed.

DNA is a structurally polymorphic macromolecule which, depending

on nucleotide sequence and environmental conditions, can adopt a

variety of conformations. As a highly negatively charged molecule,

dsDNA is considered a hydrophilic molecule.

While dsDNA only partially shows its hydrophobic domain

through its major and minor grooves or through those sites where

dsDNA is open and exposing DNA bases, ssDNA has the hydrophobic

bases freely available for their interactions with hydrophobic sur-

faces.

These structural and chemical differences between ssDNA and

dsDNA are reflected in different adsorption patterns for both the

molecules.

The greater size and the more rigid shape of dsDNA with respect

to ssDNA are other parameters affecting their adsorption. Besides

the DNA molecule and the solid support, the solvent (normally

water), in particular the ionic strength, pH, and the nature of the

solutes, plays an important role in the adsorption process, mainly

in the stabilization of the adsorbed molecule on the substrate.

The hybridization event can be detected both with label-free or

enzymatic labeling procedures.

The single-point attachment of DNA can be achieved by the

immobilization of biotinylated DNA on Av-GEB platform. In this case,

a one-step immobilization/hybridization procedure is achieved. The

capability of surface regeneration of the biocomposite electrodes



allows repeated analyses with the same electrode as the Av-GEB

platform can be considered not only the transducer but also the

reservoir for the avidin.

The same immobilization strategy based on avidin/biotin link-

age can be achieved on magnetic beads. After their efficient mod-

ification, the magnetic beads can be easily captured on an m-GEC

transducer for the electrochemical determination of the hybridiza-

tion event.

The sensibility conferred by the m-GEC electrode in connection

with the use of magnetic beads and enzymatic labeling results in a

rapid, cheap, robust, and environment-friendly device which allows

the detection of pathogenic species on food, environmental, and clin-

ical samples.

The single-point attachment of DNA can also be achieved by

chemisorption of thiolated DNA on nanoAu-GEC platform. The

capability of surface regeneration of the nanocomposite electrodes

allows repeated analyses with the same electrode.

GEC materials present a low non-specific adsorption either for

DNA probes or enzyme labels. They do not require blocking steps to

minimize the non-specific adsorption on the free sites of the trans-

ducer.

Although the non-specific adsorption issues are controlled on

GEC, stringent conditions can be achieved when using avidin/biotin

linkage than when DNA is simply adsorbed, allowing rigorous con-

ditions for hybridization over longer times.

DNA biosensors based on GEC meet the demands of genetic

analysis, especially in food, biotechnology, and pharmaceutical

industries, while also generating new possibilities for the develop-

ment of DNA biosensors that are sensitive, robust, low cost, and eas-

ily produced.

For all the aforementioned reasons, it is possible to conclude that

GEC-based materials are very suitable platforms for DNA analysis.

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Chapter 4

Gold Nanoparticle-BasedElectrochemical DNA Biosensors

Marıa Pedrero, Paloma Yanez-Sedeno, and Jose M. PingarronDepartamento de Quımica Analıtica, Facultad de Ciencias Quımicas,Universidad Complutense de Madrid, Avenida Complutense s/n, E-28040 Madrid, [email protected]

Nowadays, gold nanoparticles play a key role in the construction of

a new generation of biosensors and, in particular, electrochemical

biosensors. This chapter is devoted to electrochemical DNA biosen-

sors coupled with the use of gold nanoparticles to improve both

oligonucleotide immobilization on electrode surfaces and signal

amplification for sensitive detection of hybridization events. Recent

advances in the development of these biosensors are covered,

stressing on the improvement of the analytical performance.

Configurations used for DNA immobilization and signal transduction

and amplification strategies are treated separately.

4.1 Introduction

The use of nanomaterials for the construction of efficient and

powerful biosensing devices constitutes nowadays one of the

research lines with more activity and effort in modern bioanalytical



104 Gold Nanoparticle-Based Electrochemical DNA Biosensors

chemistry. The extremely promising prospects of nanomaterials-

based biodevices accrue from the unique properties of nanomateri-

als, making possible advanced applications where limits of detection

at zeptomolar concentrations and ultra-sensitive multiplexed detec-

tion can be achieved [1]. Different types of nanomaterials can be

employed to design and construct these biosensing devices: carbon

nanotubes, nanowires including metal, silicon, conducting polymer,

and metal oxide nanowires, nanocantilivers, quantum dots, and

nanoparticles, including metal, metal oxide, semiconductor, and

magnetic nanoparticles. A nice overview on the use of nanomaterials

for the construction of biosensors can be found in the monography

edited by Kumar [1].

In this context, the preparation of nanostructured electrode

surfaces constitutes also a priority research line with high activity

in the field of electroanalytical chemistry [2]. This electrode

modification strategy combines, on the one hand, advances in

sensor technology, offering a wide range of approaches using or

not biological systems, as well as several (bio)assay-transduction

symbiotic strategies, and, on the other hand, the applications of

nanotechnology in its wider sense as the products, processes, and

systems operating at nanometric scale. The use of nanostructured

electrode surfaces produces significant advantages from the electro-

analytical point of view. In general, they improve the kinetics of the

electron-transfer reactions, exhibit electrocatalytic ability toward

many electrochemical processes of biological significance allowing

the detection potentials to be lowered, and show an anti-fouling

capability for the products of many electrochemical reactions. These

characteristics improve basic analytical properties such as the

sensitivity and selectivity of the methods and the repeatability of the

measurements.

Although most of the nanomaterials mentioned above can be

employed for this purpose, gold nanoparticles play a key role in the

construction of a new generation of biosensors and, in particular,

of electrochemical biosensors. The ability of gold nanoparticles to

provide a stable surface for the immobilization of biomolecules

retaining their biological activity is a major advantage for the

preparation of biosensors. Moreover, gold nanoparticles allow

direct electron transfer between redox proteins and bulk electrode


Introduction 105

materials to be performed thus enabling electrochemical sensing

with no need for electron-transfer mediators. Properties of gold

nanoparticles, such as their high surface-to-volume ratio, high

surface energy, the ability to decrease the distance between proteins

and metal particles, and their performance as electron-conducting

pathways between the prosthetic groups and the electrode surface,

have been claimed as reasons to improve electron transfer between

redox proteins and electrodes [3]. Also, gold nanoparticles have

shown to constitute useful interfaces for the electrocatalysis of redox

processes of molecules such as H2O2, O2, or NADH involved in the

biochemical reactions with analytical significance [4].

This chapter is devoted to electrochemical DNA biosensors

coupled with the use of gold nanoparticles to improve both

oligonucleotide immobilization on electrode surfaces and signal

amplification for sensitive detection of hybridization events. Elec-

trochemical genosensors have demonstrated in recent years to

constitute reliable alternatives for applications directed to gene

analysis, detection of genetic disorders, tissue matching and

forensics, due to their high sensitivity, small dimensions, low cost,

and compatibility with microfabrication technology. Besides tra-

ditional electrochemical transduction of DNA hybridization events

involving electroactive indicators/intercalators or enzyme tags,

the use of nanoparticles, especially of gold nanoparticles, offers

elegant pathways for interfacing such events with electrochemical

signal transduction and for amplification of the resulting electrical

response [5].

Gold nanoparticles can be employed for improving the immo-

bilization of DNA on electrode surfaces and thus for increasing

the hybridization capacity of the modified surface [6]. The use of

gold nanoparticles supporting films constructed by self-assembling

of 16-nm diameter colloidal gold onto a cystamine-modified gold

electrode resulted in surface densities of oligonucleotides as high as

4 × 1014 molecules cm−2, allowing a detection limit of 500 pM to be

achieved.

The other fundamental advantage of gold nanoparticles-based

electrochemical DNA biosensors is the development of amplification

routes for the DNA sensing events. According to Willner et al. [7], the

concept of the amplified detection of DNA using gold nanoparticles



Figure 4.1. Schematic amplification route for DNA sensing using an

oligonucleotide-functionalized nanoparticle as an amplifying unit.

is outlined in Fig. 4.1. The primary probe-DNA recognition event

is accompanied by the coupling of a colloidal gold tag, which is

followed by acid dissolution and anodic stripping electrochemical

measurement of the metal tracer. Sensitivity can be enhanced to

sub-picomolar detection limits by catalytic enlargement of the gold

tracer in connection to nanoparticle-promoted precipitation of gold

[8] or silver [9, 10].

More complex amplification strategies involve dendritic

nanoparticles arrays or coupling with probe-coated magnetic beads.

The former approach can be visualized in Fig. 4.2 [7]. Two

types of functionalized gold nanoparticles are employed. One

of them is functionalized with a 3’-terminated oligonucleotide

complementary to the 3’-end of the target DNA, whereas the second

type of gold nanoparticle is functionalized with the 5’-terminated

oligonucleotide complementary to the 5’-end of the target DNA.

The latter approach involves hybridization of probe-coated

magnetic beads with gold-tagged DNA targets, giving rise to three-

dimensional structures of magnetic beads cross-linked together

through the DNA and gold nanoparticles. No aggregation was

observed in the presence of noncomplementary or mismatched

oligonucleotides [5].

This chapter covers recent advances in the development of DNA

electrochemical biosensors making use of gold nanoparticles to

improve their analytical performance. Aptamer-nanoparticle-based

biosensors will be also covered since using exponential selection

strategies, various RNA and DNA sequences have been identified

that bind small molecules and proteins while inducing a change


Configurations Used for DNA Immobilization 107

Figure 4.2. Dendritic amplification of DNA sensing by oligonucleotide-

functionalized gold nanoparticles.

in nucleic acid tertiary structure [11]. In order to contribute to

a better understanding of the chapter content, we have decided

to separate the covering of immobilization strategies from that of

detection strategies, although, as it will be outlined below, many

immobilization approaches are oriented to achieve amplification of

the resulting analytical signals.

4.2 Configurations Used for DNA Immobilization

As it is well known, the immobilization of DNA probes onto electrode

surfaces is one of the key steps in DNA sensor development. It

has been widely demonstrated that the DNA sensor performance

(e.g., sensitivity, selectivity, and stability) is highly dependent on

the characteristics of DNA probes’ immobilization approaches.

One of these relies on the use of nanomaterials such as gold

nanoparticles (Au-NPs), taking advantage of their unique electrical

conductivity, biocompatibility, and ease of self-assembly through



a thiol group. Large specific surface Au-NPs-modified electrodes

can enhance the amount of DNA immobilized onto the electrode

leading to an improvement of the biosensor performance. Many of

the examples reported in the literature imply a previous chemical

preparation of Au-NPs, with a subsequent modification of the

electrode surface through physical adsorption or chemical linking.

Nanomaterial-based platforms suitable to construct biosensors

are fabricated in this way. Alternatively, direct electrochemical

deposition of Au-NPs onto the electrode surface constitutes a rapid,

clean, and versatile mode to create nanomaterial platforms in situ for

the construction of DNA biosensors. The next sections revise some

illustrative examples of several designs employed to prepare such

platforms.

4.2.1 Au-Thiol Binding

Gold electrode substrates have attracted special attention for

the preparation of electrochemical DNA biosensors since DNA

can be strongly bound at the surface of gold through Au–thiol

binding. Thiolated DNA can be monolayered on gold in a self-

assembly manner, which provides stable and structurally well-

defined electrochemical interfaces. However, one major drawback

of the gold electrode is attributed to the cleaning step; in order

to obtain reproducible results, the gold electrode needs to be

mechanically polished and then electrochemically etched in acid

solutions. This time-consuming process determines a quality of

electrochemical DNA biosensors.

Various recent interesting configurations of genosensors take

advantage of the Au–thiol binding strategy with gold nanoparticles,

for example, gold nanoparticles were electrodeposited on screen-

printed electrodes which were subsequently modified with a self-

assembled monolayer of thiol-capped single-stranded DNA (capture

probe) (Fig. 4.3). This immobilization strategy was employed to

construct a genosensor for the detection of the rfbE gene, which is

specific to E. coli O157 [12].

A similar immobilization methodology has been used to develop

a DNA-sensing platform for Helicobacter pylori [13]. Rigid conduct-

ing gold nanocomposites have been also modified with this strategy.



Figure 4.3. Scheme displaying the DNA immobilization strategy involving

thiol binding [12] (adapted with permission of the American Chemical

Society).

For example, using Au-NPs-graphite epoxy composites, islands of

chemisorbing material ( Au-NPs) surrounded by nonreactive, rigid,

and conducting graphite epoxy composites are achieved to avoid the

stringent control of surface coverage parameters required during

immobilization of thiolated oligonucleotides in continuous gold

surfaces [14]. The spatial resolution of the immobilized thiolated

DNA can be easily controlled by merely varying the percentage of

gold nanoparticles in the composite.

Impedimetric genosensors were also constructed by making

use of gold nanoparticles electrodeposited on the surface of a

gold electrode, and subsequent immobilization of probe DNA

on the surface of gold nanoparticles through a 5’-thiol-linker

[15]. Electrochemical impedance spectroscopy (EIS) was used to

investigate probe DNA immobilization and hybridization. Compared

to the bare gold electrode, the gold nanoparticle-modified electrode

improved greatly the density of probe DNA attachment and the

sensitivity of DNA sensor.

Shen et al. [16] have recently reported the development of an

electrochemical DNAzyme biosensor based on DNA-Au bio-bar code

amplification, which provides a platform for fabrication of sensors

for analysis of many small molecules, especially for metal ions. For

example, a specific DNAzyme for Pb2+ was immobilized onto an

Au electrode surface via a thiol–Au interaction, taking advantage of

catalytic reactions of a DNAzyme upon its binding to Pb2+ and the

use of DNA-Au bio-bar codes to achieve signal enhancement [16].

The presence of gold nanoparticles, enhancing the active surface



area, allows a larger number of short DNA sequences to be bound,

leading to a substantial amplification of signals for ultrasensitive

detection.

4.2.2 Gold Nanoparticles: Metallic Oxide Composites

Metallic oxides have been used in combination with gold nanopar-

ticles to prepare electrode surfaces with improved stability and/or

response capacity for DNA detection. Among them, zirconia (ZrO2)

has been used in various applications due to its thermal stability,

chemical inertness, lack of toxicity, and affinity for the groups

containing oxygen. Thus, it is an ideal candidate material for the

immobilization of biomolecules with oxygen groups. The approach

used for the preparation of an electrochemical DNA biosensor

based on zirconia and gold nanoparticles is depicted in Fig. 4.4. A

gold nanoparticle film was electrodeposited onto a glassy carbon

electrode, and then a zirconia thin film was prepared on the Au-

NPs/GCE by cyclic voltammetry in an aqueous electrolyte of ZrOCl2

and KCl. DNA probes were attached onto the ZrO2/Au-NPs/GCE

due to the strong binding of the phosphate group of DNA with the

zirconia film and the excellent biocompatibility of nanogold with

DNA [17].

Thin gold films deposited by low pressure gold sputtering

or electrochemical deposition can provide a highly sensitive and

reproducible electrode for the preparation of DNA biosensors

without the requirement of the cleaning step. However, the thin

gold film directly sputtered on a substrate very easily peels off

Figure 4.4. Schematic representation of the DNA immobilization on a

ZrO2/ Au-NPs/GCE [17] (adapted with permission of Elsevier).



during immobilization or electrochemical measurement, giving less

reliability. To avoid this problem, DNA biosensors using a thin

gold film sputtered on capacitive anodic nanoporous niobium

oxide were proposed [18]. The nanoporous niobium oxide offers

a good adhesion as well as an enhancement of redox signals by

accumulation of charges in between the gold film and the niobium

oxide. The mechanism of enhancing the signal by the thin gold film

on nanoporous niobium oxide is in part attributed to capacitive

niobium oxide and is in part ascribed to the bridged thin gold

film.

4.2.3 Carbon Nanotube–Gold Nanoparticle Hybrids

Carbon nanotubes (CNTs) are widely recognized as an ideal support

for fabricating electrochemical sensors with a high sensitivity

and selectivity [19]. Due to the ability of carbon nanotubes

to promote electron-transfer reactions, and the high catalytic

activity and biocompatibility of gold nanoparticles, the developed

genosensors showed excellent reproducibility and stability under

the DNA hybridization conditions [20]. A novel DNA biosensor

was constructed by layer-by-layer (LBL) covalent assembly of

gold nanoparticles and multiwalled carbon nanotubes (MWCNTs).

Cysteamine molecules acted as a glue to connect activated MWCNTs

and Au-NPs into a three-dimensional hybrid network on a gold

electrode. Then, NH2-ssDNA was immobilized on multilayer films via

amino link at the 5’-end.

4.2.4 Polymer–Gold Nanoparticle Hybrids

Because electrochemical polymerization allows the control of film

thickness, permeation, and charge-transport characteristics by

adjusting the electrochemical parameters, this methodology has

demonstrated to be a promising approach to immobilize DNA

probes. Among other electronconducting polymers, polyaniline

(PANI) has attracted a special attention in the field of conducting

macromolecules. Due to its homogeneity, unique redox properties,

high electrical conductivity and strong adherence to electrode sur-

face, PANI has been extensively applied to develop electrochemical



nanoPAN

Dispersed withDMF and chitosan

– 0.2 V, 500 s 25 °C, 2h

DNA probeHAuCl4

Figure 4.5. Schematic diagram of the immobilization of DNA on a Au-

NPs/nanoPANI/GCE [21] (adapted with permission of Elsevier).

biosensors. Furthermore, PANI hybrid materials constituted of the

polymer and metal nanoparticles have been reported to play an

important role for the design of novel efficient electrochemical

biosensors. As an example, Fig. 4.5 displays the formation of gold

nanoparticle/polyaniline nanotube membranes on a glassy carbon

electrode for the electrochemical sensing of the immobilization and

hybridization of DNA [21]. The synergistic effect of the two kinds of

nanoparticles, nanogold and nanoPANI, enhanced dramatically the

sensitivity for the DNA hybridization recognition in this particular

case, a DNA sequence-specific phosphinothricin acetyltransferase

gene (PAT) existing in some transgenic crops.

PNA is a structural DNA analogue containing a neutral N-

(2-aminoethyl)-glycine pseudopeptide backbone to which the

nucleobases are linked. The lack of negative charges on these

molecules allows strong base-pairing interactions with ssDNA.

PNA shows very high specificity in DNA recognition. Nanogold-

modified electrodes can largely increase the ssPNA capture probe

immobilized amount leading to an increase of the electrical signal.

As an example, single-stranded PNA probes were immobilized on

a nanogold-modified electrode for the label-free detection of DNA–

PNA hybridization using a water-soluble ferrocene-functionalized

polythiophene transducer [22]. The ferrocene-containing cationic

polythiophene do not interact electrostatically with the PNA probes

due to the absence of the anionic phosphate groups. However, after

DNA–PNA hybridization, the cationic polythiophene is adsorbed on



the DNA backbone, giving a clear hybridization detection signal by

differential pulse voltammetry.

Electrochemical DNA biosensors were also prepared using

carbon nanotube-polymer hybrids in combination with gold

nanoparticles. Composite materials based on integration of CNTs

and polymers have gained growing interest because they possess

the properties of each component with a synergistic effect. In a

recent report [23], p-aminobenzoic acid (PABA), which contains

electron-rich N atom and high electron density of carbonyl group,

was electrodeposited by cyclic voltammetry on the surface of a

multiwalled carbon nanotube-modified glassy carbon electrode.

Gold nanoparticles were subsequently electrodeposited onto the

surface of the PABA/MWNTs composite film, and the probe DNA was

immobilized on the surface of Au-NPs through an Au–S bond.

4.2.5 Avidin–Biotin Affinity Reactions

Electrochemical detection of bioaffinity interactions with a gold

nanoparticles sensing platform was accomplished by using

thrombin–thrombin binding aptamer couple as a model [24]. The

aptamer was immobilized on a screen-printed electrode modified

with gold nanoparticles by avidin–biotin technology. The cathodic

peak area resulting from the reduction of previously formed gold

oxide was found to be proportional to the thrombin quantity

specifically adsorbed onto the modified electrode surface.

Dendritic polymers (dendrimers) belong to a new class of

synthetic macromolecules possessing a regularly branched tree-

like structure. A great attention has been paid to the potential

applications of polyamidoamine (PAMAM) dendrimers for the

development of biosensors, because of the dendrimer high geo-

metric symmetry, chemical stability, controllable size, and surface

functionality. Dendrimer–Au-NPs nanocomposites have been shown

to combine the physical and chemical properties of Au-NPs with

the surface reactivity of dendrimers. A recent report describes an

electrochemical approach for sequence-specific DNA detection using

PAMAM and gold nanoparticles [25]. The biosensor design consisted

of a gold electrode modified with 3-mercaptopropionic acid, which

was reacted with an amino-terminated polyamidoamine (PAMAM,



G 4.0-NH2) to obtain a thin film. Single-stranded 3’-biotin end-

labeled oligonucleotide was immobilized onto the film to obtain a

stable recognition layer through biotin–avidin combination to detect

complementary target.

4.3 Signal Transduction and Amplification Strategies

Recent literature dealing with integration of Au-NPs with DNA

detection systems shows that different strategies can be employed to

achieve improved analytical performance of the resulting genosen-

sors. As mentioned above, using Au-NPs-modified electrodes, the

amount of DNA immobilized onto the electrode can be considerably

enhanced. The efficient immobilization of DNA onto the transducer

paves the way for the design of effective signal transduction

approaches of the hybridization event. Furthermore, Au-NPs have

been employed as amplification components which when combined

with electrochemical techniques have given rise to the design of

selective and highly sensitive DNA sensors [26]. Castaneda et al. [27]

and Guo et al. [26, 28] have recently reviewed the achievements of

the electrochemical sensing of DNA using Au-NPs.

In a rather general approach, and in order to systematize the

content of this section, we have considered separately the detection

strategies that take advantage of the design of sensing platforms

integrating Au-NPs, but in which the detection methodology itself

does not involve Au-NPs as an element to obtain the signal trans-

duction, and those involving direct participation of Au-NPs in the

generation of the electroanalytical signals. In the next subsections,

some illustrative examples of each of these methodologies will be

commented, and tables summarizing other recent works using the

different detection strategies will be given.

4.3.1 Detection Strategies Not Involving DirectParticipation of Au-NPs in the Generationof the Electrochemical Signal

The immobilization methods commented in Sec. 4.2 can be

coupled with signal transduction methodologies involving different


Signal Transduction and Amplification Strategies 115

strategies such as the detection of redox markers, the detection

based on enzymatic labels, the detection based on electrochemical

labels intercalated within ds-DNA, and the use of Au-NPs as carriers

for other nanoparticles or other electrochemical labels which are

responsible for the generation of the analytical signals.

4.3.1.1 Direct detection of redox markers

Figure 4.3 showed an example involving DNA immobilization

through thiol binding onto gold nanoparticles electrodeposited on

screen-printed electrodes. This sensor architecture allowed the

development of an electrochemical sensor for the detection of E.coli O157 based on competition between the target gene (com-

plementary to the capture probe DNA) and reporter DNA-tagged,

hexaammineruthenium (III) chloride–encapsulated liposomes. The

current signal of the released liposomal [Ru(NH3)6]3+ was mea-

sured using square wave voltammetry (SWV), yielding a sigmoidal-

shaped dose-response curve whose linear portion was over the

range from 1 to 106 fmol. This liposomal competitive assay provided

an amplification route for the detection of the rfbE gene (specific to

E. coli O157) at ultratrace levels, with a detection limit of 0.75 amol

[12].

Another interesting example of direct detection of redox markers

is a ferrocene catalyzed aptamer-based thrombin sensor [29].

Figure 4.6 displays the sensor architecture and functioning. A

thrombin binding aptamer was covalently immobilized onto three-

dimensional Au-NP–doped conducting polymer nanorod electrodes

(Au-NPs/3D-CPNEs). Ferrocene was attached with anti-thrombin

through streptavidin-biotin interactions and it electrochemically

catalyzed the oxidation of ascorbic acid. Since thrombin was

sandwiched between thrombin aptamer and anti-thrombin anti-

body attached with ferrocene, the catalytic current response was

proportional to the thrombin concentration. The aptamer sensor

showed a dynamic range from 5 to 2000 ng L−1 with a detection

limit of 5 ng L−1 (0.14 pM) and it was applied to the determination

of spiked concentrations of thrombin in real human serum samples.

Another interesting approach involves the use of quantum dot

tracers. For example, Huang et al. [30] have described a DNA



3D-CPNEs

Figure 4.6. Schematic illustration of the fabrication of the Apt/3D-CPNEs-

based thrombin aptamer sensor [29] (reprinted with permission of the

American Chemical Society).

biosensor where the target DNA was immobilized on AuNPs films

bound to the surface of a chitosan-entrapped carbon paste electrode

(CPE). The probe DNA was labeled with CdSe quantum dots, and

the CdSe was loaded on the electrode surface via DNA hybridization

and then dissolved in HNO3. The released Cd2+ was detected

by differential pulse anodic stripping voltammetry (DPASV). The

dynamic detection range for 18-base DNA specific sequence of the

cauliflower mosaic virus gene was 5.0 × 10−12 to 5.0 × 10−7 M, with

a detection limit of 6.5 × 10−13 M.

Table 4.1 summarizes data reported between 2007 and 2009 for

DNA sensors where the detection strategy is related to the direct

detection of redox markers.

4.3.1.2 Detection based on enzymatic labels

Another detection strategy that takes advantage of the use of Au-

NPs in the immobilization step to improve the immobilization

and orientation of DNA strands is involving enzyme reactions in

the detection process. A nice recent example is the one reported



Table 4.1. Gold nanoparticle-based electrochemical DNA biosensors using

direct detection of redox markers

Modified Electrode Detection Detection

Analyte Substrate Technique Limit Reference

DNA synthetic strands GCE DPV (MB) 10−10 mol L−1 31

Adenosine AuE CV (Ferrocene) 2.0 × 10−8 M 32

DNA synthetic strands AuE DPV (ferrocene- 1.0 × 10−11 M 28

polythiophene)

Cocaine AuE SWV (ferrocene) 0.5 μM 33

Cancer antigen 15-3 AuE CV (Prussian blue) 0.6 ng mL−1 34

Cauliflower mosaic CCPE DPASV (Cd2+) 6.5 × 10−13 M 35

virus gene

E. coli O157 specific gene SPE SWV [Ru(NH3)6]3+ 0.75 amol 12

(carbon)

DNA synthetic strands AuE DPV [Ru(NH3)6]3+ 1 × 10−11 M 36

Hydrogen peroxide AuE Amperometry 2.0 μM 37

Thrombin 3D-CPNE CV (ascorbic acid) 5 ng L−1 (0.14 pM) 29

DNA damage GCE DPV [Ru(NH3)6]3+ 0.05 mg mL−1 38

AuE: gold electrode; CCPE: chitosan-entrapped carbon paste electrode; CV: cyclic voltammetry;

3D-CPNE: three-dimensional conducting polymer nanorods electrodes; DPASV: differential pulse

anodic stripping voltammetry; DPV: differential pulse voltammetry; GCE: glassy carbon electrode;

SPE: screen-printed electrode; SWV: square wave voltammetry.

by Brasil de Oliveira Marques et al. [39] for the electrochemical

biosensing of Salmonella sp. They used for the first time a double

tagging PCR strategy based on the double labeling of the amplicon

during PCR with a digoxigenin and a –SH set of labeled primers.

The thiolated end allowed efficient immobilization of the amplicon

on an Au-NP–modified graphite-epoxy composite electrode, while

digoxigenin allowed the electrochemical detection with the antiDIG-

HRP reporter to be performed in the femtomole range (Fig. 4.7).

An interesting detection approach combining enzymatic elec-

trochemical detection and silver precipitation is that reported

by Martınez-Paredes et al. [40]. They used the enzyme alkaline

phosphatase (AP) to catalyze the dephosphorylation of the substrate

3-indoxyl phosphate thus producing a compound able to reduce

silver ions in solution into a metallic deposit localized where

the enzymatic label is attached. The deposited silver is then

electrochemically stripped into solution and measured by anodic

stripping voltammetry (ASV). The DNA hybridization assay was



red

Figure 4.7. Schematic illustration of the amperometric detection of

DNA hybridization based on an enzymatic reaction [39] (reprinted with

permission of the American Chemical Society). See also Color Insert.

carried out on a Au-NP–structured screen-printed carbon electrode,

and the sequence chosen as target is included in the 29 751-base

genome of the SARS (severe acute respiratory syndrome)-associated

coronavirus. A linear range was found for the biotinylated target

between 2.5 and 50 pmol L−1 with a detection limit of 2.5 pmol L−1.

Table 4.2 summarizes works reported between 2007 and 2009

using an enzymatic approach to detect hybridization.

4.3.1.3 Detection based on electrochemical labelsintercalated within dsDNA

DNA sensing platforms constructed with Au-NPs (and other NPs) to

enhance the DNA immobilization and hybridization performance by

increasing the electrode area have also employed for the detection



Table 4.2. Gold nanoparticle-based electrochemical DNA biosensors using

enzymatic detection

Modified Electrode Detection

Analyte Substrate Technique Detection Limit Reference

Synthetic DNA strands CPE DPV 5.0 × 10−11 M 41

H2O2 AuE CV Chronoam- 1.3 μM 42

perometry

Salmonella sp. nanoAu-GEC Amperometry 9 fmol (60 pM) 39

Synthetic Au/SPE Amperometry – 43

oligonucleotides

Synthetic 30-mer SPCE CV 2.5 pmol/L 40

oligonucleotides

Au/SPE: gold screen-printed electrode; nanoAu-GEC: nanogold graphite-epoxy composite.

step a series of compounds exhibiting a significant difference in

their voltammetric signals in the presence of ssDNA or dsDNA. Du

et al. [44] recently fabricated a DNA biosensor by the sequential

modification of gold electrodes with Au-NPs and CdS-NPs. The

modified electrode was applied for the detection of target DNA with

Co(phen)22+ as hybridization indicator. The target DNA sequence

was quantified over the range 2.0 × 10−10 to 1.0 × 10−8 M, with

a detection limit of 2.0 × 10−11 M.

In a similar methodology, methylene blue (MB) has recently

been used as electrochemical indicator for the preparation of an

adenosine triphosphate (ATP) aptasensor [45]. Au-NPs claimed to

make more MB interact with DNA on the sensing interface. The DPV

peak current corresponding to the MB oxidation decreased with

ATP concentration in the 1 × 10–10 M to 1 × 10–7 M range, with a

detection limit of 0.1 nM. The sensor showed advantages regarding

low-cost, rapidity, simple detection, and reusability.

Intercalated doxorubicin has been also used as an electro-

chemical label in the detection of DNA hybridization events

in a genosensor built by layer-by-layer covalent attachment of

multiwalled carbon nanotubes and Au-NPs [46]. The oxidation peak

current obtained by differential pulse voltammetry showed a linear

relationship with the logarithm of the target DNA concentration in

the range 5.0 × 10−10 to 1.0 × 10−11 M, with a detection limit of

6.2 pM.



Also, intercalated adriamycin has been used as hybridization

label in a genosensor built by modifying a glassy carbon electrode

(GCE) with multiwalled carbon nanotubes with carboxyl groups and

Au-NPs [23]. Differential pulse voltammetry (DPV) was utilized to

monitor the DNA hybridization event. Under the optimal conditions,

the increase of reduction peak current of adriamycin was linear

with the logarithm of the concentration of the complementary

oligonucleotides from 1.0 × 10−12 to 5.0 × 10−9 M with a detection

limit of 3.5 × 10−13 M.

An interesting and relatively new water soluble intercalating

compound, pentaamin ruthenium [3-(2-phenanthren-9-yl-vinyl)-

pyridine] complex [Ru(NH3)5L] prepared in situ, has been also

used to detect the hybridization event. The metal provides with

a redox centre that can be used as the electrochemical indicator,

while gold nanoparticles contribute to facilitate the electron transfer

between the redox indicator and the electrode surface by acting

as tiny conduction centres. Following this detection strategy,

complementary target sequences of H. pylori were detected over the

range 40 to 800 pmol with a detection limit of 25±2 pmol [13].

Table 4.3 summarizes data reported between 2007 and 2009

for electrochemical genosensors based on dsDNA-intercalated

electrochemical labels.

4.3.1.4 Detection involving the use of Au-NPs as carriers

Au-NPs can be used as nanocarriers for other NPs or other

electroactive species. As a result, the DNA detection is notably

enhanced compared to the use of single labels. However, this

strategy requires a careful work to avoid irreproducibility in the

biosensors’ response [27].

In this context, Ding et al. [50] have recently described a sensitive

assay for sequence specific DNA detection based on bio-bar code

techniques, and using electrochemical detection of Cd ions dissolved

from CdS-nanoparticles. Figure 4.8 shows a scheme of the method

used for the amplified sensing of target DNA. Sandwich-type DNA

complexes were fabricated with a thiol-functionalized capture DNA

sequence immobilized on an Au-NP–GCE and hybridized with one

end of target DNA. The other end of target DNA was recognized with



Table 4.3. Electrochemical genosensors based on dsDNA-intercalated

electrochemical labels



DNA synthetic strands AuE DPV (MB) 0.1 pM 47

Adenosine AuE DPV (MB) 1 nM 48

Helicobacter pyroli AuE DPV [Ru(NH3)5L] 25 pmol 13

DNA sequence

Thrombin GCE DPV (MB) 0.5 nM 49

DNA synthetic strands AuE DPV (doxorubicin) 7.5 pM 20

DNA synthetic strands AuE DPV [Co(phen)2]2+ 2.0 × 10−11 M 44

ATP AuE DPV (MB) 0.1 nM 45

DNA synthetic strands AuE DPV (doxorubicin) 6.2 pM 46

DNA synthetic strands GCE DPV (adryamicin) 3.5 × 10−13 M 23

AuE: gold electrode; DPV: differential pulse voltammetry; EIS: electrochemical impedance spec-

troscopy; GCE: glassy carbon electrode; (Ru(NH3)5L: pentaamin ruthenium [3-(2-phenanthren-

9-yl-vinyl)-pyridine] complex.

Figure 4.8. Scheme displaying the fundamentals and electrochemical

detection of DNA hybridization through bio-bar code DNA probes of

amplification [50] (reprinted with permission of Elsevier).



signal DNA labeled on the surface of Au-NPs. In order to amplify

the detection signals, the Au-NPs were also modified with CdS-NPs.

Since a single Au-NP could be loaded with hundreds of signal DNA

probe strands, a significant amplification for the detection of target

DNA was achieved. The hybridization events were monitored by

measuring the Cd ions dissolved from the hybrids using differential

pulse voltammetry (DPV). The peak current values increased with

the target DNA concentration in the range 1.0 × 10−14 to 1.0

× 10−13, and a detection limit of 4.2 × 10−15 M was achieved. Two-

base mismatched sequences showed weaker peak current and non-

complementary sequences gave no response at all.

Another recent approach [51] consisted of the construction

of a DNA sandwich electrochemical biosensor based on the use

of PbS-NPs attached to AuNPs/bio-bar code–modified magnetic

microbeads (MBs). The magnetic carriers containing PbS-NPs

labeled DNA probe were immersed in an electrochemical cell

containing 1.0 M HNO3 and the released Pb2+ was measured by

anodic stripping voltammetry (ASV) at a mercury film electrode

(MFE). The target DNA gave a linear response in the range from

2.0 × 10−14 M to 1.0 × 10−12 M, with a detection limit of 5.0 ×10−15 M.

Similar detection strategies involve immobilization of DNA-

modified Au-NPs onto the working electrode surface through

hybridization, and the use of molecules binding to DNA as

electrochemical labels. As an example, Miao et al. [52] described a

sensing strategy for the detection of glutathione in fetal calf serum

based on the use of two Au electrodes and two complementary

oligonucleotides (Fig. 4.9). The surface of one AuE is modified

with one of the two oligonucleotides and then immersed in the

glutathione solution where, due to the ligand release effect, the

oligonucleotides are replaced by glutathione. When the second AuE

is immersed in the solution, the released oligonucleotide molecules

are immobilized onto this electrode surface. Then, Au-NPs modified

with the complementary oligonucleotide are added and immobilized

onto this electrode surface through hybridization. Large numbers of

[Ru(NH3)6]3+ are then localized onto the electrode surface via the

electrostatic interaction between the electrochemical species and

oligonucleotide molecules. Since the Au-NPs amplify the detection



Figure 4.9. Schematic illustration of ultrasensitive detection of glu-

tathione based on the use of Au-NPs as carriers of electrochemical labels

[52] (reprinted with permission of Elsevier).

signal, glutathione could be detected in the range from 1 × 10−12 to

1 × 10−10 M, with a detection limit of 4 × 10−13M.

Bio-bar code technique was also employed by Shen et al.[16] with an electrochemical DNAzyme biosensor. The DNAzyme

hybridizes to a specially designed complementary substrate strand

that has an overhang, which in turn hybridizes to the DNA-

Au bio-bar code (short oligonucleotides attached to 13-nm gold

nanoparticles). Upon binding of Pb2+ to the DNAzyme, the DNAzyme



catalyzes the hydrolytic cleavage of the substrate, resulting in

the removal of the substrate strand along with the DNA bio-

bar code and the bound [Ru(NH3)6]3+ from the Au electrode

surface. The release of Ru(NH3)3+6 results in lower electrochemical

signal of Ru(NH3)3+6 confined to the electrode surface. Because each

nanoparticle carries a large number of DNA strands that bind to the

signal transducer molecule Ru(NH3)3+6 , the use of DNA-Au bio-bar

codes enhances the detection sensitivity by 5 times, enabling the

detection of Pb2+ at a very low level (1 nM). The DPV signal response

of the DNAzyme sensor is negligible for other divalent metal ions,

indicating that the sensor is highly selective for Pb2+. Table 4.4

summarizes recent works found in the literature for genosensors

making use of Au-NPs as carriers for other NPs or other electroactive

species.

4.3.2 Detection Strategies Involving Direct Participationof Au-NPs in the Generation of theElectrochemical Signal

Other detection strategies involve direct participation of Au-

NPs in the generation of the electroanalytical signal used for

quantification of target DNA. These include the detection systems

based on Au-NPs dissolving, label-free electrochemical impedance

and conductimetric detection approaches, different methodologies

for signal enhancement by precipitation of silver or even gold onto

Au-NP–DNA conjugates and Au-NP enlargement strategies. Some

illustrative examples of these strategies are discussed below.

4.3.2.1 Detection based on Au-NPs’ acidic orelectrochemical dissolving

This detection procedure is based on the oxidative dissolution of

the Au-NPs bound to DNA into aqueous Au ions followed by their

electrochemical sensing. Chemical dissolution of the Au-NP tags has

been mainly carried out with a HBr/Br2 solution, this step being

followed by accumulation and stripping analysis of the resulting

Au(III) ions. Due to the high toxicity of the HBr/Br2 solution, other

oxidation methods have been also employed.

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Table 4.4. Electrochemical genosensors using Au-NPs as carriers

Analyte Modified Electrode Substrate Detection Technique Detection Limit Reference

DNA synthetic strands Silicon wafer nanogap Current-voltage measurement – 53

Thrombin Pyrolytic graphite DPV (adenine) 0.1 ng mL−1 54

EcoRI endonuclease AuE CV (ferrocene) – 55

Breast cancer-associated BRCA-1 mutant DNA AuE Chronocoulometry [Ru(NH3)6]3+ ∼fM 56

DNA synthetic strands NPGE Chronocoulometry [Ru(NH3)6]3+ 28 aM 57

Pb2+ AuE DPV 1 nM 16

DNA synthetic strands, IgG SPE DPV 30 fM (DNA) 25 fg/mL (IgG) 58

Adenosine AuE CV 1.8 × 10−10 M 59

Cytocrome c AuE CV 6.7 × 10−10 M 60

DNA synthetic strands GCE DPV (Cd2+) 4.2 × 10−15 M 50

DNA synthetic strands MFE ASV (Pb2+) 5.0 × 10−15 M 51

Thrombin AuE DPASV (Pb2+) 6.2 × 10−15 M 61

Hg(II) AuE SWV 0.5 nM (100 ppt) 62

Glutathione AuE Chronocoulometry [Ru(NH3)6]3+ 4 × 10−13 M 52

Synthetic oligonucleotides AuE DPV [Ru(NH3)6]3+ 1.4 × 10−11 M 25

ATP AuE Chronocoulometry [Ru(NH3)6]3+ 0.2 nM 63

Platelet derived growth factor AuE CV [Ru(NH3)5Cl]2+ 1 × 10−14 M 64

ASV: anodic stripping voltammetry; ATP: adenosine triphosphate; AuE: gold electrode; CV: cyclic voltammetry; DPV: differential pulse voltammetry; DPASV:

differential pulse adsorptive stripping voltammetry; GCE: glassy carbon electrode; MFE, mercury film electrode; NPGE: nanoporous gold electrode; SPE: screen-

printed electrode.



Figure 4.10. Schematic representation of the procedure for detection of

DNA hybridization using Au-NP–coated latex labels [65] (reprinted with

permission of acs).

Pinijsuwan et al. [65] loaded streptavidin-coated latex particles

with biotin-coated Au-NPs so as to increase the quantity of Au-NPs.

Then, they attached the latex particles to biotinylated DNA probes

for E. coli previously hybridized to a DNA-modified SPE (Fig. 4.10).

The detection step involved the immersion of the modified SPE in

a HBr/Br2 solution, and further differential pulse anodic stripping

voltammetry (DPASV) of Au3+ ions. Following this methodology, a

detection limit of 0.5 fM was achieved.

The procedures described by Castaneda et al. [66] and by

Zheng et al. [67] were based on the detection of Au-NPs through

their electrochemical oxidation to AuCl−4 at +1.25 V (vs. Ag/AgCl),

followed by a DPV scan resulting in an analytical signal due to

the reduction of AuCl−4 at +0.4 V. This method was applied for

the detection of DNA hybridization using two different approaches

[66]. The first one consisted of hybridization between a capture

DNA strand linked with paramagnetic beads and a target DNA

strand related to BRCA1 breast cancer gene which was coupled

with streptavidin–Au-NPs. The second design was based on a

sandwich assay where a cystic fibrosis related DNA strand was

used as the target and sandwiched between two complementary



DNA probes, the first one linked with paramagnetic beads and the

second one modified with Au-NPs via biotin-streptavidin affinity

reactions. Zheng et al. [67] employed this detection methodology

for the development of a specific electrochemical aptasensor for the

detection of thrombin. The sensor was based on a sandwich format

of magnetic nanoparticle/thrombin/Au-NP and signal amplification

by forming network-like thiocyanuric acid/Au-NPs. A detection limit

of 7.82 aM was achieved.

4.3.2.2 Label-free electrical detection

The development of label-free biosensors is a clear trend in modern

biotechnology due to the numerous technical advantages that these

types of biosensors offer when compared with those needing

chemical labeling. Related to this, a great attention is being paid

to the development of micro-biosensors based on direct electrical

measurement of impedance, resistance, capacitance, perturbation

current, or charge. In fact, due to characteristics of electrical

transduction methods such as affordable instrumentation, excel-

lent compatibility with advanced semiconductor technology and

miniaturization, direct electrical detection methods have become

suitable candidates for the next generation of DNA sensors [68].

Moreover, this kind of detection can be tailored as extremely

sensitive with a high multiplexing capability and combined with the

unique electrical properties of metal nanoparticles, make electrical

detection systems as excellent prospects for the designing of DNA

detection devices.

Impedance-based detection has been employed for the label-

free detection of target DNA by measuring the difference between

the charge-transfer resistance (Ret) value at a DNA-immobilized

polyaniline nanofibers/carbon paste electrode (PANnao/CPE) mod-

ified with nanogold and carbon nanotubes composite nanoparticles,

and that at the hybridized electrode [69]. The approach was

applied to determine the sequence-specific DNA of phosphinothricin

acetyltransferase (PAT) gene and the polymerase chain reaction

amplification of nopaline synthase gene from transgenically modi-

fied beans. The dynamic range for detecting the PAT gene sequence



was from 1.0 × 10−12 to 1.0 × 10−6 M, with a detection limit of 5.6 ×10−13 M.

Furthermore, as commented in Sec. 4,2.1, impedimetric genosen-

sors were also constructed by making use of gold nanoparticles

electrodeposited on the surface of a gold electrode, and subsequent

immobilization of probe DNA on the surface of gold nanoparticles

through a 5’-thiol-linker [15]. The difference of electron-transfer

resistance (�Ret) was linear with the logarithm of complementary

oligonucleotides sequence concentrations in the 2.0 × 10−12 to

9.0 × 10−8 M range, and the detection limit was 6.7 × 10−13 M.

In addition, the DNA sensor showed a fairly good reproducibility

and stability during repeated regeneration and hybridization

cycles.

An amplified electrochemical impedimetric aptasensor for

thrombin has been also described [70]. A nice improvement in

the detection sensitivity was achieved by constructing a sandwich

platform where the thiolated aptamers were immobilized on a

gold substrate to capture the thrombin molecules. Then, aptamer

functionalized Au-NPs were used to amplify the impedimetric

signals (Fig. 4.11). A detection limit of 0.02 nM, with a linear range

of 0.05 to 18 nM was achieved.

Figure 4.11. Schematic illustration of sandwich amplified impedimetric

aptasensor based on functionalized Au-NPs [70] (reprinted with permission

of Elsevier).



Figure 4.12. Schematic illustration of DNA conductimetric detection

enhanced by reporter DNA–Au-NP conjugates [71] (reprinted with permis-

sion of Wiley).

On the other hand, a nanoparticle enhancement approach has

been described by Dong et al. [71] to improve the detection

sensitivity of field-effect transistors based on single-walled carbon

nanotube networks (SNFETs). Figure 4.12 shows as the target DNA

was hybridized with probe DNA on the device, reporter DNA labeled

with Au-NPs flank a segment of the target DNA sequence. The

amplified change in drain current allowed a DNA concentration

down to 100 fM to be detected.

Table 4.5 summarizes recent works which appeared in the

literature on the development of label-free genosensors employing

electrochemical impedance and conductimetry as transduction

techniques.

4.3.2.3 Signal enhancement methods

A useful approach to improve the sensitivity of genosensors

involving the use of Au-NPs consists of performing silver or

even gold precipitation onto immobilized Au-NP–DNA conjugates.

Nevertheless, the achieved improvements seem to be a compromise



Table 4.5. Label-free electrochemical impedance and conductimetric

genosensors



DNA PAT transgene GCE EIS 2.4 × 10−11 M 72

Synthetic DNA strains GCE EIS 10−12 M 73

Synthetic DNA strands SNFETs Conductance 100 fM 71

Synthetic DNA strands SiO2 Conductance 5 × 10−14 M 68

PAT gene sequence GCE EIS 3.1 × 10−13 M 21

Thrombin AuE EIS 0.02 nM 70

Synthetic DNA strands Quartz crystal Conductance 5.0 fmol 74

Synthetic DNA strands Pt Conductance 25 pmol 75

CaMV35S gene fragment CPE EIS 2.3 × 10−13 M 76

Synthetic DNA strands AuE EIS 6.7 × 10−13 M 15

PAT gene sequence CPE EIS 5.6 × 10−13 69

CaMV35S: 35S promoter from cauliflower mosaic virus; CPE: carbon paste electrode; EIS:

electrochemical impedance spectroscopy; GCE: glassy carbon electrode; PAT: phosphinothricin

acetyltransferase; SNFETs: field-effect transistors based on single-walled carbon nanotube

networks.

between the wanted signal augmentation and the reproducibility of

the assays [27]. Furthermore, strategies involving enlargement of

Au-NPs have been employed as well to enhance the electroanalytical

signal monitoring DNA hybridization events.

Silver enhancement is based on the reduction of silver ions from

one solution (usually the enhancer) by another (the initiator) in

the presence of Au-NPs [77]. The reduction reaction causes silver

to build up preferentially on the surface of the Au-NPs, giving rise

to a core-shell structure. An illustrative example is the work from

Bonanni et al. [78]. They used streptavidin-coated Au-NPs and silver

enhancement kits to amplify the impedimetric signal generated in

a biosensor detecting the DNA hybridization event. The scheme

displaying the sensor preparation procedure is shown in Fig. 4.13.

A good reproducibility was achieved (RSD lower than 8.5%), the

detection limit being 11.8 pmol.

The same group discussed recently described impedimetric

detection methods for double-tagged DNA from PCR amplification

of Salmonella spp. [77]. One of these methods involved amplification

of the impedimetric signal by using a monoclonal IgG1kappa anti-



Figure 4.13. Schematic illustration of the experimental procedure fol-

lowed to obtain DNA hybridization sensors based on stretavidin-modified

Au-NPs and silver enhanced detection [78] (reprinted with permission of

Elsevier). See also Color Insert.

digoxigenin antibody (anti-DIG) from mouse able to specifically bind

the digoxigenin-modified end of the amplicon. A secondary anti-

mouse IgG labeled with Au-NPs, able to interact with the anti-

DIG, was then added for signal amplification. A silver enhancement

treatment was carried out by depositing onto the electrode surface

20 mL of a solution obtained by the combination of equal volumes

of enhancer and initiator, and allowing 7 min for the reaction to

proceed. The amplified signal was around 30% higher than the

signal obtained without amplification.

Electrooxidation of hydrazine does not occur on DNA-conjugated

Au-NPs, although it does on bare Au-NPs. However, a chemical

treatment with NaBH4 significantly enhances the electrocatalytic

activity of DNA-conjugated Au-NPs allowing a high signal current

to be obtained for compounds such as H2O2, formic acid, glucose,

or hydrazine [79, 80]. Figure 4.14 shows a schematic view of the

electrochemical DNA detection approach using the NaBH4 enhanced

electrocatalytic activity of Au-NPs. The chemical treatment produces

the adsorption/absorption of a large amount of hydrogen species

on/into Au-NPs, thereby forming an enhanced activity state. This

enhancement process may provide very fast electron-transfer kinet-

ics for hydrazine electrooxidation on Au-NP. A high signal current

was obtained at 0.7 V, whereas the low intrinsic electrocatalytic



Figure 4.14. Schematic view of electrochemical DNA detection using the

NaBH4 enhanced electrocatalytic activity of Au-NPs [79] (reprinted with

permission of the American Chemical Society). See also Color Insert.

activity of ITO electrodes allowed low background currents to be

obtained, the contribution of the attached Au-NPs to the background

current being minute due to a low surface coverage of Au-NPs. The

high signal-to-background ratio allowed a detection limit for the

sensor of 1 fM to be achieved.

Protocols involving the enlargement of Au-NPs as a means

to achieve signal amplification can be also considered in this

subsection. Liao et al. [81] described a detection strategy for

mutated papillary thyroid carcinoma DNA based on the square

wave stripping voltammetry (SWSV) measurement of gold released

from enlarged Au-NPs. As shown in Fig. 4.15, a biotinylated 30-

nucleotides probe-DNA was immobilized in a streptavidin-modified

96-well microtiter plate. After blocking with bovine serum albumin

(BSA), the biotinylated target DNA was allowed to hybridize.

Next, streptavidin-labeled Au-NPs were added, and a nanoparticle

enlargement process was performed using a gold ion solution and

formaldehyde as a reducing agent. The enlarged Au-NPs were then

dissolved in bromide and SWSV was applied to monitor the DNA

hybridization event. The enlargement process allowed a high sen-

sitivity to be achieved with a linear semi-log plot in a DNA concen-

tration range from 0.52 to 1300 aM, and a detection limit of 0.35 aM.

Enlargement of Au-NPs has also been employed as a sig-

nal amplification system in conjunction with other amplification



Figure 4.15. Schematic display of the experimental procedure for the

determination of mutated papillary thyroid carcinoma DNA-based on the

use of enlarged Au-NPs and SWSV [81] (reprinted with permission of

Elsevier).

protocols for the development of a model thrombin aptasensor

[82]. The sensing platform is illustrated in Fig. 4.16. It consisted

of a gold electrode-aptamer/thrombin/aptamer-functionalized Au-

NP (Apt–Au-NP) sandwich design, where the detection sensitivity

was improved due to the development of three-level cascaded

impedimetric signal amplification steps. A thiolated thrombin-

aptamer was self-assembled on a gold electrode and used to

capture the analyte thrombin in sample solution, obtaining the

Au/TBA/MCH/thrombin electrode. The Apt–Au-NPs were subse-

quently bound to this electrode forming the sandwich system

mentioned above and achieving the first-level signal amplification

of the electron-transfer resistance. Then, the Apt–Au-NPs bound

to the modified electrode were used as the seeds for their

catalytic enlargement thus obtaining the second/third level signal

amplification. Due to the steric-hindrance between the enlarged

Apt–Au-NPs blocking the electron-transfer of the redox probe, the

electron-transfer resistance of the Au/TBA/MCH/thrombin/Apt–

Au-NPs electrode increased, realizing the second-level signal

amplification. In addition, the negatively charged SDS used as

stabilizer capped the enlarged Apt–Au-NPs with negative charge,

which repels the negatively charged redox probe, [Fe(CN)6]3−/4−,



Figure 4.16. Schematic outline of the label-free impedimetric biosensor

for thrombin at an aptamer-functionalized Au electrode based on a three-

level cascaded signal amplification: (a) Formation of a mixed monolayer

of thiolated aptamer and 6-mercaptohexanol on the AuE; (b) thrombin

addition and binding with aptamer; (c) binding with Apt–Au-NPs to

carry out the first-level signal amplification (I); (d) the enlargement of

the SDS-stabilized Apt–Au-NPs to achieve the second/third-level signal

amplification (II/III); and (e) schematic outline of the electron-transfer

resistance of different modified electrodes [82] (reprinted with permission

of acs).

leading to an enhancement of electron-transfer resistance and

achieving the third-level signal amplification. The label-free electro-

chemical impedimetric developed aptasensor showed a detection

range from 100 fM to 100 nM and could provide a promising

model for electrochemical impedance spectroscopy detection of

proteins.

Similarly to previous detection approaches, recent works on the

development of DNA electrochemical biosensors based on the use of

Au-NPs and signal enhancement methods to improve the sensitivity,

are summarized in Table 4.6.

March

14,201220:1

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6in04-O

zsoz-c04

SignalTransductionand

Amplification

Strategies135

Table 4.6. Gold nanoparticle-based electrochemical biosensors using signal enhancement methods

Analyte Modified Electrode Substrate Detection Technique Detection Limit Reference

Silver enhancement

Hepatitis B virus DNA sequences – PSA 0.7 ng mL−1 83

Synthetic oligonucleotides GECE EIS 11.8 pmol 78

Kitasatospora strains SiO2 chip Conductimetry 1 ng/mL 84

Oligonucleotides with transcription factor NF-κB AuE ASV 0.1 pM 85

Salmonella spp. DNA Av-GEB EIS 1 fmol 77

Enhancement by treatment with NaBH4

DNA synthetic strands ITO CV 1 fmol 86

DNA synthetic strands ITO Amperometry 1 fmol 79

Mutated papillary thyroid carcinoma DNA GCE SWSV 0.35 aM 81

Enlargement of Au-NPs

Thrombin AuE EIS 100 fM 82

ASV: anodic stripping voltammetry; AuE: gold electrode; Av-GEB: avidin bulk-modified graphite-epoxy biocomposite; CV: cyclic voltammetry; EIS: electrochemical

impedance spectroscopy; GCE: glassy carbon electrode; GECE: graphite epxy composite electrode; ITO: indium-tin oxide; NF-κB: nuclear factor-kappa B; PSA:

potentiometric stripping analysis; SWSV: square wave stripping voltammetry



4.4 Conclusions and Outlook

The work carried out in the last years show fairly well that

integration of Au-NPs with DNA detection systems allows the

development of genosensors with an improved analytical per-

formance when compared with conventional DNA sensors. Au-

NP–modified electrodes permit a remarkable enhancement of the

amount of DNA immobilized onto the electrode. The efficient DNA

immobilization achieved paves the way for the design of effective

signal transduction approaches of the hybridization event making

use of the different strategies summarized in this chapter. The

amplification routes that the use of Au-NPs facilitates, combined

with electrochemical techniques, allow the design of selective and

highly sensitive DNA sensors. However, looking at the literature, a

lack of applications of these DNA sensing devices to real samples

is observed. The extremely promising prospects on sensitivity and

stability that the gold nanoparticle-based DNA platforms provide

should be validated for solving real analytical problems in order

to demonstrate their competitiveness against conventional DNA

analyses.

Another prospect that can be easily foreseen is the use of

these sensing platforms for multiplexed purposes. Integration of

the genosensors into miniaturized (or even nano) devices involving

microfluidic systems should lead to the efficient and versatile

design of genosensing platforms capable to give multiple adequate

responses to the current analytical demands in the fields such as

the rapid detection of genetic disorders, pollution alarm systems, or

forensic analysis.

Acknowledgments

The financial support of the Spanish Ministry of Science

and Innovation (MICINN) through the projects CTQ2009-

09351, CTQ2009-12650 and DPS2008-07005-C02-01 is gratefully

acknowledged.


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Chapter 5

Nanoparticle-Induced Catalysis forElectrochemical DNA Biosensors

Marisa Maltez-da Costa,a Alfredo de la Escosura-Muniz,a

and Arben Merkocia,b

aNanobioelectronics & Biosensors Group, Institut Catala de Nanotecnologia,CIN2 (ICN-CSIC), Esfera Universitat Autonoma de Barcelona, Bellaterra,Barcelona, SpainbICREA, Barcelona, [email protected]

In this chapter, the use of nanoparticles (NPs) in catalytic

electrochemical analysis of DNA as a new detection strategy

reported in recent years is revised. The topics covered here

include labeling with nanoparticles and their subsequent signal

enhancement employed for DNA hybridization detection. Direct

sensing of nanoparticle labels as well as indirect detection routes

through electrochemical sensing of label-catalyzed reactions have

been reported. Nanofabrication of platforms used for the detection

of DNA through electrochemical signal amplification has also

been revised. Some recent examples of interesting nanoparticle-

induced catalytic methodologies applied for protein detection using

electrochemical biosensors are also given, because of their potential

interest in future applications in DNA detection.



142 Nanoparticle-Induced Catalysis for Electrochemical DNA Biosensors

5.1 Introduction

Various nanomaterials, including carbon nanotubes, nanoparticles,

nanomagnetic beads, and nanocomposites, are being used to

develop highly sensitive and robust biosensors and biosensing

systems [1] with a special emphasis on the development of

electrochemical-based (bio)sensors [2, 3] due to their simplicity and

cost efficiency.

One of the main requirements for a good performance of

a biosensor is the high sensitivity of the response. This is of

great importance when, for example, it is required to use the

biosensor in clinical diagnostics for the detection of low levels of

clinical biomarkers in human fluids [4], because in most cases the

biomarker to be detected is present in very low concentrations. The

need for biosensing systems that can detect these markers with high

sensitivity without loss of selectivity, that is, low detection limits

with high reliability and superior reproducibility, is becoming an

important challenge.

The amplified detection of biorecognition events and specifically

of DNA hybridization events stands out of the biosensing field,

because it is one of the most important objectives of the current

bioanalytical chemistry. In this context, approaching the catalytic

properties of some (bio)materials appears to be a promising way to

enhance the sensitivity of the bioassays.

Catalysts are materials that change the rate of chemical reactions

without being consumed in the process. Because of their huge

economical contribution, by lowering the costs of several processes,

they are actually one of most wanted materials and can be found in

manufacturing processes, fuel cells, combustion devices, pollution

control systems, food processing, and sensor systems. Catalysts are

generally prepared from transition metals, most of them from the

platinum group, but this fact still represents a high cost due to the

material expensiveness, and thus a reduction in used amounts would

be appreciated [5, 6].

The coupling of enzymes as biocatalytic amplifying labels is a

generated paradigm in developing bioelectronic sensing devices.

The biocatalytic generation of a redox product upon binding of

the label to the recognition event, the incorporation of redox


Introduction 143

mediators into DNA assemblies that activate bioelectrocatalytic

transformations, or the use of enzyme labels that yield an insoluble

product on electrode surfaces has been extensively used to amplify

biorecognition events. Due to the several problems associated with

these techniques and the fast development in nanotechnology,

nanoparticle-assisted signal enhancement for DNA biosensors has

been greatly developed in the last decade [7, 8, 9, 10, 11].

In electrochemical sensors, electrocatalytic procedures can be

approached in two ways, either by using an electrode that have

highly or moderately electrocatalytic properties, or by exploiting

a significant change in the electrocatalytic activity of an electrode

during the detection process. Gold and platinum are commonly

employed as highly electrocatalytic electrodes. Although these elec-

trodes allow fast electron-transfer kinetics for most electroactive

species, their background currents are high and fluctuate with the

applied potential, which may make difficult to obtain the high signal-

to-background ratios, required to achieve low detection limits. In

recent years, moderately electrocatalytic electrodes have been used

to obtain high signal-to-background ratios. Such electrodes can be

obtained by modifying a poorly electrocatalytic electrode with a low

coverage of a highly electrocatalytic material. For example, indium-

tin oxide (ITO) electrodes modified with a partial monolayer of

ferrocene, carbon nanotubes, or gold nanoparticles (Au-NPs) have

been employed [7, 11].

The actual knowledge concerning the special properties of NPs

arises from the numerous studies related to the effects of changes

in shape and size on the general properties of materials. From the

electroanalysis point of view the major features resulting from these

studies are enhancement of mass transport, high catalytic activity,

high effective surface area, and control over local microenvironment

at the electrode surface [8, 12, 13, 14].

The development of nanotechnology during the last decades has

led scientists to fabricate and analyze catalysts at the nanoscale.

These nanostructured materials are usually high-surface-area met-

als or semiconductors in the form of NPs with excellent catalytic

properties due to the high ratio of surface atoms with free valences

to the cluster of total atoms. The catalysis takes place on the

active surface sites of metal clusters in a similar mechanism as the



conventional heterogeneous catalysis [12] and in general, this is a

process that occurs at the molecular or atomic level independent

of the catalyst dimensions [6, 14]. There is a considerable amount

of research articles and interesting reviews in what concerns to

the study of nanoparticle-catalyzed reactions, but the application

of these reactions in electrochemical analysis is not so well

documented.

Employing NPs in electroanalysis can induce more sensitive

and selective sensors as well as more cost-effective and portable

systems. Their application as catalysts in electroanalytical systems

can decrease overpotentials of many important redox species,

inducing discrimination between different electroactive analytes,

and also allowing the occurrence and reversibility of some redox

reactions, which are irreversible at common modified electrodes

[15]. The catalytic effect can be explained through the enhancement

of electron transfer between the electrode surface and the species in

solution, by enhancement of mass transport or also by the NPs’ high

surface energy that allows the preferred adsorption of some species

that by this way suffer a change in their overpotentials (Fig. 5.1).

The most exploited materials in catalysis are the metals from

platinum group, but with the introduction of nanotechnology some

other elements that in bulk state did not attract a lot of attention,

either due to their lack of reactivity toward some analytes or due to

their high costs in production, are now emerging.

Figure 5.1. Schematic illustration of the processes that affect the

electrocatalytic oxidation by Au-NP when functionalized with DNA strands

(adapted from Ref. 7 with permission). See also Color Insert.


Catalysis Induced by Gold Nanoparticles 145

5.2 Catalysis Induced by Gold Nanoparticles

Gold nanoparticles (Au-NPs) and silver nanoparticles (Ag-NPs) are

of particular interest in DNA sensors and immunosensors due

to their advantageous properties, such as hydrophilicity, standard

fabrication methods, excellent biocompatibility, unique character-

istics in the conjugation with biological recognition elements, and

multiplex capacity for signal transducer. Therefore, a large number

of published methods use Au- or Ag-NPs in DNA [16, 17, 18] protein

[19] and even cell [20] electrochemical detection besides optical

detections like ICP-MS [21] or their use as ELISA enhancer [22].

Metallic gold was thought to be very stable and useless for some

catalytic systems, but by the reduction of size to the nanoscale range,

gold has been proved to be a very reactive element and it has been

extensively used in sensing and biosensing systems as a catalyst

for some interesting electroanalytical applications. For instance, a

sensitive NO sensor was developed through the modification of

a platinum microelectrode by Au-NPs in which they catalyze the

electrochemical oxidation of NO with an overpotential decrease of

about 250 mV [15]. An SO2 gas sensor was also developed using

Au-NPs to catalyze the electrochemical oxidation of SO2 when the

gas diffuses through the pores of the working electrode [23].

Based on the selective catalysis of Au-NPs, selective electro-

chemical analysis could also be achieved as, for example, in the

dopamine electrochemical detection in presence of ascorbic acid.

In this case, Au-NPs can be used as selective catalysts since their

presence induces the decreasing of ascorbic acid overpotential and

the effective separation of the oxidation potentials of ascorbic acid

and dopamine [13].

5.2.1 Electrocatalytic Activity of Gold Nanoparticle Labelson Silver Deposition

Wang et al. [24] first reported a DNA hybridization detection method

based on the precipitation of silver on Au-NP tags and subsequent

electrochemical stripping detection of the dissolved silver. The

assay employed a sandwich-like protocol with streptavidin–Au-NPs

labeling the biotinylated-breast cancer gene (BRCA1) sequences.



Figure 5.2. Schematic diagram of the silver chemical deposition on Au-

NP labels applied for the electrochemical detection of DNA hybridization.

Voltammograms correspond to DPV responses of the Au-NP–labeled

oligonucleotide probes in presence of (A) complementary, (B) single-base

mismatch, and (C) non-complementary oligonucleotides (adapted from

Ref. 25 with permission). See also Color Insert.

After the silver precipitation on the gold, the silver was dissolved

and detected at a disposable thick film carbon electrode using

potentiometric stripping. This method coupled the inherent signal

amplification of nanoparticle-promoted silver precipitation and the

stripping metal analysis with effective discrimination against non-

hybridized DNA. Cai et al. [25] reported a similar assay based on

the silver deposition onto Au-NP–labeled oligonucleotides and sub-

sequent electrochemical detection of Ag ions anchored onto Au-NPs

connected to hybrids through differential pulse voltammetry using

a glassy carbon electrode (Fig. 5.2). With this assay they obtained a

detection limit of 50 pM of complementary oligonucleotides.

Later on, Lee et al. [26] reported the electrocatalytic effect of

Au-NPs on silver electrodeposition upon ITO-based electrodes

(Fig. 5.3), in absence of pre-oxidation steps, and its successful

application to the DNA hybridization detection obtaining a signal-

to-noise ratio of 20 that presented a great improvement in relation

to their previous works under similar conditions.

5.2.2 Electrocatalytic Activity of Gold Nanoparticle Labelson Other Reactions

The specific binding of highly electrocatalytic labels to a biosensing

layer immobilized on poor electrocatalytic electrode enhances the

electrocatalytic current signal. If these labels are arranged in a

low coverage level, the induced change in background current will

be small which results in a large change in signal along with


Catalysis Induced by Gold Nanoparticles 147

Figure 5.3. Schematic representation of the electrocatalytic effect of

Au-NPs on silver electrodeposition on ITO-based electrodes applied for the

DNA hybridization detection (adapted from Ref. 26 with permission). See

also Color Insert.

low backgrounds, that is, possibility to achieve very low detection

limits. However, the conjugation of the electrocatalytic labels with

biomolecules may decrease their electrocatalytic activity, and a long

distance between the electrode and the labels may also produce an

undesired slow electron tunneling between them, even if the label

exhibits a high electrocatalytic activity itself. To overcome these

problems, it is possible to enhance the electrocatalytic activity of

labels by electrochemical, thermal, or chemical treatment. Thermal

and electrochemical treatments may damage the sensing layers

during the detection process, since, respectively, high temperatures

and extreme applied potentials are often necessary. But mild

chemical treatments can be a desirable option [11].

When Au-NP labels are present near an electrode they can act

as (electro)catalytic agents. However if the electrocatalytic reaction

is not reproducible, which jeopardizes the achievement of low

detection limits, the electrocatalytic reaction should be minimized

and the electrochemical signal should arise only from the catalytic

reaction [7]. The latest can be done by limiting the electron

transfer between nanoparticles and the electrode through the use

of nonconductive spacers like other particles, organic monolayers,

etc. [11, 27, 28].

Selvaraju et al. [11] reported the use of Au-NPs as catalytic labels

to achieve ultrasensitive DNA detection via fast catalytic reactions

involved in p-nitrophenol reduction in presence of NaBH4. In order



to minimize the electrocatalytic oxidation of NaBH4 by Au-NPs they

used magnetic beads as capture probe immobilization platforms

that acted also as spacers between Au-NPs and the ferrocene-

modified ITO electrode, achieving this effect only when the density

of Au-NPs at magnetic beads surface is low. The Au-NPs used as

labels, were modified with a monolayer of DNA detection probe,

without a significant loss in catalytic activity of Au-NPs for signal

amplification. By the conjugation of all the mentioned techniques

they achieved good signal amplification with low background

current and a detection limit of 1fM for DNA target with good target

discrimination.

Recently, Yang’s group [7] reported a novel strategy for Au-NP–

based signal enhancement by the improvement of electrocatalytic

activity of labels. The DNA layer on the Au-NPs does not significantly

limit the mass transfer of small molecules and ions such as

p-hydroquinone and Ag+, or inhibits the catalytic reduction of

p-nitrophenol. However, the distance between the electrocatalytic

Au-NP label and the ITO electrode is too long and the enhancement

by electrochemical treatment requires extremely applied potentials.

To overcome this problem, this same group applied a simple chem-

ical treatment of Au-NPs by using NaBH4 instead of electrochemical

treatment (Fig. 5.4). The results showed that NaBH4 treatment could

significantly enhance electrocatalytic activity of DNA-conjugated

Au-NPs toward the hydrazine current on the ITO electrodes, without

damaging the biosensing layers. This result, in combination with the

electrode modification with Au-NPs, allowed a high signal current

Figure 5.4. Schematic view of electrochemical DNA detection using

the enhanced electrocatalytic activity of Au-NP labels toward hydrazine

(NH2NH2) reduction on ITO electrodes (adapted from Ref. 7 with

permission). See also Color Insert.


Catalysis Induced by Platinum and Palladium Nanoparticles 149

whereas the low intrinsic electrocatalytic activity of ITO electrodes

minimized the background current. With this method, 1fM of target

DNA in an electrochemical DNA sensor was detected without the

need of target or enzymatic signal amplification [29].

5.2.3 Electrocatalytic Activity of Gold Nanoparticles Usedas Modifiers of Electrotransducer Surfaces

Another application of metal nanoparticles in electrochemical

detection of DNA is their incorporation with composites used as

electrode surface modifiers. Even though these modified electrodes

can show higher background signals than the unmodified electrodes,

the incorporation of Au-NPs can be used to promote selective

immobilization spots to well-oriented DNA detection probes.

Liu et al. [9] have recently reported the application of composites

of Au-NPs and multi-walled carbon nanotubes (Au-NPs/MWCNT)

for enhancing the electrochemical detection of DNA hybridization.

Au-NPs were deposited onto the surface of MWCNTs by one-step

reaction and then a thiolated-DNA probe was immobilized onto the

Au-NPs/MWCNTs–modified glassy carbon electrode (GCE) through

the strong gold–sulfur linkage, which could control the molecular

orientation of probe DNA. On the basis of DNA detection it was

found that the Au-NP/MWCNT composites could highly improve the

sensitivity of DNA biosensor due to their enhanced conductivity and

increased effective surface area. Furthermore, it was revealed that

selectivity and reproducibility of the DNA sensor were also excellent,

which resulted in a significant platform for the hybridization

detection of DNA.

5.3 Catalysis Induced by Platinum andPalladium Nanoparticles

5.3.1 Electrocatalytic Activity of Platinum NanoparticleLabels

Despite the high cost of this metal in the bulk state, the subsequent

saving that reducing the metal size implies placed platinum

nanoparticles (Pt-NPs) in the centre of attention of scientists due to

their ability to be used as catalyst for many industrial processes [13].



Pt-NPs are used as catalysts for electrochemical hydrogen perox-

ide (H2O2) detection, where they act as modifiers of the electrode

surface and electrocatalyze the oxidation of H2O2 observed by

a lower oxidation peak potential when compared with the bulk

platinum electrode [30]. As the H2O2 is a product of many enzymatic

reactions, this electrode has a vast potential application as an

electrochemical biosensor for many substances [15].

Pt-NPs have also been used as catalysts in gas sensors like nitric

oxide (NO) sensor making use of the electrocatalytic effect in the

oxidation of this specie [31]. In conjugation with carbon nanotubes

(CNTs) and glutaraldehyde, Pt-NPs also allowed the development of

a carbon-based electrode as a sensor for glucose, in a similar system

as one of the reported H2O2 sensors [13].

Regarding its application in DNA sensors, Polsky et al. [10]

used nucleic acid functionalized Pt-NPs as catalytic labels to

amplify the electrochemical detection of both DNA hybridization

and aptamer/protein recognition. The assay was based on the

catalytic effect of the Pt-NPs on the reduction of H2O2 to H2O, using

gold slides as electrodes. The amperometric measurement of the

electrocatalyzed reduction of H2O2 detected DNA with a LOD of

1 ×10−11 M.

N. Zhu et al. [32] reported in 2005 the use of Pt-NPs combined

with nafion-solubilized MWCNTs as electrode-surface modifiers

for fabricating sensitivity-enhanced electrochemical DNA biosensor.

The hybridization events were monitored by DPV measurements

of the intercalated daunomycin (Fig. 5.5). Due to the ability of

MWCNTs to promote electron-transfer reactions and the high

Figure 5.5. Schematic representation of the electrochemical detection of

DNA hybridization based on Pt-NPs combined with MWCNTs (adapted from

Ref. 32 with permission). See also Color Insert.


Catalysis Induced by Platinum and Palladium Nanoparticles 151

catalytic activities of Pt-NPs for chemical reactions, the sensitivity

was remarkably improved achieving a detection limit of 0.1 pM

of target DNA. The results showed that this DNA hybridization

biosensor responded more sensitively to target DNA than those

based on Pt-NPs or MWCNTs only.

5.3.2 Electrocatalytic Activity of PalladiumNanoparticle Labels

Palladium belongs to the platinum group of metals, and, due to

its similar features in terms of electrocatalytic alctivity toward

numerous redox reactions, it has been used in electrode modifica-

tion processes in several electrochemical sensors [33]. Palladium

nanoparticles (Pd-NPs) were applied in several electrochemical

biosensors. For instance, a glucose biosensor based on codeposition

of Pd-NPs and glucose oxidase onto carbon electrodes [34],

encapsulated channels for protein biosensing and the reduction

of H2O2 [35], and a DNA-template preparation of Pd-NPs onto

ITO for H2O2 reduction and ascorbic acid oxidation, has been

reported [33].

In the work reported by Chang et al. [36], Pd-NPs in combination

with MWCNTs were used to fabricate an electrochemical DNA

biosensor with enhanced sensitivity. Methylene blue (MB) was used

as hybridization indicator and a method with high sensitivity and

effective electrochemical discrimination against complementary

DNA, by coupling the large surface area and effective electron

transfer of MB redox from MWCNTs and the catalysis of the MB

redox reaction by Pd-NPs, was achieved. The Pd-NPs/MWCNTs

significantly increased the DNA hybridization signal to push down

the detection limits and facilitate potential manipulation of the

modified glassy carbon electrode. The LOD obtained was 0.12 pM

for target DNA. The catalytic activity of Pd-NPs employed in this

work is related to their ability to adsorb/release the involved

hydrogen atoms promoting the electronic transfer during the MB

redox reaction.

An ultrasensitive DNA sensor using the rapid enhancement

and electrocatalytic activity of DNA-conjugated Pd-NPs on NaBH4

hydrolysis was reported by Yang’s group [37]. After a previous



Figure 5.6. Schematic representation of the electrochemical DNA detec-

tion using the catalytic and electrocatalytic oxidation of NaBH4 on Pd-NP

labels onto ITO electrodes and the rapid enhancement of electrocatalytic

activity of DNA-conjugated Pd-NPs (adapted from Ref. 37 with permission).

See also Color Insert.

similar work with Au-NPs as electrocatalytic labels [11], they

recently reported Pd-NPs, activated by NaBH4, as ideal electrocat-

alytic labels for DNA detection (Fig. 5.6) that work even at high

pH levels with reduced incubation time. The high pH is necessary

in order to avoid the self-hydrolysis of NaBH4 at lower values,

even though the catalytic hydrolysis of NaBH4 can be slower at

this pH. The resulting sensor achieved an LOD of 10 aM (BRCA1

associated gene sequence) with a detection range of 10 orders

of magnitude, using an ITO electrode as substrate and following

the hybridization process by linear sweep voltammetry. The rapid

enhancement comes from the fast catalytic hydrolysis of NaHB4

onto Pd-NPs’ surfaces and subsequent fast hydrogen sorption into

Pd-NPs. The electrocatalytic activity of DNA-conjugated Pd-NPs

allows high currents for the electro-oxidation within the potential

windows.

5.4 Catalysis Induced by Other Nanoparticles

5.4.1 Electrocatalytic Activity of Titanium DioxideNanoparticle Labels

Metal oxides are emerging as important materials because of their

versatile properties such as high-temperature superconductivity,


Catalysis Induced by Other Nanoparticles 153

ferroelectricity, ferromagnetism, piezoelectricity, and semiconduc-

tivity [38].

Recently, nanostructured TiO2 particle (TiO2-NP) preparation

and their applications in photovoltaic studies, photocatalysis, and

environmental studies have attracted much attention mostly in the

emerging sensor technology based on nanoparticles and nanocom-

posites with chemical and biological molecules [33]. In protein-

based biosensors the efficient electrical communication between

redox proteins and solid electrode surfaces is still an important

request, and many methods have been tried in order to obtain

direct electrochemical responses of proteins embedded in surface

modifier films. An example for the latter is the work presented

by Zhou et al. [39] where the photovoltaic effect of TiO2-NPs,

induced by ultraviolet light, can greatly improve the catalytic activity

of hemoglobin as a peroxidase with increased sensitivity when

compared to the catalytic reactions in the dark, which indicates a

possible method to tune the properties of proteins for development

of photocontrolled protein-based biosensors. The method claims an

enhancement in the catalytic activity of hemoglobin, by a specific

interaction with 35 nm TiO2-NP, toward the H2O2 reduction. This

catalytic effect was not observed by other comparative experiments

with films containing nanostructured CdS or ZnO2.

The advances in hybrid nanotechnology involving nucleic acids

are mostly linked with sequence-specific nucleic acid interactions.

TiO2-oligonucleotide nanocomposites retain the intrinsic photocat-

alytic capacity of TiO2 as well as the bioactivity of the oligonucleotide

DNA; therefore, the developments in this area have been oriented

toward cellular imaging and protein or DNA sensor microarrays

[38].

Lo et al. [38] reported a nanocomposite biosensor for the

amperometric detection of H2O2 based on thionin incorporated

bilayer of DNA/nano-TiO2 film-modified electrode. Furthermore,

this system showed electrocatalytic activity toward the O2 and H2O2

reduction in physiological conditions.

A nano-TiO2 substrate in combination with Au-NP–modified

DNA probe was used by Lu et al. [40] to develop a novel

photoelectrochemical method for quantitative detection of the

linear DNA hybridization. In the detection process schematized in



Figure 5.7. Schematic illustration of the fabrication of Au-DNA probe

modified TiO2 electrode and the detection of target DNA (A) and the

photo-induced processes of electron-hole generation and charge transfer

processes (B) (adapted from Ref. 40 with permission). See also Color Insert.

Fig. 5.7, the probe immobilization and the following hybridization

induced the photocurrent change of the TiO2 electrode that was

enhanced with the Au-NP–DNA probe immobilization, and then

gradually decreased with increasing the concentration of the target

DNA. They could effectively discriminate the hybridization from un-

hybridization processes, and potentiate this system as a biosensor

to study a wide variety of biological processes.

A very recent work from Hu et al. [41] proposes a direct

electrochemical detection procedure for DNA hybridization using

the electrochemical signal changes of conductive poly(m-amino-

benzenosulfonic) acid (PABSA)/TiO2 nanosheet membranes, which

were electropolymerized by pulse potentiostatic method (see

scheme in Fig. 5.8). The polymerization efficiency is greatly

improved by the use of TiO2-NPs, and their combination with

PABSA resulted in a highly conductive composite membrane with

unique and novel nanosheet morphology (80 nm thick ramified

membrane) that provides more activation sites and enhances

the surface electron-transfer rate. Furthermore these nanosheets

presented good redox activity and electroconductivity even in

neutral environment (PBS solution of pH 7.0), and the DNA probes

could be easily covalently immobilized, so that the hybridization


Catalysis Induced by Other Nanoparticles 155

Figure 5.8. Schematic representation of the immobilization and

hybridization of DNA on the PABSA/TiO2 nanosheets (adapted from Ref. 41

with permission). See also Color Insert.

event could be monitored through impedance measurements. The

LOD obtained was 1.7 pM of target probe (CaMV35S gene sequence)

with a RSD of 4.91% (for 1.0 μM of target DNA) and the biosensor

selectivity was tested with non-complementary and double-base

mismatched sequences. Since this hybridization detection does not

require labeling of the oligonucleotide probe or target prior to the

assay, this procedure results in an advantageous method in terms of

simplicity, non-invasiveness, and low costs.

5.4.2 Electrocatalytic Activity of Osmium OxideNanoparticle Labels

Isoniazid-capped 25 nm osmium oxide nanoparticles (OsO2-NPs)

were reported by Gao and Yang [42] as successfully electrocatalytic

tags in a microRNA ultrasensitive detection system, schematized

in Fig. 5.9. The assay employs an ITO electrode with immobilized

capture probes (antisense to microRNAs for testing) and after

hybridization with periodate-treated microRNAs, the OsO2-NP tags

are brought to the electrode through a condensation reaction

between isoniazid molecules, grafted onto the nanoparticles and

the 3’-end dialdehydes of the microRNA in a hydrazine PBS buffer.

The readout oxidation potential of hydrazine was directly correlated

to the concentration of the hybridized microRNA and the assay



Figure 5.9. Schematic illustration of miRNA assay using electrocatalytic

OsO2-NPs (adapted from Ref. 42 with permission). See also Color Insert.

reported a linear relationship between current and concentration

from 0.3 to 200 pM microRNA with a measureable signal reported

for as low as 80 fM microRNA in 2.5 mL droplets following 60 min

hybridization. Successful attempts were made in the microRNA

expression analysis of HeLa cells.

Additionally the assay can easily distinguish between a single

base mismatch, with a signal detected for fully matched microRNA,

and less than 25% signal reported for mismatched microRNA.

These results are comparable to the previous electrochemical

microRNA detection by this group [43], and offer many of the same

advantages over more conventional methods, such as PCR-based

and Northern blot techniques. The use of OsO2-NPs in preference

to the electrocatalytic moieties presented previously by this group

offers additional advantages for the electrocatalytic quantification of

microRNA. These advantages include control over the choice of the

capping groups on the nanoparticle, which simplifies their ligation

to the microRNA and the improved catalytic effect on the oxidation of

the hydrazine that results in improved signal. However, the authors

do not address the efficiency and reliability of the conjugation of the

nanoparticle tags to the microRNA [43].

5.4.3 Electrocatalytic Activity of Other Nanoparticles

Other non-metal particles have also been described as possible

catalysts in electroanalytical systems [13, 33]. For example, copper

oxide nanoparticles (CuO-NPs) of 5 nm size were mixed with


Conclusions 157

graphite powder and used as catalysts for the electrochemical

detection of amikacin antibiotic oxidation, achieving a 40 times

higher current than with a bulk CuO-modified carbon paste

electrode [15]. Cu2O hollow spheres (150–220 nm sized spherical

aggregations of small Cu2O-NPs) were applied by Zhu et al. [44] in

electrochemical DNA sensing using a carbon paste electrode and

MB as the hybridization indicator. They make use of these particles

as enhancers to the ssDNA probe immobilization on the electrode

surface to obtain a sensitive detection of Hepatitis B virus DNA

sequences by differential pulse voltammetry.

More recently, Prussian blue nanoparticles (PB-NPs) were found

to catalyze the electrochemical reduction of H2O2 when immobilized

in the form of layers on ITO electrodes [14]. The application of

PB-NPs as catalytic labels for highly sensitive detection of DNA

hybridization was also reported based on their catalytic effect

toward H2O2 when embedded in polystyrene spheres and loaded

onto a gold-disk electrode [45].

Iron and iron oxide nanoparticles (Fe and Fe3O4-NPs)

also present catalytic properties in electroanalysis. Fe-NPs were

described as efficient and selective catalysts in the electrochemical

detection of H2O2 in the presence of O2 by facilitating the electron

transfer between adsorbates and the glassy carbon electrode

surface. Fe3O4-NPs were used to modify a crystalline gold electrode

for the electrochemical detection of dopamine. They showed good

catalytic activity by lowering the dopamine oxidation overpotential,

allowing the dopamine and ascorbic acid peaks to become separated

and resolvable and with even lower detection limits than the Au-NP

system referred above [13].

5.5 Conclusions

The induced catalysis by NPs is showing special interest in the DNA

biosensing technology. The application of NPs as catalysts in DNA

detection systems is related to the decrease of overpotentials of

the involved redox species including also the catalyzed reduction

of other metallic ions used in labeling-based hybridization sensing.

Although the most exploited materials in catalysis are the metals



from platinum group, with the introduction of nanotechnology

and the increasing interest for biosensing applications, gold nano-

particles, due to their facile conjugation with biological molecules,

besides other advantages, are being shown to be the most used.

Their applications as either electrocatalytic labels or modifiers

of DNA related transducers are bringing important advantages

in terms of sensitivity and detection limits in addition to other

advantages.

Ag-NPs are not so commonly used as Au-NPs but nevertheless

their catalytic properties in electrochemical detection have also

been exploited. For instance, they were reported as promoters for

electron transfer between the graphite electrode and hemoglobin

in a NO sensor system where they also act as a base to attach the

hemoglobin onto a pyrolytic graphite electrode while preserving

the hemoglobin natural conformation and therefore its reactivity

[46]. With respect to the application of silver catalytic properties

on DNA hybridization detection, the published works refer mostly

to its use in combination with Au-NPs by means of chemical or

electrochemical silver deposition onto them [29].

The catalytic properties of nanoparticles used in protein detec-

tion can also be extended to DNA analysis. For example, the

selective electrocatalytic reduction of silver ions onto the surface of

Au-NP reported by our group and applied for protein detection

can be extended to DNA analysis too [47]. The hydrogen catalysis

reaction induced by Au-NPs [48] and applied even for cancer cells

detection [20] is expected to bring advantages for DNA detection as

well.

The reported studies suggest that the use of nanoparticles as

catalysts in electroanalysis in general, and particularly in DNA

sensing is not confined to metal nanoparticles only. The conjugation

of nanoparticles with electrochemical sensing systems promises

large evolution in actual electroanalysis methods and is expected to

bring more advantages in DNA sensing overall in the development

of free PCR–DNA detection besides other applications that may

include microfluidics and lateral flow detection devices. These

works are under way at our and other laboratories. Their successful

application in DNA detection in real samples would require a

significant improvement of cost-efficiency of nanoparticle-based


References 159

detection systems, in general and those based on nanoparticle-

induced electrocatalysis, in particular.

Acknowledgments

We acknowledge funding from the MEC (Madrid) for the projects

MAT2008-03079/NAN, CSD2006-00012 “NANOBIOMED”

(Consolider-Ingenio 2010) the E.U.’s support under FP7 contract

number 246513 “NADINE” and the NATO Science for Peace and

Security Programme’s support under the project SfP 983807.

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Chapter 6

Application of Field-Effect Transistors toLabel-Free Electrical DNA BiosensorArrays

Peng Li,a Piero Migliorato,a and Pedro Estrelab

aDepartment of Engineering, University of Cambridge,Electrical Engineering Division, Cambridge CB3 0FA, United KingdombDepartment of Electronic & Electrical Engineering, University of Bath,Bath BA2 7AY, United [email protected]; [email protected]; [email protected]

Biotechnology is in great need of low-cost intelligent biochips

capable of massive parallel detection to be used in portable

instrumentation. One way this may be achieved is by exploiting

mature semiconductor technologies for the development of biosen-

sor arrays. We review here two highly promising techniques for

label-free electrical detection of DNA hybridization: potentiometric

detection and electrochemical impedance spectroscopy. Field-effect

transistor technologies can play an important role in the develop-

ment of these techniques in biosensor microarrays.

6.1 Introduction

The ability to detect biomolecular interactions is crucial in med-

ical, pharmaceutical, and biotechnological applications. The most



164 Application of Field-Effect Transistors to Label-Free Electrical DNA Biosensor Arrays

commonly employed techniques for the detection of such inter-

actions are based on optical methods, in particular fluorescence

detection of labeled biomolecules. Large arrays of 500,000 spots

per chip are currently used for high-throughput screening of DNA

sequences, where a large volume of genomic data is obtained

with a single experiment. The parallel detection of biomolecular

interactions in large microarrays is of great scientific and economic

importance. Depending on the analyte, which can be DNA, proteins,

peptides, etc., applications of microarrays include gene expression

monitoring, pharmacogenomic research and drug discovery, clinical

diagnostics, including infectious and genetic diseases, cancer diag-

nostics, and viral and bacterial identification. It is also important for

the detection of biowarfare and bioterrorism agents, and for forensic

and genetic identification. To fully exploit these opportunities,

biosensors should provide a combination of high sensitivity and

selectivity, speed, low cost, and portability.

Although a large level of success has been achieved with

fluorescent-labeled DNA microarrays, these methods are difficult to

implement in portable instrumentation, so that their use is limited

to specialized laboratories. Electrical detection of biomolecular

interactions is highly desirable due to its suitability to low-

cost portable sensors that can be used in the field by non-

specialized personnel. The use of label-free techniques has the

added advantages of reducing costs and avoiding the need for

sample pre-treatment.

Over the past few decades, effort has been devoted to exploit

semiconductor field-effect transistors (FETs) in chemical and

biological sensors due to the potential of these devices to meet some

of the requirements discussed above. Most of this work concerned

the development of the ion-sensitive field-effect transistor (ISFET)

for the detection of specific ions and analytes using appropriate

ion-selective or enzymatic membranes. One of the advantages of

the ISFET is that it operates in equilibrium conditions. Due to the

presence of the insulating layer on top of the semiconductor, no

current flows across the biological layer.

More recently, field-effect devices have been investigated for

the detection of DNA hybridization and protein interactions. It is

expected that a full understanding of the mechanisms involved will


Field-Effect Transistors 165

result in optimal device designs and create a generic platform for the

detection of any biomolecular interactions that produce a change in

the charge distribution at the surface of a transistor gate.

Besides these potentiometric-based methods, a series of electro-

chemical techniques can be applied to the detection of biomolecular

interactions. Depending on the desired dynamic detection range and

the specific properties of the system under study, techniques such

as electrochemical impedance spectroscopy, voltage step capaci-

tance measurements, amperometry, differential pulse voltammetry,

square wave voltammetry, AC voltammetry, and chronopotentiomet-

ric stripping analysis can be used for label-free detection of DNA,

proteins, and peptides [1]. Often these techniques require the use of

redox mediators. Electrochemical impedance spectroscopy (EIS), in

particular, is a very promising technique for DNA biosensing [2, 3].

Of particular interest for FET-based chemical and biological

sensors is the use of thin-film transistors (TFTs). For example, the

polycrystalline silicon (poly-Si) TFT, which can provide the drive

logic as well as the switching transistors, is a very interesting

technology for the development of low cost, disposable biosensors,

with a large number of parallel channels. By employing poly-Si TFTs,

a microarray of over 105 channels, with integrated logic drivers,

would require only a few tens of electrical connections to the rest

of the system. These could be provided by edge connectors, thereby

enabling easy insertion and removal of the sensor array from the

external reading system and, therefore, single use of a complex

microarray. Furthermore, poly-Si TFTs with a special extended

gate structure have been used as potentiometric sensors for DNA

hybridization [4]. The construction of TFT-addressed biosensor

microarrays with integral scan and readout circuits constitutes, in

our opinion, one of the great future challenges for TFT-integrated

electronics.

6.2 Field-Effect Transistors

Potentiometric chemical and biological sensors detect the electric

potential which arises at the surface of a solid material when placed

in contact with an electrolyte. Field-effect semiconductor devices



Figure 6.1. Structure of a metal–oxide–semiconductor field-effect tran-

sistor (MOSFET) and an ion-sensitive field-effect transistor (ISFET).

(a) Cross section of an n-type MOSFET. (b) An ISFET is created by

replacing the metal gate of the MOSFET by an electrolyte and a reference

electrode.

can be used as potentiometric chemical and biological sensors.

The basic structure is the metal–oxide–semiconductor field-effect

transistor (MOSFET).

A single crystal silicon based n-channel enhancement mode

MOSFET is shown in Fig. 6.1a. It consists of a p-type single crystal

silicon semiconductor substrate with two heavily doped n-type

regions (named source and drain), a gate dielectric, and a metal

gate electrode on top of the gate dielectric [5]. When the voltage VG

applied to the metal gate is lower than the threshold voltage VT, the

p–n junction between the drain and the substrate is reverse biased

and no current flows between source and drain. For VG >VT, the

electric field induced by the gate voltage is large enough to convert

the lightly doped p-type silicon substrate into n-type (inversion): an

n-type channel is created at the insulator–semiconductor interface

and current can flow between source and drain. Due to the

presence of the insulating layer, no current flows from the gate into

the semiconductor. The amplitude of the current flowing through

source and drain is modulated by the electric field set up by gate

voltage, which is determined by the charge on the metal gate

electrode.

By its working principle, the MOSFET amplifies the input signal

VG with an intrinsic gain given by the transconductance gm. In

the linear region where VG is small and in the saturation region



where VG is sufficiently large, gm is given by the following equations,

respectively:

gm = ∂ ID

∂VG

∣∣∣∣

VD=const

= WL

μC VD (6.1)

gm = ∂ ID

∂VG

∣∣∣∣

VD=const

= WL

μC (VG − VT) (6.2)

where ID is the drain current, μ the carrier mobility of the substrate

material, C the gate capacitance per unit area, W and L the width

and length of the conducting channel, respectively. Hence the ampli-

fication power of a MOSFET device is closely related to the mobility

of the semiconductor material and can be tuned by the design of the

transistor. The sensitivity of the drain current to the charge on the

gate electrode can hence be explored for sensor applications.

If the metal gate of a MOSFET is removed from the field-effect

transistor and the gate dielectric placed in contact with a liquid

solution, as shown in Fig. 6.1b, ions can adsorb on the surface of the

gate dielectric, which generates an electric field similar to applying

a voltage at the metal gate [6, 7]. When an external gate voltage is

applied through a reference electrode in the solution, the electrical

field introduced by the adsorbed ions leads to a shift on the device

characteristic. As the shift is quantitatively linked to the type and

density of the adsorbed ions, this new device is hence named an ion-

sensitive field-effect transistor. Selectivity of ISFETs can be induced

by the appropriate incorporation of certain pH-sensitive insulators

or ion-selective membrane.

Successful application of ISFETs in pH meters has generated

great interest regarding the possibility of using the well-understood

FET technology to produce amplifying devices that would respond

to larger and more complex molecules in solution or gas phase,

such as DNA, enzymes, antibodies, or antigens, or even whole

tissue layers [4, 7–11]. Numerous biosensors have been developed

based on similar principles, with a large variety of targets, gate

materials, and device structures. More recently, FETs with a metal

gate functionalized with a biological recognition layer have also been

developed [4, 7, 11].



6.2.1 Field-Effect Transistor Technologies

One of the major advantages of employing FETs in sensor

applications is their mature manufacturing technology. Due to the

development of the microelectronics industry, the microfabrication

process has been well established allowing FETs to be mass-

produced with extremely high yield. Thin layers of materials can be

deposited on large areas of substrates and patterns of the device

can be created by lithography, through customized masks which

can be reused. The cost of each device is mainly determined by

the substrate area and production volume, making it possible to

fabricate complex sensor arrays at affordable costs. This is especially

attractive for biosensor applications, as disposability is a highly

emphasized feature to avoid contaminations.

6.2.1.1 Single crystalline silicon and CMOS

Traditional FET transistors are fabricated on a single crystalline

silicon wafer of a few hundred micrometer thickness. The silicon

crystalline framework is homogenous and continuous with very low

levels of defects. The electron mobility,μ, is therefore at a high level,

ranging from few hundreds to over a thousand cm2 V−1 s−1, enabling

high performance devices to be fabricated. In addition with the

abundance of material, cost-efficiency, and well-understood device

physics, single crystalline silicon has been the most widely used

substrate material in the microelectronics industry.

Complementary metal–oxide–semiconductor (CMOS) is a sin-

gle crystalline silicon-based semiconductor fabrication technology,

which distinguishes itself from other types of fabrication technolo-

gies by providing both n-type (as shown in Fig. 6.2a) and p-type

MOSFETs on the same substrate. It has been used predominantly in

microprocessors, memories, and other digital logic circuits due to

its low power consumption and unmatched production yield. CMOS

technology is also used for a wide variety of analog circuits such as

image sensors, data converters, and transceivers.

Driven by the microelectronics industry, the CMOS fabrication

process has been continuously refined to make smaller MOSFETs,

which are both faster and more cost-efficient. The state-of-the-art



Figure 6.2. Schematic structures of (a) a single-crystal Si MOSFET,

(b) amorphous silicon TFT, and (c) polycrystalline silicon TFT.



CMOS transistor today has gate dimensions as small as 45 nm and

working frequencies up to a few GHz. On the other hand, with the

high purity substrate material and advanced fabrication process,

the yield of the CMOS process is extremely high, making it possible

to include hundreds of millions of transistors in a single device.

Although the silicon MOSFET transistor does not have the best

noise and speed performance as other semiconductor devices in the

field of electronics, the well established CMOS technology certainly

makes it an obvious choice for biosensor applications.

Despite the high performance of CMOS, its manufacturing

process requires very high-cost equipment, clean room facilities,

and expensive high purity single-crystal silicon wafers. Those

limitations have set up the barrier to further reduce the fabrication

costs and hindered the use of CMOS technology in large area

electronics such as displays.

6.2.1.2 Thin-film transistors

Besides using a CMOS process, which employs single crystalline

silicon as a substrate, FETs can also be fabricated on thin films

of semiconductors such as amorphous (α-Si) or polycrystalline

silicon. A direct benefit of these technologies is to replace expensive

single crystalline silicon wafers with cheaper insulators supporting

a thin layer of deposited semiconductor as substrate, which

substantially reduces the manufacturing costs. A thin-film transistor

is a metal–insulator–semiconductor field-effect transistor (MISFET)

fabricated on an insulating substrate by employing entirely thin-

film constituents. The total thickness of the transistor is normally

less than 1 μm [12]. There are variations in TFT design, but the

basic device structures for both amorphous silicon and polycrys-

talline silicon technologies are depicted in Figs. 6.2b and 6.2c,

respectively.

Normally TFTs are operated like enhancement-mode MOSFETs. A

typical drain current ID vs. gate voltage VGS characteristic is shown

in Fig. 6.3. When the gate voltage VGS (with respect to the source) is

low, very little current flows between the source and drain, because

of the high resistance of the active layer. When the gate voltage is

high, charge is induced near the oxide–semiconductor interface, and



Log

I D (

A)

Ion

IoffVGS

ID

VGS (V)

VDS

Figure 6.3. Typical drain current vs. gate–source voltage characteristics

for a TFT. The circuit elements are indicated in the inset. The curve is for

fixed VDS.

a conductive path (channel) is established between the source and

drain. Hence, the TFT can operate as a switch, controlled by the gate

voltage.

Despite of its much reduced manufacturing cost and versatile

form factor, the main drawback of TFTs compared with single

crystalline silicon devices is the low electrical performance. This is

a direct result of the low electron mobility of the semiconductor

material employed for TFT fabrication.

In the case of the amorphous silicon TFT, the conducting channel

is created in the amorphous silicon layer, in which the long range

order of lattice is absent and the atoms form a continuous random

network. Due to this disordered nature of the material, amorphous

silicon has a high level of defects which is normally passivated

and reduced by hydrogen to prevent anomalous electrical behavior.

Consequently the electron mobility is reduced to 1–10 cm2 V−1 s−1,

compared with a few hundred for single crystalline silicon. This

essentially ruled out amorphous silicon TFT for analog circuits and

high speed logic circuits, where high internal gain and large fan out

of transistors are required.



While amorphous silicon TFT suffers from low electronic

performance, it is very flexible in application and manufacturing.

One important advantage is that amorphous Si can be deposited

at temperatures as low as 75◦C. This makes it possible for the

device to be made not only on glass, but also on plastics. In

addition, amorphous silicon can be deposited over very large

areas by plasma-enhanced chemical vapor deposition (PECVD) with

standard industrial equipments. Both features make mass-scale

production of amorphous silicon TFT-based devices relatively easy

and economic. The main application for amorphous silicon TFT is

on liquid crystal displays (LCDs), in which each pixel is individually

driven by a TFT transistor.

Polycrystalline silicon is a material consisting of multiple small

silicon crystals with sizes ranging from nanometers to micrometers,

widely used as a gate material of FET and interconnection in

integrated circuits. Depending on the size of the crystals or

grains, the electron mobility in polycrystalline silicon lies between

that of amorphous and crystalline silicon, ranging from 10 to

100 cm2 V−1 s−1, and providing device performance good enough for

electronic circuits. It was the ability to fabricate integrated drive cir-

cuits [13] that stimulated the initial interest in polycrystalline silicon

for active matrix displays. The technology, now well developed, has

been for long time applied in LCD displays for projectors and is now

being used for mobile phones. Poly-Si TFTs have also been employed

to make static random-access memories (SRAMs) and operational

amplifiers.

The fabrication of a polycrystalline silicon film can be achieved

through various CVD methods or crystallization of amorphous

silicon. But these processes require high temperatures of at least

300◦C, making the deposition only possible on glass but not

plastic. A relatively new technique called laser recrystallization has

been devised to crystallize a precursor amorphous silicon film by

localized heating without damaging the plastic substrate. A transfer

process has also been developed to fabricate poly-Si TFT circuits on

plastic substrates [14].

In recent years, organic or polymer semiconductor materials

have been intensively researched to make TFTs. These organic

TFTs can be manufactured with very low cost using much simpler



processes which don’t require clean room facilities, making it

suitable candidates for disposable biosensor applications. However,

as the development of those devices is still in its infancy and the

manufacturing processes have not been well established, organic

semiconductor TFTs will not be included further in this discussion.

6.2.2 Field-Effect Transistor Arrays

FETs are frequently employed on arrays in a variety of applications

like memories, displays, and sensor arrays such as charge-coupled

devices (CCDs). In these applications, FETs are used to construct cir-

cuit elements performing certain functions, which are then repeated

in a network. The nature of lithographic fabrication processes

makes FETs ideally suitable for large-scale array applications. Since

FETs are manufactured in batch mode with patterns transferred by

lithographic masks, the increased number of devices and complexity

only requires the alteration of the mask, while other manufacturing

steps essentially remain the same. Under mass production, the

fabrication cost of each array is hence determined by the area of

the substrate material consumed, and practically independent of

the number of array elements. Arrays integrating a large number

of elements require active logic addressing circuits to reduce the

number of interface connections. According to the requirements on

performance and cost, various types of FETs find their application in

different areas.

For high performance applications such as dynamic random

access memory (DRAM) and CCDs, where high density, high working

frequency, or high sensitivity is required, CMOS FETs are used for the

circuit elements. The peripheral circuits which address and read the

array cells are also built with CMOS and monolithically integrated

with the array elements to achieve high speed. Due to the relevant

high cost of CMOS process, these arrays are often highly integrated

with millions of array elements arranged on a substrate with an area

of about 1 cm2.

Liquid-crystal displays normally employ a matrix of amorphous

silicon TFTs to control the voltage applied to the individual

pixels. In order to drive an active-matrix addressed flat-panel

LCD, it is necessary to make contact to each of the row and



column connections, which typically amounts to over 2000 external

connections. However, the logic circuits driving the TFT matrix have

to be made by conventional single crystal silicon microchips, since α-

Si TFTs cannot provide logic drivers with the necessary speed, due

to the low electron mobility (<1 cm2 V−1 s−1).

Monolithic integration of logic drivers on the active matrix array

plate has the great advantage of reducing the number of electrical

connections between the array and the rest of the system, which

is of particular relevance when compact construction is a premium

to overcome space limitations. Polycrystalline silicon TFTs have a

much higher mobility (>100 cm2 V−1 s−1) than α-Si TFTs and

can therefore be used to provide the drive logic as well as the

pixel transistors. Complete integration reduces the total number of

external connections to ∼20 for power, clock, and input data signal

lines [15].

The above properties make poly-Si TFTs a very interesting

technology for the development of low-cost disposable biosensors,

with a large number of parallel channels. A microarray of 100,000

channels, with integrated logic drivers, would require only a few tens

of electrical connections to the rest of the system. These could be

provided by edge connectors thereby enabling easy insertion and

removal of the sensor array from the system and, therefore, single

use of a complex microarray.

6.3 Field-Effect DNA Sensing

Similarly to the working principle of ISFETs, the sensitivity of

FET devices to the charge on its gate electrode can be utilized to

develop sensors for the detection of charged biological species. In

general, biologically sensitive FETs (BioFETs) can be constructed

from MOSFET structures by functionalizing the gate electrode with

different biological recognition elements. A change in the charge

density of a biolayer immobilized on an electrode induces a change

in the electrode surface charge density, σ0, which in turn alters

the surface potential, ϕ0, that is, the open circuit potential (OCP).

A change in the surface potential may be generated by a catalytic

reaction product, surface polarization effects, or the change in dipole


Field-Effect DNA Sensing 175

Figure 6.4. DNA immobilization and hybridization on the gate metal of a

FET. See also Color Insert.

moments occurring with bio-affinity reactions. It can also be due

to potential changes arising from biochemical processes in living

systems, such as the action potential of nerve cells. The FET acts as

a potentiometric transducer.

In the case of DNA, the increase in negative charge in a

layer of immobilized DNA probes upon hybridization with target

oligonucleotides causes a significant change in ϕ0 (Fig. 6.4). If

immobilization is on the gate of an FET, hybridization causes a

shift in the flat-band potential, Vfb, of the semiconductor. This

causes a shift in the current–voltage (I –V ) characteristic of the FET

[4, 7, 9].

Field-effect DNA biosensors have been fabricated with very

different approaches to immobilization strategies, hybridization,

rinsing, and measurement conditions. These have had varying levels

of success, achieving different immobilization densities, hybridiza-

tion efficiencies, amount of non-specific binding, and stability. For a

high sensitivity, a large voltage shift upon hybridization is needed.

This requires a large increase in surface charge density upon

hybridization, requiring a large surface density of probes that still

allows high hybridization efficiency. To achieve a stable, high-density

probe layer resulting in high efficiency hybridization, end-tethered

covalent attachment is necessary. Many designs are based upon

functionalization of the gate dielectric of an ISFET. However, since

the pH selectivity of the gate oxide is not required, functionalization

of a gate metal is an option that allows immobilization using

thiol chemistry. This enables easy and reproducible fabrication of

high-density and highly stable mixed self-assembled monolayers of



thiolated oligonucleotides, using only a single biochemical step.

It also eliminates various problems that may occur using semi-

conductor or insulator surfaces, which are prone to uncontrolled

modifications, contaminations, or hydration. These may lead to

a change in the intrinsic properties of the insulator, such as its

dielectric constant, which are critical for the stable operation of

FETs.

Polycrystalline silicon thin film transistors have also been

employed for the detection of DNA hybridization [16]. A mixed

self-assembled monolayer of thiolated DNA probes and mercapto-

hexanol was immobilized onto the gold gate of an extended gate

poly-Si TFT. A shift of the I –V characteristics on the order of 300 mV

was obtained upon hybridization of the immobilized probe with a

fully complementary strand. The shift is independent of electrode

area, so microarrays can be constructed where a known DNA probe

is immobilized on each FET. The inherent miniaturization and com-

patibility with microfabrication technologies make the technique

highly promising for the development of low-cost portable devices.

6.3.1 Physical Mechanisms of Detection

A better understanding of the physical mechanisms involved in the

field-effect detection of DNA is fundamental in the development

of reliable DNA microarrays based on FETs. Several aspects play a

role in the detection mechanism. Counterion condensation theory

can be used to evaluate the effective charge density of the DNA

layer in contact with an electrolyte, which partly screens its charge,

its dependence on the ionic strength of the electrolyte, and the

reduction of the charge fraction observed upon hybridization.

Mathematical models have been used to describe the observed shifts

in the I –V curves of the field-effect transistors.

The immobilization of the nucleic acid probe is crucial in deter-

mining the performance of the biosensor. To achieve high sensitivity

and selectivity, the hybridization efficiency must be maximized

and the non-specific adsorption minimized. Immobilization should

produce a stable layer of well-defined probe orientation, readily

accessible to the target. There are a wide variety of immobilization

methods, depending on the transducer surface and application. For



devices with a gold metal gate, mixed self-assembled monolayers of

thiols are usually chosen since they give rise to highly organized,

stable, and reproducible films in which the surface density of

the oligonucleotides can be controlled in order to eliminate

steric hinderance effects and increase the hybridization efficiency.

To achieve fast hybridization kinetics and a high hybridization

efficiency, a probe density of ≤3 × 1012 cm−2 is required [17]. To

obtain the greatest shift in gate potential (VG) in a field-effect sensor,

there will be a trade-off between greater hybridization efficiency

and greater counterion screening of the DNA charge as the probe

density is reduced. In addition, if the DNA layer is considered as a

plane charge, the voltage shift depends non-linearly upon the charge

density through the Grahame equation, so that an increase in the

density of probes may lead to a large increase in the charge density

upon hybridization, but only a small increase in the voltage shift.

Hybridization kinetics can be promoted with a high ionic

strength buffer, with specificity achieved by washing with a low ionic

strength buffer. A low ionic strength measurement buffer is required

for field-effect sensing to give little screening of charge. However,

the stability of the DNA duplex in these low ionic strengths must

be considered. To give greater hybridization efficiency and sequence

selectivity and to increase stability at low ionic strength, PNA probes

can be utilized.

6.3.1.1 Description of the electrochemical system

When an electrolyte is in contact with an electrode, an electrochem-

ical double layer forms. In the Gouy–Chapman–Stern model of the

electrochemical double layer [18], it is assumed that the solvent

provides a continuous dielectric medium with dielectric permittivity

equal to its bulk value, that charges of discrete ions are smeared out

into a continuous distribution of net charge density, and that ion–

ion interactions can be neglected so that all ions in solution are free

to contribute to the charge density. Due to their finite size, ions may

not approach the electrode closer than the outer Helmholtz plane

(OHP). Since there is no charge between the electrode and OHP,

the electric field E is constant in this region, and the electrostatic

potential ϕ varies linearly. Outside the OHP, the potential may be



determined by considering the solution to be divided into laminae

parallel to the electrode. The laminae are in thermal equilibrium, but

at differing energies due to the potential ϕ, so the concentration ni

of species i with valence zi is related to its bulk concentration n0i by

the Boltzmann factor

ni = n0i exp(−zieϕ/kT ) (6.3)

The net charge density ρ(x) is related to the potential by the

Poisson equation

ρ(x) = εε0

d2ϕ

dx2(6.4)

where ε is the relative dielectric permittivity, ε0 is the permittivity of

free space, and x is the distance from the electrode. Use of boundary

conditions leads to the non-linear Poisson–Boltzmann equation.

For ϕ � kT /e, the linearized Poisson–Boltzmann equation results.

Alternatively, the non-linear Poisson–Boltzmann equation may be

solved for a symmetrical electrolyte that contains only one cationic

and one anionic species, both with charge magnitude z, giving the

Grahame equation for the charge per unit area on the electrode σ1:

σ1 = −εε0

dϕ

dx

∣∣∣∣

OHP

=√

8kTεε0n0 sinh|z|eϕOHP

2kT(6.5)

6.3.1.2 DNA charge fraction

dsDNA is a semi-flexible chain with persistence length ∼100 nm,

where the persistence length is the distance in which tangent

vectors decorrelate, a measure of the rigidity of a polymer. Short

duplexes can be considered as cylinders of 2.0 nm diameter and

axial length per base pair of 0.34 nm. The corresponding parameters

for ssDNA have not been established. Stacking interactions between

hydrophobic bases tend to produce a stiff single-stranded helix and

ssDNA has been modeled as a cylinder of diameter ∼1.4 nm and axial

length per monomer of 0.34 nm [19]. However, if ssDNA is assumed

to be a freely jointed chain with a length per base of 0.43 nm [20],

its persistence length varies from 5 nm at 1 mM ionic strength to 0.8

nm at 100 mM ionic strength [21]. This is consistent with a much

stronger rigidity of dsDNA compared to ssDNA.



Manning’s counterion condensation (CC) theory [22] is the

asymptotic Poisson–Boltzmann solution for straight polyelec-

trolytes of infinite length at infinite separation and zero salt

concentration. If the axial charge spacing b is less than the Bjerrum

length lB (the distance at which two unit charges have a Coulomb

interaction energy equal to the thermal energy kT ), a fraction θ of

the polyelectrolyte charge is compensated by counterions localized

to the polyelectrolyte, reducing its net charge:

θ = 1 − ξ−1 (6.6)

where ξ is the Manning parameter, i.e., the number of unit charges

per Bjerrum length, given by

ξ = lB/b = e2/εε0kT b (6.7)

If b < lB, counterion condensation occurs and the net axial

charge density of the polyelectrolyte is reduced to one charge per

Bjerrum length (equal to 0.714 nm for water at 25◦C). CC remains

valid as long as the polyelectrolyte length is greater than the Debye

screening length λD and b � λD. At greater salt concentrations,

excessive counterion condensation is expected. CC holds for helical

charge lattices, with the counterion fraction still dependent upon the

axial charge spacing [23]. For dsDNA b = 0.17 nm, giving a charge

fraction of 24%. For ssDNA b ≈ 0.43 nm, giving a charge fraction

of 60%, and the same effective charge per unit length. Due to the

reduction in length upon duplex formation, dsDNA is expected to

have a lower net charge than ssDNA. This is valid as long as b � λD;

so counterion condensation is expected to remain valid at ionic

strengths much less than 500 mM, corresponding to a Debye length

of 0.43 nm. This charge fraction value of ∼25% for dsDNA has been

confirmed experimentally [24].

Molecular dynamics solutions have shown that as the separation

of polyelectrolytes is decreased from infinite, the counterion fraction

increases slightly from the Manning limit. At low salt concentrations,

CC is qualitatively unchanged. The layer of condensed counterions

contracts, but the amount of condensation is only marginally

increased. Increasing salt concentration leads to a crossover

between Manning condensation and charge screening when the

Debye length becomes smaller than the radius of the condensed



layer [25]. However, the Poisson–Boltzmann theory fails to describe

the physical situation if the electrostatic interactions are strong,

the counterions are multivalent or the density of DNA is high [26].

Monte Carlo studies of oligonucleotides have indicated that the local

cation concentration is expected to decrease sharply as either end

of the molecule is approached, due to coulombic end effects [27].

Due to end effects and dense packing of oligoelectrolytes, counterion

condensation may not give an accurate approximation of the charge

fraction for oligomers immobilized in a SAM. Molecular dynamics

studies of single-grafted ssDNA and dsDNA oligomers show that

counterion condensation increased with both longer chain lengths

and added salt [28]. For 16 bases oligonucleotides at zero salt

concentration, 30% of counterions were contained within 1.6 nm

of the oligonucleotide for dsDNA, compared to 15% for ssDNA.

Although dsDNA has a smaller charge fraction, its net charge will

be 65% greater than for ssDNA. Addition of 5 mM salt increased the

fraction of counterions within 1.6 nm of the ssDNA to 45%. Results

on salt addition were not given for dsDNA. In the single-chain limit

studied, a significant portion of counterions lies beyond the chain

length from the surface. However, for a strong polyelectrolyte brush,

counterions are expected to be contained within the brush with

electroneutrality satisfied locally [29].

The DNA charge will also be affected by its confinement to a SAM.

The ionization of acidic or basic groups in a SAM is less favored and

for acid groups pK a will increase by approximately 1 unit [30].

6.3.1.3 Quantitation of the field-effect device signal

A variety of different approaches to calculate the shift in the

I –V characteristics upon DNA hybridization or due to charge

redistribution upon antibody–antigen binding have been presented

in the literature. It has been suggested that the accumulation of

charged molecules at a surface might be electronically detected

as responses to a Donnan potential, which is built up during the

attachment of the molecules [31]. This should only be possible if the

sensor exhibits a pH sensitivity smaller than the Nerstian response.

If this is the case, the output signal should depend on the gate

material.



The relationship �VG = −�Q DNA/C ox, where Q DNA is the

charge of the DNA SAM and C ox is the gate insulator capacitance

has been proposed in the literature [32]. This relationship is of

the same form as that of a shift in flat-band potential due to a

fixed oxide charge Q f located at the Si–SiO2 interface in a MOSFET.

However, the DNA charge is instead located at the metal–solution or

insulator–solution interface. Therefore, to return the semiconductor

to the state it would be in the absence of the DNA charge requires

charging the double-layer capacitance. Approximating the double-

layer capacitance as constant, the relationship should be �VG =−�Q DNA/C dl.

Other authors equate �VG to the change in electrochemical

double-layer surface potential resulting from the change in surface

charge, calculated using the Grahame equation [33]. The solution

and semiconductor are coupled by the electric field in the oxide,

ESiO2. If VG is adjusted to operate the FET at constant current, ESiO2

and the potential drop across the semiconductor and oxide remain

constant, and the only changes in the system occur in the double

layer. Therefore, �VG is equal to the change in potential across

the double layer, and no consideration of semiconductor physics is

necessary [7].

If the biomolecular probe is immobilized onto a metal electrode,

such as the metal gate of a MOSFET, a contact can be made to this

electrode and the open-circuit potential EOC measured against the

reference electrode. Since VG = ϕsolid state − EOC, where ϕsolid state is

the constant potential difference between the FET source and the

solid–solution interface, EOC corresponds to the shift of the I –Vcharacteristics from those measured by direct connection between

the gate and source or back contact. Therefore, the FET is simply

being used to measure the change in open-circuit potential, taking

advantage of its high-input impedance, low-output impedance, and

small size.

A one-dimensional model for electrolyte–insulator–metal–

oxide–semiconductor and electrolyte–insulator–semiconductor

structures modified with a charged membrane has been presented

[34]. It was shown that the largest sensitivity occurs at low

electrolyte concentrations, and that the signal from hybridization

is expected to be smaller than that from probe immobilization,



assuming a doubling of the membrane charge density. Proto-

nation/deprotonation of surface sites significantly reduces the

magnitude of the variation of the surface potential with respect

to the bulk electrolyte by effectively pinning the insulator surface

potential. At 10 mM salt concentration, the difference in potential

upon hybridization saturates with increasing probe density, so

increasing the probe density above 1 × 1012 cm−2 is not expected

to further increase the shift upon hybridization. At these high probe

densities, a −19 mV shift is calculated for full hybridization on an

uncharged surface, 6 times greater than the −3 mV change with

an amphoteric Al2O3 surface. At salt concentrations of 10 mM or

greater, where the thickness of the charged layer is significantly

greater than the Debye length, for uncharged surfaces the Donnan

potential was shown to give a good approximation of the double-

layer potential.

Finite element modeling of DNA functionalized electrodes was

applied to calculate the interfacial potential, and used to identify

conditions for maximum potential change with target hybridization

[35]. Using different models such as the Donnan potential model

[34] and numerical solution of the Poisson–Boltzmann equation

for a three-dimensional model, the authors estimate a maximum

potential variation of −17 mV for 100% hybridization efficiency at

the optimized DNA probe density of 3 × 1012 cm−2 even at low ionic

strength.

Even though larger shifts have been reported in the literature,

the simulations give a good insight on the variation of the signal

with probe density and ionic strength. The signal decreases rapidly

at probe densities lower than 1 × 1012 cm−2, while increasing the

probe density above the optimal value has little effect due to the

reduction of hybridization efficiency. Decreasing the ionic strength

on the other hand, has little effect on the signal at high probe

densities but increases the signal at low probe densities.

The value of the interfacial potential with ssDNA is significantly

larger than the change in potential resulting from hybridization.

In addition, decreasing the ionic strength significantly increases

the potential but not the variation in potential upon hybridization.

If uncharged PNA probes are used instead of DNA probes, the

interfacial potential before hybridization is expected to be approx-

imately zero, independent of ionic strength. Therefore, significantly



greater potential changes with hybridization are expected, and these

changes are enhanced by the use of low ionic strengths. PNA probes

also have the advantage of a PNA–DNA duplex stability that is

approximately independent of ionic strength. A much larger value

for the interfacial potential change of −100 mV has been calculated

for a PNA probe density of 2 × 1012 cm−2 at low ionic strengths

[35], suggesting that PNA probes are likely to provide reliable

potentiometric DNA sensors with low limits of detection.

6.3.1.4 Equivalent electrical circuit model of functionalized FET

The impedance of a FET with the gate immersed in solution

and potential applied to a reference electrode in solution may be

represented by the equivalent circuit shown in Fig. 6.5. The circuit

consists of the silicon resistance RSi, space-charge capacitance C SC,

oxide capacitance C ox of the FET, and the Randles equivalent circuit

for the double layer, where Z W has been omitted since there are

no redox molecules in solution. In the absence of redox molecules,

Rct is large and Z imag can be considered to result from the series

combination of the three capacitances.

When the biomolecular interaction happens at the solid–solution

interface, it changes the value of C dl. At fixed applied potential,

this would introduce charge redistribution between C dl and C ox,

where the change of potential across C ox depends on the ratio of

the two capacitors, C dl/C ox. The value of this ratio is fixed when the

biomolecular probe is immobilized directly on the gate dielectric or

on the gate electrode directly on top of the dielectric. In an extended

gate structure, a sensing pad is electrically connected to the gate

electrode. The area of the sensing pad can be much larger than

Figure 6.5. Equivalent circuit for a field-effect device with gate immersed

in solution.



that of the transistor. In this configuration, the ratio C dl/C ox can

be largely improved by increasing the double-layer area, offering a

larger voltage shift for the measurement [4].

6.3.2 Differential OCP Measurement

Taking advantage of its high-input impedance, low-output

impedance, and miniaturization, the metal-gate FET is being used

to measure variations in the open-circuit potential that occur

upon interaction. Recently, direct OCP measurements using an

instrumentation amplifier have been performed resulting in reliable

detection of protein interactions [36].

The open-circuit potential was measured in real time by using an

ultra-low input bias current instrumentation amplifier, providing an

accurate differential measurement of voltage. The very high input

impedance and very low input bias current minimize the effect of

the measurement on the OCP. The gain of the amplifier was set to 1

in order to eliminate instability effects, temperature drift, etc., of the

external resistor needed to set a higher amplifier gain.

The functionalized gold electrode and the reference electrode

were connected to the amplifier differential inputs (see Fig. 3.6).

The amplifier output voltage, equal to the open-circuit potential for

V0INA116

G=1

RE

Au

–9V

+9V

Figure 6.6. Schematic instrumentation amplifier set-up for the open

circuit potential measurement (RE, reference electrode).


Electrochemical Impedance Spectroscopy 185

unit gain, was recorded using a potentiostat. The amplifier output

reference terminal was grounded to ensure good common-mode

rejection.

The simplest on-chip circuit that can be conceived is for differ-

ential OCP measurement, although complex electrostatic discharge

protection needs to be incorporated.

6.4 Electrochemical Impedance Spectroscopy

Many electrochemical biosensors rely on the reduction and oxi-

dation (redox) processes that occur at a functionalized electrode.

These sensors are engineered so that a biomolecular interaction

induces a change in the redox current. These amperometric

techniques rely on the measurement of output currents upon a

voltage-driven electrochemical event. The measurement of elec-

trochemical currents requires the use of a potentiostat with a

three-electrode cell arrangement since a current flowing through

the reference electrode creates an electrochemical reaction at its

surface and, consequently, alters the applied potential. The voltage

is applied through a reference electrode connected to a high-

impedance input of the potentiostat so that no current flows

through it, and the current is measured with the help of a counter

electrode.

Many standard electrochemical techniques can be used, depend-

ing on the biological system to be studied. In the presence of redox

markers in solution, modification of the electrode resulting from

biomolecular interaction affects the impedance of the system, which

can be measured by using electrochemical impedance spectroscopy

(EIS). EIS is a very promising technique, in particular for the

detection of DNA hybridization.

In EIS, the impedance of the system is measured by applying a

small ac signal and by the frequency scanned (typically between 10–

100 kHz and 1 Hz or less). Stable impedance spectra can be obtained

with electrically charged redox markers in solution. The data can be

fitted with an equivalent electrical circuit, where the most important

components are the charge transfer resistance Rct and the double

layer/biolayer capacitance C dl.



The charge of a biolayer immobilized onto an electrode will

create an electrostatic barrier to, e.g., the negatively charged

[Fe(CN)6]3−/4− redox couple in solution, which is reflected in the

value of the charge transfer resistance. Upon interaction, the charge

distribution of the biolayer will change, causing a modification in the

electrostatic barrier and, therefore, in the value of Rct. An increase in

Rct can be related to an increase in negative charge or a decrease in

positive charge at the biolayer. Reverse charge changes will cause a

decrease in Rct. Another important factor to take into account when

interpreting charge-transfer resistance changes is the fact that some

areas on the Au surface which are accessible to the redox couple, will

be blocked upon the biomolecular interaction due to the relatively

large volumes of target molecules, such as proteins. This effect will

result in an increased Rct.

On the other hand, a change in capacitance is expected upon

biomolecular interactions. When a large target biomolecule inter-

acts with the immobilized probe, the biolayer thickness increases,

causing a decrease in the total capacitance of the system.

In the case of DNA, hybridization at the electrode results in

a significant increase in the negative charge of the DNA layer.

Therefore, the electrostatic barrier to the negatively charged redox

couple becomes stronger upon hybridization, causing an increase

in the charge transfer resistance. A typical Nyquist plot (−Z imag vs.

Z real) is shown in Fig. 6.7 for a Au electrode after immobilization

of single-stranded DNA probe and after hybridization with its

complementary strand. The charge transfer resistance corresponds

to the diameter of the semi-circle in the Nyquist plot. For the sample

in Fig. 6.7, a 5 k� increase in Rct is observed upon hybridization [37].

The technique is robust and large signal discrimination upon

hybridization can be obtained with optimization of the DNA

probe density and the measurement conditions. Keighley et al.[37] report on the optimization of co-immobilization of thio-

lated oligonucleotides and mercaptohexanol to form mixed self-

assembled monolayers on gold. Specifying the solution mole ratio of

the thiol components provides an effective and easily implemented

method to accurately control the oligonucleotide surface density.

A linear relationship between mole ratio and probe density was

observed for the range (1.3–9.1) × 1012 probes/cm2. With this



Figure 6.7. Electrochemical impedance spectroscopy characteristics for a

Au electrode with ssDNA and after hybridization.

method the sample-to-sample variability was reduced as compared

to previously reported immobilization methods. The ratio on the

surface was approximately equal to that in the solution only for DNA

mole fractions lower than 0.3%.

Electrostatic repulsion between the immobilized negatively

charged oligonucleotide probes and negatively charged ferri/

ferrocyanide redox couple in solution results in a modulation of the

charge transfer resistance with probe surface density. The increase

in negative charge at the sensor surface upon hybridization only

results in a modulation of charge transfer resistance at probe

densities above 2.5 × 1012 cm−2 [37]. This threshold is probably due

to counterion screening of the oligonucleotide charge resulting in

channels between probes through which the mass transport of the

ferri/ferrocyanide redox couple is unaffected.



The maximum percentage change of charge transfer resistance

upon hybridization with fully complementary target oligonu-

cleotides was obtained with samples prepared by co-immobilization

of oligonucleotide probes and mercaptohexanol with a DNA mole

fraction of 20%. This corresponds to a mean probe surface density

of 5.4 × 1012 cm−2.

6.4.1 PNA-Based Sensing

The electrostatic barrier to the negatively charged redox markers

changes upon DNA hybridization, causing the EIS signal. The use

of PNA probes yields much larger EIS signals upon hybridization.

Since PNA is uncharged, the potential barrier before hybridization

is negligible resulting in a very low charge transfer resistance value;

upon hybridization with the charged DNA target, the potential

barrier is strongly felt resulting in a particularly large variation of

Rct.

Optimization of PNA surface density resulted in a massive

enhancement of the fractional change in Rct upon hybridization,

without the use of additional biochemical amplification steps [38].

A fractional change 100-fold larger than previously reported has

been achieved. Another relevant aspect is that the optimization

of PNA surface density in a mixed PNA/MCH SAM results in a

small initial Rct, controlled by the mercaptohexanol regions of

the SAM. For a given electrode area and overpotential, a smaller

Rct gives a greater current density. For a given sensitivity of the

detection electronics, higher current densities enable a reduction

of the minimum sensing electrode area and therefore an improved

detection limit. A detection limit of 25 fmol target was demonstrated

by Keighley et al. [38]. This is likely to be further improved by

reduction of the electrode area and sample volume. For example,

reducing the electrode diameter from 2 mm to 100 μm, a 400-fold

decrease in area, would increase the initial Rct to around 4 M�. A

10 mV AC overpotential would result in an approximately 2.5 nA

AC current, feasibly measured in a portable detection system. This

would allow the sample volume to be scaled to 3 nl, reducing the

detection limit to 3 amol. This shows electrochemical impedance



spectroscopy with PNA probes to be a very promising technique for

portable DNA detection applications.

6.4.2 Modeling of the Signal

The optimization of the EIS signal for DNA sensing can be achieved

through modeling of the DNA layer potential and charge changes

upon hybridization. To consider the effect of discrete charge sites,

a geometry model was composed to represent the DNA structure

at the surface [35, 39]. The ssDNA probe or DNA/DNA duplex

(or PNA/DNA duplex when PNA probe is used) was modeled as

a cylinder with diameter of 2 nm, perpendicular to the electrode

surface and linked by a spacer, as shown in Fig. 6.8a. The negative

charge of the phosphate backbone was considered as a uniform

surface charge evenly distributed on the side of the cylinder. The

DNA strand was spaced from the surface by the linker molecules—

in the case presented, 2.7 nm long to represent a linker consisting

of 6 polyethylene glycol (PEG) groups. As this distance is longer

than the Debye length of solutions with ionic strength above 15 mM,

the effect of the metal electrode on the electric field around DNA

strand can be neglected. When the spacer molecule is uncharged, the

electric field is not affected by the SAM and a symmetry plane can be

Figure 6.8. Geometry model for the simulation of modification layer with

discrete charged sites: (a) side view of the structure, (b) cross section

showing the simulation plane with dimensions representing probe density

of 3×1012 molecules/cm2.



found at half of the height of the DNA probe, as shown with the dash

line in Fig. 6.8a. Since the length of ssDNA probe is normally much

longer than that of its width, as an approximation the model can be

simplified into a two-dimensional simulation on this plane. Since

the top and bottom sides of the cylinder do not carry any charge

in the model, the electric field on the simulation plane is hence the

largest potential which determines the charge transfer resistance.

Figure 6.8b shows the view on the symmetry plane for simulation.

The probe was assumed to be arranged in a homogeneous hexagonal

lattice with a center-to-center spacing determined by the probe

density.

Similar to that of a uniformly charged layer, the Poisson–

Boltzmann equation was solved numerically in two dimensions

within the domain surrounded by the dash line in Fig. 6.8b. A

typical result with a probe density of 3×10 12 molecules/cm2 and

measurement ionic strength of 50 mM is shown in Fig. 6.9 [39]. Upon

bonding of DNA target, the increase of the charge density further

enhances the electric field around the hybridization site, resulting

in a change of Rct, which can be measured as the sensing signal of

target hybridization.

Detection of DNA target can also be achieved with ssPNA as the

probe. Using DNA or PNA as sensing probe presents two different

situations for the change of charge upon the hybridization. For DNA

probe, the ssDNA itself carries charge before the target binding

and the target hybridization increases the surface charge. While for

a PNA probe, target binding converts an uncharged surface to a

charged surface. The signal range of the sensor is defined by the

signal measured with the probe fully hybridized by the target and

the signal of the un-hybridized probe.

From simulation results, using PNA probe yields larger signal

range for all probe densities [39]. The difference is more pronounced

with high probe densities, when the hybridization signal with PNA

probe can be over 10 times larger than that of DNA probe measured

with the designated ionic strength.

As shown in Fig. 6.9, the electric field generated by the

charged probes extends laterally, until screened by the supporting

electrolyte. From the Debye theory, the screening length is also

a function of the ionic strength. For the same modified surface,



Surface: Electric potential Contour: Electric potential

–6 –4 –2 0 2 4 6

× 10–9

6

4

2

0

–2

–4

–6

× 10–9

Min: –0.124 Min: –0.117

–0.02

–0.03

–0.04

–0.05

–0.06

–0.07

–0.08

–0.09

–0.1

–0.11

–0.12

Max: –0.0195 Max: –0.022

–0.027

–0.032

–0.037

–0.042

–0.047

–0.052

–0.057

–0.062

–0.067

–0.072

–0.077

–0.082

–0.087

–0.092

–0.097

–0.102

–0.107

–0.112

–0.117

Figure 6.9. Simulated electric potential produced by DNA probe immo-

bilized with the mixed SAM structure with an uncharged spacer. The DNA

probe density is set to be 3×1012 molecules/ cm2 and ionic strength 50 mM.

decreasing the ionic strength of the measurement solution can result

in a larger and more extended potential field, which leads to a larger

impedance signal.

The signal, defined as the ratio Rct(duplex)/Rct(probe), is

estimated to have very different ranges for the situation where DNA

or PNA are used as probes: with PNA the signal increases drastically

upon hybridization from 1.1 to 2 × 106 when the ionic strength is

reduced from 1000 to 1 mM; under the same conditions, using a DNA

probe only yields an increase from 1.05 to 3.5. The signal range using

DNA probe saturates when the ionic strength is lower than 10 mM,

since Rct of both ssDNA and dsDNA increases with similar amplitude.

For the PNA probe, as the probe itself is not charged, decrease of

ionic strength always gives increased signal range. When the ionic

strength is sufficiently low (∼50 mM for the probe density studied),



a linear relationship is observed between the signal range and ionic

strength on the log scale. These results further support that PNA is

better as sensor probe than DNA for EIS hybridization detection.

6.5 Application of FETs on Biosensor Arrays

Application of biosensors in areas such as pathogen identification

and gene expression requires a large number of sensor elements

to work simultaneously in an array format. As an example, current

fluorescence-based optical DNA microarrays for genotyping and

gene expression often involve ∼500,000 spots, where an individual

probe is deposited at each sensor element [40, 41]. High level of

integration and performance is clearly required in those devices.

The development of optical biosensor arrays is limited by the optical

scanner’s high cost, the unreliability of the optical labeling process,

and the complex data processing procedures. Considerable efforts

have been devoted to the development of alternative biosensor

array platforms suitable for low-cost production and higher level of

integration.

Fully integrated label-free electronic biosensor arrays based on

well-established microfabrication methods are believed to be able

to adequately address the disadvantages of optical arrays. Label-free

electrochemical characterization techniques can be implemented

directly using integrated electronics, achieving significant cost

reduction and better system integration. These electronic biosensor

arrays can be easily connected to simple handheld readers for point-

of-care applications. Moreover, studies have shown that electrical

stimulation can significantly affect the kinetics of biomolecular

interaction at solid–liquid interface [42–44], which is easily achieved

with electronic biosensor arrays.

6.5.1 FET-Addressed Biosensor Arrays

Besides the biomolecular probes and the packaging components, the

electronic components of a fully integrated biosensor array can be

divided into three categories: transducer, array addressing circuit,

and measurement unit. As the most widely used microelectronic

devices, FETs play an important role in all these three categories.


Application of FETs on Biosensor Arrays 193

As previously introduced, the FET itself can be used as a

potentiometric biosensor transducer to translate biomolecular

interactions into variations of flat-band voltage or source–drain

current. The ability of miniaturization makes FETs ideal candidates

for applications on biosensing arrays, as the signal-to-noise ratio

is independent of the geometry size. This feature allows FET-

based biosensors to be integrated on extremely high-density arrays,

with the limit of detection determined by the immobilization of

biomolecular probes and practical operations. The performance of

an FET-based potentiometric transducer depends on the internal

gain of the FET, which is measured by the transconductance, and on

the fabrication geometry, which determines the ratio of the double

layer capacitance to the gate dielectric capacitance.

Independently of the electrochemical technique employed for

DNA sensing, FETs can have an important role in the development

of electronically addressed biosensor arrays. Acting as switches for

individual cell elements, there are two basic requirements for the

addressing circuits. First, the electronic switch attached to each

sensor must have a high on/off current ratio. This is to make sure

that when the designated sensor element is measured, interference

from other sensor elements does not affect the characterization.

The second requirement is that the logic circuit, which translates

the input signal into the address information and selects the sensor

element, must work at a high enough frequency. As the biomolecular

reaction is often a dynamic process when the measurement is

carried out, all the sensor elements need to be characterized in a

relatively short time window, typically a few seconds. The driver

logic circuit needs to switch on all the sensor elements sequentially

within this time window to allow the measurements.

Two possible architectures for TFT-addressed biosensor arrays

are illustrated in Figs. 6.10 and 6.11. For potentiometric sensing,

the biosensing pad is connected to the gate of the TFT (see

Fig. 6.10), which acts as the transducer. A dummy transistor, where

no biomolecular interaction occurs, can be used for differential

measurements [45]. For current detection, the sensing pad needs

to be connected to the source or the drain of the TFT as shown in

Fig. 6.11.

As the biomolecular interaction delivers a very weak electronic

signal, integrated amplification and noise canceling are often



Figure 6.10. TFT switching matrix for potentiometric detection. Each

cell is composed of a sensing transistor (S), a reference dummy transistor

(D), and switch transistors for the sensor (ST-S) and the dummy (ST-D).

needed, which requires build-in reference and measurement circuits

on the same chip. FETs are ideal to build various analog circuits

including differential input, voltage reference, operational ampli-

fiers, and potentiostat circuit. Requirements on the performance

of those circuits largely depend on the type of application and

specification of sensor arrays.

6.5.2 Specifications of the Biosensor Arrays

Although the ultimate performance of any integrated biosensor

is limited by the properties of the affinity-based biomolecular

interaction, the method of detection and fabrication impose certain



Figure 6.11. TFT switching matrix for amperometric and EIS detection.

Each sensing electrode is connected to a transistor. A potentiostat is

required for current detection.

requirements on the transducers, the electronic circuits, and the

manufacturing process. A systematic analysis of those requirements

is necessary in the development of integrated electronic biosensor

arrays and to identify the suitable technology to use.

The first consideration for a miniaturized biosensor array is the

size and surface topology of the sensor element. The preparation

of biomolecular probes on the surface of each individual sensor

element involves manipulating very small volumes of sample, which

is normally achieved by the use of microspotters. Most advanced

robotic-based liquid dispensing spotters nowadays have resolutions



down to tens of micrometers, which sets the minimum size of the

sensing area of each individual sensor element, and consequently

the density of the sensor array [40].

The transducing methods also limit the size of sensor elements.

For potentiometric detection, although the signal-to-noise ratio is

independent of the FET dimension, having a larger extended gate

structure can significantly improve the sensitivity of the transducer

[4]. According to the requirements on the sensitivity, it is preferred

to have a ratio between the areas of the extended gate and the

FET gate at 10–100, which results in individual sensor dimensions

of 10–100 μm based on modern microfabrication technologies.

For amperometric detection methods such as EIS, reducing the

size of sensor electrodes leads to a decrease of the current to be

measured. The resistance of the biomolecular layer varies within a

large range—typically between 30 k�cm2 and 5 M�cm2. A reliable

measurement of sub-pA current requires very high performance

electronic devices and complicated circuit design. Therefore, the

typical dimension of amperometric sensors based on ac methods

cannot be smaller than tens of micrometers.

Electronic biosensor transducers, either potentiometric or

amperometric/EIS-based, also require atomically flat surfaces or at

least surfaces with controlled roughness. The underlying consider-

ation is the density of immobilized biomolecular probes and hence

the target captured by the probes in the biomolecular interaction.

It has been shown that the immobilization density depends on the

microscopic area of the sensor surface, which is determined by both

the geometry area and the roughness factor [39]. For measurement

techniques where the amount of charge is of concern, such as

potentiometric detection or chronocoulometric detection, a uniform

surface with regular roughness factor is needed for the entire sensor

array.

Another important consideration is the working frequency. To

characterize the sheer number of sensors in the same array in

real time, both the switching circuit and measurement units need

to work at high enough frequency. The speed needed eventually

depends on the nature and kinetics of the biomolecular interaction

to be measured. For example, considering a typical array with 1,000

elements to be measured in 1 second, the logic circuit to address the



sensor array needs to work at frequencies of ∼104 Hz. Moreover, if

an ac method such as EIS is used as the characterization technique,

all the sensor elements need to be measured in a single period. The

frequency requirement will then be the basic addressing frequency

multiplied by the highest frequency used in EIS, which raises the bar

to around 107 Hz in practice.

Other factors that need to be taken into consideration include

temperature variations induced by the power consumption of the

circuit, lifetime in solution which is determined by the passivation

material, and overall chip size and packaging for practical handling.

Those factors are less important in terms of the use of FET and need

to be reviewed for each application.

6.5.3 Development of Biosensor Arrays Based on FETs

Due to their advantages over conventional optical arrays, electronic

biosensor transducers and arrays have attracted intensive research

interest in recent years. The vast majority of those efforts, however,

are focused on the use of FETs as transducers and only a few

groups have successfully prototyped their array devices [46–54].

The obvious reason for this is the high cost of mask making and chip

fabrication. Among them, the CMOS process dominates due to the

easy access to commercial CMOS foundries.

A configurable electrochemical sensor microarray system-on-a-

chip fabricated in a standard CMOS process has been presented

in the literature [48, 49]. The array had 5 × 10 elements, each

occupying an area of 160 μm × 120 μm and containing a differential

electrochemical transducer with a programmable sensor. The sensor

had a digitally configurable topology capable of performing different

electroanalytical measurements including voltammetry and field-

effect sensing.

In another report, a DNA sensor array of 16 × 8 sensor elements

with pitch size of 250 μm has been fabricated using a 0.5−μm

CMOS process [50, 51, 55]. The DNA hybridization is measured

through the change of interfacial capacitance and then converted to a

digital output signal by the integrated electronics. The chip was post-

processed with a gold layer to facilitate the attachment of probe DNA



using the thiol-gold chemistry. Successful discrimination between

complementary and mismatched target DNA was demonstrated.

Other examples of CMOS-based field-effect sensor arrays are

used for the monitoring of extracellular electrophysiological signals

or pH changes [56]. Those devices normally involve only ISFETs and

hence are less complicated in structure than those required for DNA

or protein arrays.

AC techniques such as EIS, which demand high electronic

performance, have also been shown on CMOS circuits integrated

with biosensor arrays. Unlike with potentiometric techniques,

the signal of current sensing techniques such as amperometry

and EIS naturally decreases with the electrode size. In addition,

EIS detection requires currents to be measured for a range of

frequencies, which could make the time needed to read the signals

from the entire microarray impractically long. The use of a wide

band stimulus coupled with a fast Fourier transform algorithm has

been proposed to overcome this problem [57]. A saving feature is

that the frequency range to be measured is below 100 Hz.

The aforementioned examples with CMOS processes achieved

success to a certain level, either in the electronic performance or

in measurements with actual biological samples. However, due to

the different fabrication factors, applications and characterization

methods, it is impractical to compare the performance of the

biosensor arrays.

The use of TFTs in electronic biosensor array is still limited.

Various TFT-based DNA and protein transducers have been devel-

oped either with poly-Si or amorphous silicon TFTs [4, 58]. Although

proved to be successful as sensor transducers, working TFT-based

sensor arrays have not reported in the literature. This is largely

attributed to the fact that TFT foundries are mostly specialized

for the manufacturing of LCD backplanes and not commercially

available to researchers.

6.5.4 Fabrication Technologies and Future Trends

From the point of view of electronic biosensor array appli-

cations, both CMOS and TFT technologies clearly have advan-

tages and disadvantages. CMOS represents the state-of-the-art for



microfabrication and can provide devices at tens of nanometers

working at GHz frequency ranges. However, for biosensor arrays this

high performance is far over-specification as previously discussed.

On the other hand, considering manufacturing and convenience of

operation, both the sensor element itself and the whole chip cannot

be made too small. In the existing examples, the size of the chip is

4 mm2 for 50 sensor spots using a 0.18-μm process and 20 mm2

for 128 sensor spots using a 0.5 μm process. Even with this small

number of sensors in the array, using CMOS technology leads to

substantial cost on the manufacturing of the chip, typically a few

dollars in these two cases, excluding the costs of the biomolecular

probes, post-processing, design, and installation fee for the masks.

This cost is mainly due to the expensive single crystal silicon

substrate, and hence would scale up when a larger number of sensor

elements or a larger area for each sensor element is needed. As

disposability is highly desired for biosensor arrays, the high cost of

CMOS process makes it impractical for large scale applications such

as diagnostics and disease screening.

On the other hand, although TFTs cannot provide such high

performance electronic devices, it can be manufactured on much

cheaper substrates such as glass and even plastics, making the

technology an ideal candidate for biosensor arrays in the view

of cost. The main limitation of TFTs is the low mobility of the

semiconductor material. This does not only affect the performance

when it is used as a transducer, but, in case of amorphous TFT, it also

prevents its use for the addressing logic and measurement circuit.

To be used as the addressing matrix switches for individual

sensor elements, the on/off state current ratio is the parameter to be

considered. For a biosensor array with thousands of sensor elements

the off-state resistance must be at least 3 orders of magnitude larger

that of the on-state to secure precise measurement of data. This can

be easily achieved by the use of a single FET based on either CMOS

[5], poly-Si TFT [15], or amorphous-Si TFT [59].

For the logic driving circuit, the TFT needs to work at 104 Hz with

normal sequential measurement, and 107 Hz if time multiplexed EIS

is to be implemented. The highest working frequency of a FET is

mainly determined by the mobility of the semiconductor material,

as well as by its geometry size and fabrication process. It has been



estimated that due to its low mobility, amorphous silicon TFT would

not work at more than 104 or 105 Hz, which rules it out for the use

of addressing logic circuits in future high density arrays [15]. Poly-

Si TFT, however, has a mobility of over 100 cm2 V−1 s−1 and can

adequately cover the desired frequency range.

For the measurement circuit, the amorphous silicon TFT has not

been considered suitable for analog circuit or high frequency digital

circuit, due to its low mobility and transconductance, while the poly-

Si TFT has been developed into a large variety of analog circuits with

moderate performance. The suitability of the three FET technologies

for biosensor array applications is summarized in Table 6.1.

Overall, polysilicon TFTs can provide all the key components, so

the application to the proposed integrated biosensor arrays is within

the capabilities of the technology. Furthermore, it seems to provide

the proper balance between the performance and cost for future

biosensor array applications, although its current development is

hindered by the lack of commercial foundries for research purposes.

Table 6.1. Advantages and disadvantages of CMOS, poly-Si TFT, and amor-

phous TFT technologies for the development of the different components in

biosensor arrays

Application in biosensor arrays

Addressing Measurement

Transducer switches Driving logic circuit

CMOS FET Pros High internal High speed, High speed High electronic

gain, smaller size high on-off performance,

current ratio compact in size

Cons Expensive to have None None None

larger extended

gate or electrodes

Poly-Si TFT Pros High internal gain High on-off Moderate Moderate

ratio speed performance

Cons Device uniformity None None None

Amorphous Pros Low cost, low Moderate None None

TFT temperature on-off ratio

manufacturing

Cons Very low gain, None Low speed Low electronic

device uniformity performance


Conclusions 201

6.6 Conclusions

In conclusion, we have reviewed two highly promising techniques

for label-free biosensor technology. Potentiometric detection offers

the advantage of a simple electrode arrangement, since only two

electrodes are needed. Furthermore, the signal is independent

of the electrode area, which facilitates scaling. Signal readout

and conditioning is straightforward, owing to the in-built cell

amplification.

Noticeable progress has been made in recent years in the appli-

cation of electrochemical impedance spectroscopy to biosensors.

Compared to potentiometric detection, it requires a more complex

electrode arrangement (three electrodes) and a more demanding

detection circuit (potentiostat). In addition, the signal decreases

with the electrode area and the measurements are taken over a

range of frequencies. It is likely that both techniques are used in

the future for different applications. For instance, potentiometric

detection is particularly suitable for real-time detection, while EIS

offers information for both charged and uncharged species.

For both types of techniques, FET technology can provide the

switching matrix and the integrated measurement circuitry. Three

FET technologies including CMOS, poly-Si TFT, and amorphous TFT

have been reviewed and discussed for their use in future disposable

electronic biosensor arrays. Both technical and economic aspects

have been covered to evaluate the future application of these

technologies. Although current research is predominantly focused

on CMOS-based arrays, poly-Si seems to present the best balance

between performance and cost for real-world applications. The

implementation of an all poly-Si FET microarray appears to be

within the capability of the technology. However, the non-scalability

of the EIS technique and the long data acquisition time pose

considerable challenges for the designer and the technologist.

Acknowledgments

The authors would like to thank Dr. S. D. Keighley (Cambridge

University) for help with the experiments and valuable discussions.



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Chapter 7

Electrochemical Detection of BasepairMismatches in DNA Films

Piotr Michal Diakowski, Mohtashim Shamsi,and Heinz-Bernhard KraatzDepartment of Chemistry, University of Toronto at Scarborough, 1265 Military Trail,Toronto, Ontario, M1C 1A4 [email protected]

In recent years, interest in the development of electrochemical

strategies for the detection of basepair mismatches in DNA has

increased dramatically. Electrochemistry-based methods present

a promising alternative for optical detection schemes, and are

attractive because they offer the potential for high speed, high

sensitivity and high throughput detection of mismatches at a

minimal cost. Moreover, electrochemical sensors offer tremendous

advantages in terms of ease of integration and miniaturization,

especially in comparison to their optical counterparts. In this chap-

ter, we provide an overview over recent electrochemical mismatch

detection strategies and summarize the state of the art in this field.

We begin our discussion with the preparation of surfaces and the

immobilization of a capture strand and continue with an overview

of detection strategies that exploit the direct electrochemistry of

nucleobases, the conductive properties of DNA or use hybridization

indicators, intercalators and groove binders. Methods employing



206 Electrochemical Detection of Basepair Mismatches in DNA Films

synthetic DNA analogues such as peptide nucleic acids (PNA) are

also discussed. Finally, protein and enzyme mediated biosensors,

nanoparticle based sensors, metal ion amplified sensors and a range

of miscellaneous methods is discussed.

7.1 Introduction

The determination of nucleic acid sequences for analytical purposes

has remained a strong research focus for years. Effective and

efficient high-throughput technologies are needed to screen for

genetic defects, identify organisms, and forensic applications. At

present, fluorescence-based techniques are the most commonly

employed. However, wide spread applications of such methods

is limited by low speed, high cost, size, number of incubations

steps, and the need to chemically label the DNA target. In addition,

such systems are far from being foolproof and in some cases false

positives or negatives are observed, making the data interpretation

difficult. Also, integration of the entire optical system into single

portable device is not simple and requires sophisticated fabrica-

tion processes. In contrast, an electrochemistry-based approach

is promising for point-of-care applications and on-site testing

using portable analyzers. What makes such approach attractive

are its inherent advantages of high speed, low cost, simple

instrumentation, and ease of miniaturization of the biosensing

components.

In recent years, numerous electrochemical DNA detection and

sensing methods have been described in the literature. Most

electrochemical detection schemes involve the immobilization of

an oligonucleotide (ODN) onto a transducer surface. Upon the

hybridization of the complementary target sequence to the capture

strand, the binding event is detected in form of an electrochemical

signal. Methods making use of ODN labeling and label-free methods

have been reported. Labels include the use of redox-modified

oligonucleotides, electroactive DNA intercalators, enzymes, metal

complexes and nanoparticles. On the other hand, label-free

approaches are reported that exploit the intrinsic electroactivity of

the DNA bases (guanine and adenine) or monitor changes in the

interfacial properties of the sensing surface, such as changes in the


Surface Immobilization 207

capacitance or electron transfer resistance of the film as a function

of hybridization. The latter method is highly versatile and highly

sensitive to the presence of mismatches as will be shown in this

chapter. Synthetic oligonucleotides are often used as the capture

probe. They are readily obtained in high purity and at low cost and

the base sequence can be adjusted to suit a particular target. Peptide

nucleic acids (PNA) can also be used as a capture probe. PNA has

a higher stability and improved binding affinity in comparison to

nucleic acids, but is costly.

Often, as in the case of redox-labeled oligonucleotides, the

covalent attachment of a redox label such as a ferrocene (Fc) group

is achieved by imine and amide formation using Fc-carboxaldehyde

or Fc-carboxylic acid, respectively, and also by Sonogashira coupling

with the corresponding Fc-alkyne derivative [1, 2]. Redox labels

can be introduced either on the monomer stage, by a metal-

catalyzed reaction, or after assembly of the oligomer sequence

[3]. For example, Fc-conjugated nucleotides can be conveniently

used as building blocks in automated oligonucleotide synthesis [4].

Similarly, the Fc group can be introduced after ODN synthesis by

amide coupling of Fc-COOH to a 5’-amino group of a synthetic

ODN, as was reported by Ihara and co-workers [5]. However, the

introduction of a Fc-label into a DNA oligomer can decrease the

stability of the duplex. And the position of the Fc group, the nature

of its linkage to the ODN, and the nucleobase will all influence the

“melting point” of the duplex. The interested reader is referred to a

review [6], where different Fc oligomers are discussed.

In order to be useful for the detection of nucleotide basepair

mismatches, the electrochemical signature of the mismatched ds-

ODN must be significantly different from that of the fully hybridized

ds-ODN. In this chapter, we summarize the state of the art in this

field and provide an overview over capture strand immobilization

strategies and various mismatch detections schemes.

7.2 Surface Immobilization

To design a functional DNA biosensor DNA, capture strands have to

be immobilized on an electrode surface. Thin film formation is often

accomplished by covalent attachment, adsorption or affinity binding



Figure 7.1. Schematic representation of different DNA immobilization

strategies: (a) covalent attachment of ODN thiols or disulfides by self-

assembly onto gold surfaces resulting in Au-S bond formation; (b) imm-

obilization by adsorption relies on electrostatic interactions between

negatively charged sugar-phosphate backbone of DNA and positively

charged electrode surface, and/or the interaction with the nucleobases; (c)

affinity binding of biotinylated oligonucleotides onto streptavidin modified

electrode surfaces.

(Fig. 7.1). The immobilization is essential for the development of

a robust biosensing interface and maintaining control over the

immobilization step is necessary to ensure proper orientation,

accessibility, and stability of the capture strands on the sensor

surface. We begin our review with an overview of immobilization

strategies that have been successfully employed in DNA biosensors.

7.2.1 Covalent Attachment

A number of covalent immobilization methods have been reported.

Among them, the self-assembly of thiol or disulfide containing ODNs

onto a gold surfaces is probably the most popular immobilization

strategy, as shown in Fig. 7.1. Thiols react with Au resulting in

the formation of a gold-thiol linkage as indicated in the following

equation: R-SH + Au → R-S-Au + e + H+.

For example, Mirkin has demonstrated that Fc-ODN films

attached through a gold-thiol linkage display reversible redox

behavior [7]. In addition, the surface coverage of a DNA probe can be

controlled using alkylthiol diluents, as shown in Fig. 7.2. The surface

coverage of a ss-DNA capture strand has a dramatic effect on the

hybridization efficiency since sufficient space between the capture


Surface Immobilization 209

Figure 7.2. (a) DNA film formed by self-assembly of thiol containing ODN

onto Au surfaces. (b) DNA film formed by self-assembly followed by a

dilution step with an alkylthiol diluents.

probes is required to control repulsion of the targets strands and

the steric effects between the probe strands on the surface [8].

Functionalized surfaces can also be used for covalent attachment

of modified DNA strands. For instance, DNA molecules were cova-

lently immobilized onto carbon paste electrode surfaces that was

activated using a carbodiimide (1-[3-(dimethylamino)-propyl]-3-

ethylcarbodiimide hydrochloride) and N -hydroxysulfosuccinimide

[9]. In another covalent attachment strategy, individual DNA strands

were attached to a carbon nanotube (CNT) layer supported on a gold

surface. Again, amide coupling between the carboxylic acid groups

on the CNTs and the 5’-amino group of DNA resulted in the formation

of a stable amide linkage and the resulting conjugate proved stable

to the electrochemical experiment [10].

7.2.2 Adsorption

Adsorption is the simplest method of immobilization as it does not

involve the formation of covalent bond formation between the ODN

and the surface (see Fig. 7.1b). Instead, it relies on electrostatic

interactions between negatively charged sugar-phosphate skeleton

of DNA and positively charged electrode surface and/or interactions

involving the nucleobases and the surface. Physical adsorption is

often achieved on electrochemical oxidized carbon electrodes [11–

14] (HOPGE, GCE, CPE) and less often on gold [15] or ITO [16]

surfaces. For examples, cationic polymers, such as chitosan, have



been successfully used to modify carbon electrodes [17]. Also, a

positive potential bias can be applied to the electrode to improve

the adsorption of the DNA. However, the main disadvantage of this

approach is the need for a strong affinity of the DNA to the surface,

which results in a multipoint attachment of the capture strand, thus

affecting the hybridization efficiency as the probe is restricted by

multipoint immobilization. In some cases, problems are associated

with the stability of the films.

7.2.3 Affinity Binding

Strong interactions between avidin and biotin can be exploited

in the preparation of useful sensing surfaces [18]. The stability

of the avidin-biotin binding is on par with that of covalent

attachment. Typically, avidin (or streptavidin) is first immobilized

on the transducer surface followed by binding of the biotinylated

oligonucleotides (see Fig. 7.1c) [19]. For instance, avidin can be

covalently bound to gold [20] or physically adsorbed on gold [21] or

carbon electrodes [22]. In one of the examples, avidin was adsorbed

onto a silica surface before immobilizing a biotinylated molecular

beacon (MB) [23]. Alternatively, biotin can be immobilized on the

surface followed by avidin binding allowing for further attachment

of biotinylated DNA probes. In one of the examples, polypyrrole

(PPy) was formed on the electrode, and the biotin units attached to

the film were used as anchoring points for the avidin immobilization

providing three binding sites on the avidin [24].

7.3 Detection Strategies

Numerous electrochemical strategies have been developed for the

detection of mismatches in DNA. These vary from the use of electro-

active DNA intercalators to enzymatic signal amplification schemes,

or redox-modified oligonucleotides. In the following sections, we

will focus on the discussion of a range of electrochemical mismatch

detection schemes.


Detection Strategies 211

7.3.1 Direct DNA Electrochemistry

Native electrochemical properties of DNA were first described by

Palecek [25]. The oxidations of either guanine and adenine are

irreversible multistep processes (see Fig. 7.3) [26–28].

Unfortunately, the oxidation of nucleobases is not desirable

under normal circumstances as it often results in the formation

of reactive species that lead to DNA decomposition. For instance,

guanine can be electrochemically oxidized [29], but practical

application of guanine oxidation as detection method is limited

to the use of G-free capture strands. Nevertheless, despite high

oxidation potentials and irreversibility of the oxidation process

several interesting mismatch detection schemes based on direct

nuclobase electrochemistry are worth mentioning. For instance,

Napier et al. [30] demonstrated the detection of the hybridization

of products of the polymerase chain reaction using electron transfer

from guanine to a transition-metal complex. The hybridization assay

involved recording of cyclic voltammograms of [Ru(bpy)3]2+ (bpy,

Figure 7.3. Differential pulse voltammetry of guanine ( 5th scan) at

pH 4.5 in 0.2 M acetate buffer at a glassy carbon microelectrode. (a’)

0.5 mM guanine; (b’) 50 μM guanine (... 1st scan; −−− 2nd scan, after

transferring the microelectrode to supporting electrolyte) at a scan rate

of 5 mV/s. Differential pulse voltammetry of adenine ( ) at pH 4.5 in 0.2

M acetate buffer at a glassy carbon microelectrode. (a”) 1 mM adenine;

(b”) 10 μM adenine (... 1st scan, after transferring the microelectrode

to supporting electrolyte) at a scan rate of 5 mV/s. Reprinted from

Bioelectrochemistry, 55, A. M. Oliveira-Brett, V. Diculescu and J. A. P. Piedade,

Electrochemical oxidation mechanism of guanine and adenine using a glassy

carbon microelectrode, 61–62, 2002, with permission from Elsevier.



2,2’-bipyridine) in the presence of an unhybridized probe strand

containing only A, T, and C. Upon hybridization to a complement

that contained seven guanines, a high catalytic current was observed

due to the oxidation of guanine by [Ru(bpy)3]3+. The metal complex

acting as a mediator was activated at potentials accessible in the

neutral aqueous solutions. This sensor design was tested in a

model system, which showed a charge increase of 35 ± 5 μC for

complementary strand and only 8 ± 5 μC for non-complementary

DNA strand. Furthermore, for PCR-amplified genomic DNA from

herpes simplex virus type II, 35–65 μC and 2–10 μC increase in

charge was observed for complementary and non-complementary

DNA respectively. Another interesting application of direct guanine

and adenine oxidation for mismatch detection was reported by Wei

and coworkers [31]. Catalytic guanine and adenine oxidation was

achieved using tris(2,20-bipyridyl)ruthenium(II) modified glassy

carbon (GC) electrodes, resulting in DNA detection by electro-

chemiluminescence (ECL). Interestingly, the modified GC electrodes

were prepared by casting a CNT/Nafion/Ru(bpy)32+ composite

film on the electrode surface. The method allowed for sensitive

single-base mismatch detection of the p53 gene sequence segment

(3.93 × 10−10 mol/L) by employing cyclic voltammetry stimulation.

Consequently, the observed ECL signal for a C/A mismatched ODN

was 1.5 times higher than that of the fully matched ODN.

7.3.2 Charge Transduction Through DNA

A different approach to sensing DNA mismatches takes advantage

of the distinctive electronic properties of DNA and potential

differences that exist between fully matched and mismatched ODNs.

Long range charge transport facilitated by the DNA π -stack is often

exploited in various DNA mismatch detection schemes as it has

been shown to be dependent on the presence of mismatches in

the double-strand. Furthermore, DNA mediated reactions weakly

depend on distance but are extremely sensitive to perturbations

in the base stack. Single-base mismatches induce only small

changes in the duplex structure/stability, but they create significant

perturbations in the electronic structure of the base-pair stack [32–

34]. Detection schemes based on the charge transport through DNA



often involve incorporation of electro-active centers into the ODNs.

For example, it was demonstrated by Barton [32] that double-

helical DNA films on gold surface display a marked sensitivity

to the presence of base mismatches within the immobilized

duplexes. Moreover, it has been observed that mismatch detection

is possible regardless of DNA sequence composition and mismatch

identity. The presence of mismatches was elucidated based on

the electrochemical characteristics of the redox active intercalators

bound to the DNA-modified gold surfaces. Coupled redox reactions

were employed to induce an electro-catalytic current and thus

increase the method’s sensitivity (Fig. 7.4).

The effect of intervening mismatches on long-range charge

transport through DNA was comprehensively studied by Bhat-

tacharya et al. [35]. It was established that DNA mediates charge

transport and the resulting oxidative damage are extremely sensitive

to the presence of intervening mismatches. A series of DNA

oligonucleotides that incorporate a ruthenium intercalator linked

covalently to the 5’ terminus of one strand and containing two

Figure 7.4. (a) Schematic representation of electrocatalytic reduction of

[Fe(CN)6]3− by methylene blue (MB) at a DNA-modified electrode. LB+

is leucomethylene blue, the product of the electrochemical reduction.

(b) Cyclic voltammetry at a gold electrode modified with DNA of 2 mM

[Fe(CN)6]3− (curve 1), 2 μM MB (curve 2), and 2 mM [Fe(CN)6]3− and 2 μM

MB (curve 3). Reproduced from S. O. Kelly, E. M. Boon, J. K. Barton, N. M.

Jackson, and M. G. Hill, Single-base mismatch detection based on charge

transduction through DNA, Nucleic Acids Research, 1999, 27(24), 4830–

4837, by permission of Oxford University Press.



5’-GG3’ sites in the complementary strand were employed in this

study. Single base mismatches were introduced between the two

guanine doublet steps, and the efficiency of charge transport

through the mismatches was determined through measurements of

the ratio of oxidative damage at the guanine doublets distal versus

proximal to the intercalated ruthenium oxidant. The damage ratio

of oxidation at the distal versus proximal site for the duplexes

containing different mismatches varied in the following order GC

∼ GG ∼ GT ∼ GA > AA > CC ∼ TT ∼ CA ∼ CT. The authors

suggested that that this ordering may be ascribed in part to local

changes in helical stability. However, these changes cannot be easily

explained through an increased solvent accessibility associated

with a mismatch. Marques et al. [36] demonstrated methodology

based on perturbation of the double helix π -stack introduced

by a mismatched nucleotide. In this investigation CYP3A4*1B

oligonucleotides were immobilized on the surface of a gold electrode

and hybridized with fully complementary oligonucleotide sequences

as well as with mismatched sequences corresponding to the

CYP3A4*1A reference sequence. The methodology developed could

identify CYP3A4*1A homozygotes by the 5 μC charge attenuation

observed when compared with DNA samples containing at least one

CYP3A4*1B allele. In another investigation, Boal et al. [37] employed

the DNA-modified gold electrodes to monitor the electrocatalytic

reduction of DNA-bound methylene blue for a wide range of

base analogues and DNA damage products. It was found that

the efficiency of DNA-mediated charge transfer is independent of

the thermodynamic stability of the helix. However, modifications

to the hydrogen bonding interface in a given Watson-Crick base

pair and added steric bulk yielded a substantial loss in charge

transfer efficiency. Base structure modifications that induce base

conformational changes and those that bury hydrophilic groups

within the DNA helix also appeared to attenuate charge transfer

in DNA. Addition and subtraction of methyl groups that do not

interfere with the H-bonding interactions of the bases did not appear

to have any significant effect on the CT efficiency. Importantly, the

system was capable of detecting base pair mismatches and base

damage products. Inouye et al. [38] reported an electrochemical

DNA sensor for the detection of single-nucleotide polymorphism.



Figure 7.5. Electrochemical discrimination of single-nucleotide mismatch

with Fc-ODN: (a) probe hybridized to its complementary strand, (b) probe

hybridized to single-nucleotide mismatched strand, and (c) uncorrected

SWV profiles at the gold working electrodes modified with two different

fully matched duplexes (curves 1 and 3) and mismatched duplex (curve

2). Reproduced by permission from M. Inouye, R. Ikeda, M. Takase, and T.

Tsuri, J. Chiba, Proc. Natl. Acad. Sci. U.S.A., 2005, 102, 11606. Copyright 2005

National Academy of Sciences, U.S.A.

A π -conjugated Fc-modified nucleoside analogue was connected at

the 5’ end of single-stranded oligonucleotide. After hybridization to

the complementary strand, the 3’ end of the probe DNA strand was

attached to gold electrode by Au-thiol chemistry, Fig. 7.5.

Consequently the electrochemistry of the Fc marker can be

observed, allowing for the detection of complementary DNA. The

presence of a single-nucleotide mismatch in the duplex causes,

presumably, a blockage of the conduction pathway through the base

stack at the position of the base-pair mismatch. These results in a

dramatic reduction of the electrochemical response, see Fig 7.5c. In

addition, a comparison of different DNA probes containing an iso-

meric Fc-diamidopyridine conjugate for electrochemical mismatch

detection was carried out by the same authors in a separate study

[39]. It was concluded that despite different stereochemistries of

the Fc label, all conjugated DNA probes were capable of providing

satisfactory electrochemical response for mismatch discrimination.

In another study, anthraquinone monosulfonic acid (AQMS) was

employed as an electroactive intercalator allowing to differentiate

between a complementary target DNA sequence and one containing

either C-A or G-A single mismatches [40]. The electrochemistry



resulting from electron transfer through the DNA to intercalated

AQMS is readily distinguished from that of AQMS on the electrode

surface. The difference in the chemical environment between

free and intercalated AQMS greatly affects its reduction potential,

allowing monitoring of DNA hybridization in real time. In another

study, Gorodetsky et al. [41] utilized DNA duplexes functionalized

with pyrene to fabricate DNA-modified electrodes on highly oriented

pyrolytic graphite (HOPG). The reduction of DNA-bound intercala-

tors was observed as a consequence of a DNA-mediated reaction.

The reduction of the intercalator was attenuated in the presence of

the single-base mismatches, CA and GT, independent of the sequence

composition of the ODN. Sensitivity to single-base mismatches

is enhanced when methylene blue reduction is coupled in an

electrocatalytic cycle with ferricyanide. Furthermore, utilization of

HOPG as electrode material allowed authors to investigate the

electrochemistry of previously inaccessible metallointercalators,

[Ru(bpy)2dppz]2+ and [Os(phen)2dppz]2+, at the DNA-modified

HOPG surface. It was shown that HOPG presents a suitable and

reproducible surface for electrochemical DNA sensors exploiting the

charge transport properties of DNA. Again, Gorodetsky et al. shown

that DNA-mediated electrochemistry can promote reactions at a

distance on the DNA sugar-phosphate backbone [42]. It was pointed

out that relative current densities for DNA-mediated disulfide

reductions of 1.8 μA/cm2 differed significantly from that for well

stacked intercalator reduction of about 80 μA/cm2.

7.3.3 Hybridization Indicators, Intercalators and GrooveBinders

Various molecules are capable to bind to the DNA duplex or to single-

stranded DNA. The application of DNA binding molecules for the

detection of base-pair mismatches is discussed below. For instance,

Millan et al. demonstrated sequence-selective electrochemical DNA

sensing using hybridization indicators [9]. In this detection scheme,

DNA capture strands were covalently immobilized on a glassy

carbon 14 electrode and [Co(bpy)3]3+ and [Co(phen)3]3+ served

as hybridization indicators that display reversible redox behavior.

Presumably, electrostatic interactions with the negatively charged



phosphate backbone allows pre-concentration of the complex in the

double-stranded DNA layer at the electrode surface and enables

detection of the hybridization event voltammetrically. In another

example, Barton examined a number of intercalators and groove

binders (see Scheme 7.1) as probes for the detection of mismatches

within DNA films [32].

Scheme 7.1. Chemical structures of the intercalators: [Ir(bpy)(phen)

(phi)]3+, daunomycin (DM), methylene blue (MB); and groove binders:

[Ru(NH3)5Cl]2+ and [Fe(CN)6]4−. S.O. Kelly, E.M. Boon, J.K. Barton, N.M.

Jackson, M.G. Hill, Single-base mismatch detection based on charge

transduction through DNA, Nucleic Acids Research, 1999, Vol. 27, No. 24,

4830–4837, by permission of Oxford University Press.

It was found that probes that intercalate into the DNA base stack

appear to be necessary for mismatch detection. In contrast, probes

that associate with DNA purely through electrostatic interactions do

not yield measurable differences in the electrochemical response

in the presence of base mismatches. The signals obtained from

the intercalators DM, MB and [Ir(bpy)(phen)(phi)]3+ are affected

by the presence of a mismatch. However, the response for groove

binding agent was found almost identical for fully matched and

mismatched films. It is possible that the reduction of the ruthenium

complex (Scheme 7.1) proceeds through the facilitated diffusion

of the complex along the double helix, while the intercalated

species participate in electron transfer mediated by the stacked

bases. Experimental evidence indicates that the bulkier intercalators

exhibited smaller CA/TA charge ratios. Nevertheless, the detection

of base mismatches was accomplished using direct electrochemistry

of molecules bound to DNA films. Subsequently, Yamashita and

coworkers employed ferrocenyl naphthalene diimide (FND) as



redox active intercalator to detect presence of mismatches in 20-mer

(sequence of the lac Z gene) double-stranded ODNs immobilized on

gold electrodes [43]. The FND concentrates at the sensor/solution

interface upon formation of the double strand giving rise to electro-

chemical signal proportional to amount of DNA target. FND does not

bind to the vicinity of mismatched bases resulting in lower current in

the presence of a mismatch. Different mismatches were detected by

differential pulse voltammetric measurements in this study. Another

group reported the detection of hybridization using [Co(byp)3]3+ as

redox active intercalator by cyclic voltammetry measurements [44].

The sensing interface was prepared on gold colloid modified glassy

carbon electrode. The study involved a thorough optimization of

the experimental conditions, including the preparation of the ODN

probes, the hybridization with targets, and of the electrochemical

conditions. The investigation showed that an electrochemical signal

was observed only in the presence of ds-DNA and that 5, 3 and 1 base

mismatches could be clearly discriminated from a fully matched ds-

DNA film. In another report, Kara et al. covalently immobilized 22-

mer single stranded ODN capture probes related to both HSV Type I

and Type II sequences on pencil graphite electrodes [45]. The extent

of hybridization between probe and target sequences obtained from

PCR was determined by DPV in the presence of Meldola Blue (MDB)

as hybridization indicator. Interactions between MDB and the DNA

at the electrode surface resulted in a significantly lower signal

in the case of a 4-base mismatch sequences than in the case of

fully matched sequence. Again, MB was employed by Ostana and

coworkers to electrochemically screen DNA for lesions caused by

de-amination of nucleobases [46]. The damaged DNA was modeled

by 18-mer ODNs containing a different number of mismatched

target bases (uracil instead of cytosine). It was shown that the

amplitude of the reduction signal corresponding to ferricyanide ions

considerably increases in the presence of MB. This electrocatalytic

effect allowed the detection of changes in electrochemical properties

of DNA caused by dUd dG mismatches. Using differential pulse

voltammetry and cyclic voltammetry, the authors showed that the

electron transport from the electrode through the double-stranded

DNA to MB and then to ferricyanide ions is suppressed by the

presence of mismatches in the ODN sequence. MB was also used



by Gorodetsky et al. who utilized duplex DNA functionalized with

pyrene to fabricate DNA-modified electrodes on highly oriented

pyrolytic graphite (HOPG) [41]. As expected, the reduction of the

intercalator was attenuated in the presence of the single-base

mismatches, CA and GT, independent of the sequence composition

of the oligonucleotide. Furthermore, the extended potential range

afforded by the HOPG surface has allowed the authors to investigate

the electrochemistry of previously inaccessible metallointercalators,

[Ru(bpy)2dppz]2+ and [Os(phen)2dppz]2+, at the DNA-modified

HOPG surface. These results support the application of DNA-

modified HOPG as a convenient and reproducible surface for elec-

trochemical DNA sensors using DNA-mediated charge transport. MB

was also used in practical sensor design utilizing a CeO2/chitosan

composite matrix to increase the loading of the ss-DNA probe and

to enhance the biosensor’s response performance [47]. The use

of an interesting ruthenium complex as a sensitive and selective

electrochemical indicator in DNA sensing was reported by Garcia

et al. [48] The ruthenium complex, Ru(NH3)5-[3-(2-phenanthren-9-

yl-vinyl)-pyridine] generated in situ incorporates dual functionali-

ties. The Ru center provides a redox probe and the ligand provides

a fluorescent tag. The presence of the aromatic groups in the

ligand endows the complex with an intercalative character and

makes it able to bind to ds-DNA more efficiently than to ss-DNA.

Combination of spectroscopic and electrochemical studies indicated

fundamentally intercalative interactions between the complex and

ds-DNA. The ligand-based fluorescence allows the characterization

of the complex formation and monitoring of duplex melting.

The metal-based redox center is employed as an electrochemical

indicator to detect the hybridization event in a DNA biosensor.

The sensing surface was prepared by incubation a Au electrode

with a thiolated ss-DNA based on a short DNA sequence from

Helicobacter pylori. With the use of this approach, complementary

target sequences of H.pylori can be quantified with a detection limit

of 92 pmol. In addition, this approach allows the detection of not

only a single mismatch but also its position in a specific sequence

of H. pylori, due to the selective interaction of this bifunctional

ruthenium complex with ds-DNA. A new electroactive intercalator,

Cd(II)-morin, (Scheme 7.2) was reported by Niu et al. [49].



Scheme 7.2. Formula of Cd (C15H9O7)2·2H2O. Reprinted from Bioelectro-

chemistry, 73, S. Niu, M. Wu S., S. Bi, S. Zhang, Reaction of Cd(II)–Morin

with dsDNA for biosensing of ssDNA oligomers with complementary, GCE-

immobilized ssDNA, 64–69, 2008, with permission from Elsevier.

Its interaction with salmon sperm ds-DNA was investigated using

electrochemical methods. The binding stoichiometry (m = 1.76) and

equilibrium dissociation constant K = 2.5×10−5 M were evaluated

according to the Hill model for cooperative binding. Moreover,

Cd(morin)2 was used as an indicator that allowed selective detection

of the target ss-DNA fragment. The target ss-DNA was quantified

over a linear range from 2.69 × 10−8 M to 9.16 × 10−7 M with a

detection limit of 9.30 × 10−9 M. In another report, interactions

of promethazine hydrochloride (PZH) with films prepared from

thiolated ss-DNA and ds-DNA on gold electrodes were studied by

Wei et al. [50]. The binding of PZH to the ss-DNA film is purely

based on an electrostatic interaction. However, the interaction of

the probe with the ds-DNA film is a combination of electrostatics

combined with intercalation into the duplex. The latter results in

an increased peak current for PZH oxidation and a larger electron

transfer coefficient and a faster standard rate constant.

The use of [Cu(dmp)(H2O)]Cl2 (dmp = 2,9-dimethyl-1,10-

phenanthroline) as a new electrochemical hybridization indi-

cator was recently demonstrated by Li and coworkers [51].

[Cu(dmp)(H2O)]Cl2 can intercalate into the base stack of ds-DNA

and has found applications for the detection of a Hepatitis B sensor

based on a synthetic 21-mer ODN sequence.



7.3.4 Peptide Nucleic Acids (PNA)

The performance of “classic” DNA sensors is affected by hybridiza-

tion efficiency, which depends on a number of factors such as

temperature, ionic strength, probe length, and others. However,

many of these problems can be minimized by the use of peptide

nucleic acids (PNA). PNAs are artificial DNA analogues in which

the ribose phosphate ester backbone is replaced by pseudo-peptide

backbone (see Scheme 7.3) [52].

Scheme 7.3. Structures of DNA and PNA, where the ribose phosphate

diester backbone (DNA) is replaced by pseudo-peptide backbone and the

nucleobases are attached to this backbone via methylene carbonyl bonds

(PNA).

Nucleobases are linked to the PNA backbone by methylene

carbonyl bonds. The PNA undergoes sequence-selective binding to

RNA and DNA [53]. Since the backbone of PNA contains no charged

phosphate groups, there are no electrostatic repulsions between

the backbones, enabling a stronger interactions for PNA/DNA

compared to the corresponding DNA/DNA. In addition, the stability

of the PNA/DNA duplexes is virtually unaffected by the ionic

strength of the medium, making it an interesting alternative in

DNA biosensing and mismatch detection [54]. The first use of PNA

as recognition layer for DNA biosensors was reported by Wang



[55], who demonstrated that the PNA film retains its efficient

hybridization properties under a variety of conditions and therefore

offers significant advantages over “classic” DNA based capture

probes.

Faster hybridization, minimal dependence on ionic strength,

and higher specificity and sensitivity (including discrimination

for single-base mismatches) were demonstrated. Electrochemical

detection of a single nucleotide base pair mismatch was achieved

using a mixed film composed of PNA and 6-mercapto-1-hexanol

as diluent [56]. Figure 7.6 outlines the principle of the PNA

biosensor proposed by Aoki et al. Binding of the complementary

oligonucleotide to the PNA probe increased the negative charge

at the electrode surface resulting in an increased electrosta-

tic repulsion between the monolayer and the redox marker

[Fe(CN)6]4−/3− present in solution. In essence, the redox reaction of

the redox probe was hindered upon hybridization with the target

DNA, Fig. 7.6b. Subsequently, Wang and coworkers reported an

electrochemical impedance spectroscopy (EIS) study on these mixed

alkanethiol/PNA films [57], providing insight into the repulsive

interactions between [Fe(CN)6]3−/4− in the presence of matched

films and those containing a single nucleotide mismatch containing

PNA/DNA hybrids. Hashimoto et al. used PNA as part of an electrode

array sensor [58]. Synthetic PNA probes modified with the thiol-

containing amino acid cysteine were immobilized on the gold

electrodes of the array. Hoechst33258 is known as a minor groove

binder and specifically binds to ds-DNA and was exploited in this

study. In contrast to other DNA binding molecules that often bind

not only to the hybrids but also to the single strands, Hoechst33258

only binds to ds-DNA. The array was used for detection of the

PCR amplified cancer gene ras. The PNA showed stronger binding

affinity for complimentary DNA than for strands with a single base

mismatch allowing the detection of point mutations. In another

investigation, nanogold-modified electrodes were used to increase

the amount of immobilized ss-PNA capture probes leading to

an increase in the electrochemical signal [59]. Fc-functionalized

polythiophene was used as a cationic hybridization indicator that

adsorbed onto the negatively charged DNA backbone, giving rise

to a clear hybridization signal in the CV and DPV. The method



Figure 7.6. (a) Working principle of sensor for oligonucleotides based

on the PNA probe immobilized on gold electrodes. Electrostatic repulsion

between the negatively charged marker (represented as an octahedron) and

the PNA/DNA duplexes at the electrode surface hinders the redox reaction

of the marker. (b) Cyclic voltammograms of [Fe(CN)6]4−/3− measured

with the gold electrode modified with a mixed monolayer of PNA probe

and 6-mercapto-1-hexanol before (A, dashed line) and after incubation

in a solution of 100 mM one-base mismatch oligonucleotide at room

temperature (B), 37◦C (C), and 47◦C (D) for 40 min. H. Aoki, P. Buhlmann,

and Y. Umezawa, Electrochemical detection of a one-base mismatch in

an oligonucleotides using ion-channel sensors with self-assembled PNA

monolayers, Electroanalysis, 2000, 12, 1272–1276. Copyright Wiley-VCH

Verlag GmbH & Co. KGaA. Reproduced with permission.

allowed for discrimination against complementary and four-base

mismatch DNA. An interesting reagentless PNA-based sensor was

reported by Reisberg et al. [60]. The working principle of the sensor

is summarized in Fig. 7.7.



Figure 7.7. Working principle of the DNA electrochemical sensor based

on a PNA functionalized conductive polymer. Reprinted from Talanta, 76,

S. Reisberg, L. A. Dang, Q. A. Nguyen, B. Piro, V. Noel, P. E. Nielsen, L. A. Le,

and M. C. Pham, Label-free DNA electrochemical sensor based on a PNA-

functionalized conductive polymer, 206–210, 2008, with permission from

Elsevier.

Here, the PNA capture probe was covalently attached to a

quinine-based electroactive polymer. Changes in flexibility of the

PNA probe strand upon hybridization generate electrochemical

changes at the polymer-solution interface. A reagentless and

direct electrochemical detection was achieved by detection of the

electrochemical changes using square wave voltammetry (SWV). An

increase in the peak current of quinone is observed upon hybridiza-

tion of probe to the target, whereas no change is observed with non-

complementary sequences. In addition, the biosensor can effectively

discriminate a single mismatch on the target sequence. A different

PNA based sensor that does not require probe immobilization

was proposed by Luo et al. [61]. This method involves solution

phase hybridization of a Fc-labeled PNA and its complementary

DNA sequence, followed by the electrochemical detection of Fc-

PNA-DNA hybrid on indium tin oxide (ITO)-based substrates. Due

to the electrostatic repulsion between the negatively charged ITO

surface and the negatively charged DNA reduced electrochemical

signal was observed in respect to signal observed for neutral Fc-

PNA conjugate. However, when the ITO electrode was coated with

a positively charged poly(allylamine hydrochloride) (PAH) layer, the



electrostatic attraction between the sensor surface and the Fc-PNA-

DNA hybrid caused a significant increase of the electrochemical

signal, which is proportional to the amount of complementary DNA

present. Importantly, the PAH-modified sensor was found to be

more sensitive (with a detection limit of 40 fM) than the bare

ITO substrate (with a detection limit of 500 fM). The method was

further validated by discrimination of fully matched and mismatch

DNA strands at elevated temperatures and detection of unpurified

PCR amplicons with detection limit of 4.17 aM. Recently, a new

strategy was reported that makes use of the minor groove binding of

singly reduced cation radical viologen (V) groups C12VC6VC12 [62].

In the presence of complementary PNA-DNA hybrids, the V 2+/+

redox couple of C12VC6VC12 exhibited a unique double-wave cyclic

voltammogram, with the formal potential shifted by –100 mV from

the E f in the presence of single base mismatched DNA-PNA hybrids

or PNA probes alone.

Without a doubt, unique properties make PNA an interesting,

although sometimes synthetically challenging and expensive, alter-

native for design of biosensors.

7.3.5 Protein Mediated DNA Biosensors

MutS is a 97 kDa protein that is part of the DNA repair “engine”

in E. coli. The protein binds to many of single nucleotide DNA

mismatches and has been used for label-free nucleotide mismatch

detection. There are several reports of utilizing MutS to detect

single nucleotide mismatches by a number of different analytical

tools. However, electrochemical techniques have recently been used

due to inherent sensitivity. It was observed that for alkanethiol-

diluted ds-DNA on gold, the charge transfer resistance Rct increases

considerably after binding of MutS to a A-C mismatch, while no

change in Rct was observed when measuring the electrochemical

impedance of matched DNA duplex in the presence of MutS since the

enzyme does not bind to fully matched ds-DNA (see Fig. 7.8) [63].

Palecek et al. [64] and Masarik et al. [65] detected a G-T

mismatch at CPE and HMDE using chronopotentiometric stripping

analysis (CSA) and squarewave voltammetry (SWV) in the presence

of MutS. Cho et al. [66] found that the binding affinity of MutS for



Figure 7.8. Electrochemical impedance measurements in the presence

of [Fe(CN)6]3−/4− of (a) matched and (c) single mismatched ds-DNA films;

(b) and (d) schematically represent MutS interactions with matched and

mismatched films, respectively. Please note the significant increase in the

impedance signal as a function of MutS bound to the surface, causing a

significant increase in the charge transfer resistance. C.-Z. Li, Y.-T. Long, J.

S. Lee, and H.-B. Kraatz, Chem. Commun., 2004, 574–575. Reproduced by

permission of the Royal Society of Chemistry.

different mismatches in the order of GT>CT>CC by CV and EIS on

modified gold electrodes. Han et al. [67] approached the problem in

a slightly different way. Instead of tagging DNA mismatched duplex,

they tagged gold electrode through a histidine-Ni-nitriloacetate

complex and measured the current decrease due to the electrostatic

repulsion between the anionic redox probe and polyanionic DNA

bound to MutS on the electrode. The detection limit of about 500

fM for a G-T mismatch is certainly encouraging. Furthermore, the

signal strength varies with the nature of the mismatch according

to TG>GG>AC=AA=AG>TT>CT>CC. A more complex approach

was adopted by Chen et al. [68] involving the binding of

methylene blue labeled mismatched DNA to MutS immobilized



(A) (B)

Figure 7.9. (A) Signal generation in a pseudo-knot E-DNA sensor. Binding

of complementary target DNA causes conformational changes in the redox-

labeled, electrode-bound capture probe. (B) Optimal signal gain (relative

current change) observed in the presence of perfectly matched (PM),

single (1MM), double (2MM) and triple (3MM) mismatches. Reproduced by

permission from K. J. Cash, A. J. Heeger, K. W. Plaxco, and Y. Xiao, Anal. Chem.,2009, 81, 656–661. Copyright 2009 American Chemical Society.

on a AuNP layer on a gold electrode followed by impedance

measurement.

7.3.6 DNA Stem-Loops

Tyagi et al. [69] developed the concept of a “molecular beacon”

for DNA mismatch detection, consisting of a hairpin-like DNA stem–

loop structure having a fluorophore and a quencher at opposite

terminals. Upon hybridization with a complementary target strand,

the conformational change associated with strand binding and con-

version of the stem–loop into linear duplex results in an increased

distance between the fluorophore and the quencher proximity and

resulting in emission. Subsequently, this strategy was developed into

an electrochemical DNA sensor (E-DNA) with the help of a redox

label attached to the stem-loop. Conformational changes induced by

hybridization significantly alter the distance between the electrode

and redox label, resulting in a change of electron transfer efficiency,

readily detectable by CV [70]. Plaxco and Heeger [71] used E-

DNA for the detection of different mismatches (such as C-A, C-

C, C-T) as well as single and multiple mismatches in presence of

organic/inorganic contaminants in a “signal-off” format. E-DNA is

selective to the target sequences in presence of contaminants since



the sensing event depends on the conformational change rather than

on adsorption to the electrode surface.

In addition, it was shown that E-DNA can readily detect single

nucleotide mismatches and can be recycled for the multiple assays.

A “signal-on” type E-DNA sensor based on a pseudo-knot is shown

in Fig. 7.9. In this system, hybridization induced conformational

change brings the redox label close to the electrode surface

and thus enhances electron transfer efficiency [72]. The system

is characterized by the 5’-end being attached to the transducer

surface while the redox-labeled 3’-end forms a pseudo-knot loop

that hybridizes on top of the hairpin loop. The stability of the

pseudo-knot loop at the 3’-terminus depends on the number of

base pairs, which was found to be 7 bp for maximum stability.

The signal response was found to be enhanced twofold with a

more flexible poly(T) loop as compared to poly(A) loop for all

systems investigated, including for fully-matched, single, double and

triple mismatches. The pseudo-knot-based sensor was found to be

selective in presence of serum and sensitive up to 30 pmoles. More

recently, a method was reported based on unlabeled stem loop

structures. Hybridization to the stem loop and opening of the stem

loop will alter the film structure, generally resulting in an increase

in the film thickness.

The charger transfer resistance Rct for electron transfer from

the anionic redox probe [Fe(CN)6]3−/4− through the film will

be greatly influenced by this conformational change. Importantly,

differences in the film caused by the presence of single nucleotide

mismatches are sufficiently large that they cause differences in the

Rct. In particular, the addition of Zn2+ ions amplifies the resistive

differences allowing the detection of single nucleotide mismatches

at concentrations as low as 10 pM [73] (see Fig. 7.10). The effect of

the metal ions is discussed in more detail in the section on metal-ion

amplified sensors.

7.3.6.1 Enzyme-mediated sensors

There are a number of reports of mismatch detection strategies, in

which enzymatic reactions are exploited to amplify the electrochem-

ical signal.



Figure 7.10. (A) An unlabeled stem-loop structure immobilized on gold

electrode opens up in the presence of target DNA, forming a film of

matched and mismatched ds-DNA, respectively. (B) Nyquist plots shows

in the increase in the charge transfer resistance of the DNA film after

hybridization; Rct of hairpin (a), mismatched duplex (b) and matched duplex

(c). Inset shows the modified Randle’s equivalent circuit used to fit the

electrochemical data. (C) Relationship between �Rct and the concentration

of the target strand showing sensitivity up to 10 pM. Y. Wang, C. Li, X. Li, Y. Li,

H.-B. Kraatz, Anal. Chem., 2008, 80, 2255–2260. Copyright 2008 American

Chemical Society.

In a sandwich-type enzyme sensor, the target strand is

hybridized to the immobilized capture probe and then the hanging

part of target is further hybridized with a label-conjugated detection

probe. Thus, this detection format eliminates the modification of

the target strand. Heller et al. [74] reported the first enzyme-

amplified DNA mismatch detection using an 18-mer capture probe.

7 μm carbon electrodes coated with a polymer containing a cationic

Os-complex were used as transducer surfaces. Next, soybean

peroxidase (SBP) labeled target DNA was hybridized to the capture

strand, bringing the redox polymer and the enzyme in close contact.



Figure 7.11. (a) Formation of SAM on Au electrode (a), immobilization of

Fc-D (b), immobilization of thiolated capture probe (c), hybridization with

target(d), hybridization with biotinylated detection probe (e), association

with avidin-alkaline phosphatase (f), electrocatalytic reaction of p-AP via

electronic mediation of Fc-D (g). (B) Cyclic voltammogram of enzyme-linked

electrodes in the case of hybridization with (a) complementary target,

(b) single basepair mismatched target, (c) non-complementary target and

(d) without hybridization with target and detection probe. Reproduced by

permission from E. Kim, K. Kim, H. Yang, Y. T. Kim, and J. Kwak, Anal. Chem.2003, 75, 5665–5672. Copyright 2003 American Chemical Society.

This, in turn, switches the film property from catalytically inactive to

an active catalyst for H2O2 electro-reduction, which was measured

amperometrically. Kim et al. [75] approached the detection problem

through sandwich-type by immobilizing a capture probe on a Fc-

tethered dendrimer (Fc-D) modified gold electrode. The enzyme,

alkaline phosphatase (ALP), was attached on the other end of

detection probe through the avidin-biotin conjugation system

(described in a previous section). ALP generates the electroactive p-

aminophenol (p-AP) from p-aminophenyl phosphate (p-APP), which

is catalytically oxidized on electrode surface mediated by the redox-

active dendrimer. Fig. 7.11 shows the CV response of the system

as a function of target strand concentration. In a separate study,

the signal was amplified by deposition of Ag particles on electrode

surface by electrochemical reduction through p-AP. This strategy

enhanced the sensitivity up to 100 aM [76]. David et al. [77] used

a direct-type sensor in which capture probe was immobilized on



a screen-printed carbon electrodes through avidin-biotin coupling

and hybridized with a labeled target strand. Catalytic currents

generated by ALP transformation were quantified voltammetrically

and decreased in case of a base pair mismatch with a sensitivity

up to 0.49 fM. Liu et al. [78] utilized the stem-loop capture

strand, a prototype of E-DNA as discussed above, in which the

capture strand was initially labeled with DIG (digoxigenin) which

was sterically shielded from a bulky horseradish peroxidase (HRP)

due to the particular structural conformation of capture strand. The

hybridization to target DNA makes the DIG accessible to the anti-

DIG-HRP. The successful hybridization event can be easily evaluated

electrochemically. In presence of a single base pair mismatch, the

current is significantly reduced and decreases further in presence of

multiple mismatches. Impedance measurement can be a method of

choice to detect the enzyme-amplified signals because of its inherent

sensitivity [79].

7.3.7 Nanoparticle-Based Sensors

A number of reports appeared in 2001 by Authier and Wang

et al. describing magnetic beads/nanoparticles based electrochem-

ical detection of DNA hybridization using stripping voltammetry

[80, 81]. Magnetic bead based DNA sensors for mismatch detection

circumvents nonspecific adsorption effects of protein, RNA, and non-

complementary oligomers through magnetic separation. Typically,

a prototype magnetic bead based sensor relies on (a) an inosine-

substituted capture probe sequence linked to streptavidin-coated

magnetic particles, (b) hybridization and magnetic removal of non-

hybridized oligonucleotides, (c) alkaline treatment to release the

hybrid from the magnetic particles and denaturing of the duplex,

and finally (d) potentiometric stripping detection of the target

strand’s guanine oxidation peak [81]. This approach can be linked

to enzymatically coupled reactions [82], binding of the metal and

amplified electrochemical detection of the dissolved AuNPs [83],

AgNPs [84], CdSNPs [85], as well as solid state stripping of AgNPs

[86] and multi-target analysis [87] as indicated in Fig. 7.12. This

approach does not give a signal for non-complementary target and

only low signal for a target with single or few mismatches as



(A) (B) (C) (D) (E)

Figure 7.12. Magnetic beads/nanoparticles based protocols for electro-

chemical detection of DNA. These assays involve the introduction of the

probe-coated magnetic beads, addition of the target/hybridization event,

magnetic removal of unwanted materials, binding of the metal and amplified

electrochemical detection of the dissolved gold (A), silver (B) and cadmium

sulfide (D) nanoparticles. (C) Solid-state stripping and (E) multi-target

detection protocols. Reprinted from Analytica Chimica Acta, 500, J. Wang,

Nanoparticle-based electrochemical DNA detection, 247–257, 2003, with

permission form Elsevier.

compared to complementary target. In addition, this approach does

not discriminate between different types of mismatches. Wang et al.developed a new type of sensor that relies on the mononucleotide

linked nanocrystals, i.e., A-ZnS, C-CdS, G-PbS, and T-CuS that bind

with their complementary nucleotide bases at mismatch sites on

dsDNA modified magnetic bead and thus results in each mutation

with specific nanocrystal-mononucleotide tags.



Subsequently, the mismatches can be identified with the voltam-

mogram peak potentials of their nanocrystal-mononucleotide tags

[57].

7.3.8 Metal-Ion Amplified Sensor

For practical application, an ideal biosensor must be as straightfor-

ward as possible with least number of synthesis and analytical steps.

Kraatz and coworkers introduced a simple, label-free and sensitive

electrochemical sensor for single nucleotide mismatch detection.

This approach relies on the diffusive property of the negatively

charged redox probe [Fe(CN)6]3−/4− and its interplay with matched

and mismatched DNA films.

Again, the charge transfer resistance Rct for electron transfer

from the solution based anionic redox probe to the transducer

surface is used as a quantifiable measure and is evaluated in the

presence and absence of Zn2+ using EIS as indicated in Fig. 7.13. The

presence of Zn2+ in the electrochemical experiment is significant

in that it influences the ability of the [Fe(CN)6]3−/4− to diffuse into

the DNA film. In the presence of Zn2+ the metal ion will interact

with the phosphate backbone, lowering the electrostatic repulsion

with the anionic redox probe, resulting in a lower charge transfer

resistance. In addition, there are significant differences between

the Rct for matched and mismatched DNA films. Presumably this

is due to differences in packing within the film. Generally, for a

mismatched film, the Rct value will be lower since the mismatched

film will be less densely packed, allowing a better penetration of

the redox probe into the film. Differences in Rct are evaluated in

the presence and absence of Zn2+ and in the presence of absence

of a mismatch [88]. This approach allows the detection of single

nucleotide mismatches down to 10 fM level. The method is tolerant

to protein contaminations and also to heterozygote DNA mixtures.

In the absence of Zn2+, the mismatch detection limit is in the

order of 100 nM [89, 90]. The sensitivity produced by metal ions

in ds-DNA film was further confirmed by K’Owino et al. [91]

who showed that the addition of Ag+ to a ds-DNA film gives a

stronger electrochemical response compared to the response for

a ss-DNA film. The simple label-free approach described shows a



Figure 7.13. (A) Schematic showing the electron transfer process across

the dsDNA film between the negative redox probe [Fe(CN)6]3−/4− and gold

transducer surface. Electron transfer process is facilitated by the addition of

metal ion that neutralizes the phosphate backbone of DNA and allows the

enhanced diffusion of the redox probe. As a result the differences in charge

transfer resistance Rct before and after metal ion addition are significantly

different and are in fact affected by the presence of a single nucleotide

mismatch. (B) Nyquist plot showing the charge transfer resistance across a

matched and a mismatched film in absence and presence of Zn2+ in the form

of semicircle. Inset shows the modified Randle’s equivalent circuit used to

fit the data. (C) The plot showing the detection limit of the system as low

as 10 fM. Reproduced by permission from X. Li, J. S. Lee, H.-B. Kraatz, Anal.Chem. 2006, 78, 6096–6101. Copyright 2008 American Chemical Society.

high potential for applications also in an array electrode format

and has allowed to detect a range of different mismatches [92, 93].

Scanning electrochemical microscopy (SECM) studies were critical

to elucidate the mechanism of this process and rationalize the

differences in Rct in terms of the diffusive properties of the probe

molecules (see Fig. 7.14) [94, 95]. Using SECM, the heterogeneous

electron transfer constants were evaluated and it was shown that

in the presence of Zn2+ the ket increases from 4.6 × 10−7 cm/s (no

Zn2+) to 5.0 × 10−6 cm/s (Zn2+ added).

Based on the initial SECM results, it was postulated that it should

be possible to evaluate differences in Rct directly by SECM and

monitor the amperometric feedback current in the presence and

absence of Zn2+. The presence of SNP caused an increase in electron

transfer rate constant, presumably due to better penetration of



Figure 7.14. Schematic diagram of SECM measurement of electron

transfer through DNA duplexes that allowed to provided a possible

mechanism for the differences in charger transfer resistances before and

after Zn2+ addition to dsDNA films. B. Liu, A. J. Bard, C.-Z. Li and H.-B. Kraatz

J. Phys. Chem. B 2005, 109, 5193–5198. Copyright 2005 American Chemical

Society.

the redox probe into the film and are sensitive not only to the

presence or absence of a single nucleotide mismatch but also to

its position. Figure 7.15 shows measurements with mismatches in

three different positions within the ds-DNA. All three systems give

a distinct amperometric response, which was amplified after the

addition of Zn2+. Moreover, impedimetric study also corroborates

the SECM results [96]. Recently, SECM has shown strong potential

towards the application for species identification [97, 98].

Recently, the effects of various metal ions on the electrochemical

impedance spectra of 25-mer dsDNA films were reported. These

metal ions include Mg2+ and Ca2+, known to have high affinity for

the phosphate backbone of DNA, the trivalent Al3+ and La3+, and

divalent transition metal ions Ni2+, Cu2+, Zn2+, Cd2+ and Hg2+. In all

cases, the presence of metal ions decreases the Rct of ds-DNA films,

presumably due to their coordination with the backbone phosphate

and potentially association with one or more of the exocyclic



Figure 7.15. SECM image with corresponding current profile recorded for

a matched and three different mismatched dsDNA microarrays on a gold

substrate in the absence (a) and presence of of Zn2+ (b). Please note the

differences in the normalized current as a function of mismatch position,

which is enhanced in the presence of Zn2+. P. M. Diakowski, H.-B. Kraatz,

Chem. Commun., 2009, 1189–1191. Reproduced by permission of the Royal

Society of Chemistry.

N -atoms of the purine bases. The �Rct of the different metal ions

was found in the order of Ca2+ > Mg2+ = Hg2+> Cd2+ >Ni2+ >

Cu2+ > Zn2+ which is inversely proportional to their free energies

of hydration (see Fig. 7.16) [99].

7.3.9 Miscellaneous Methods

Recently, Zhu et al. [100] exploited PAMAM dendrimers to tag the

target strand which increase the Rct on hybridization with surface

immobilized capture strand. The �Rct can easily distinguish ds-

DNA without PAMAM tag, mismatched and non-complementary

sequences from PAMAM tagged complementary ds-DNA with

picomolar sensitivity (Fig. 7.17).

Watanabe et al. [101] detected DNA mismatches through a

strand exchange reaction in which the duplex consisting of a

capture probe and a redox-labeled probe strand are immobilized on



Figure 7.16. The relationship between �Rct and free energy of hydration

of divalent metal ions. X. Bin, H.-B. Kraatz, Analyst, 2009, 134, 1309–1313.

Reproduced by permission of the Royal Society of Chemistry.

Figure 7.17. (a) Schematic representation of a gold surface modified with

a ssDNA capture strand followed by hybridization with a ssDNA-PAMAM

target and the formation of dsDNA-dendrimer hybrid; (b) AFM image of

the PAMAM on the mica surface. N. Zhu, H. Gao, Y. Gu, Q. Xu, P. He and

Y. Fang, Analyst, 2009, 134, 860–866. Reproduced by permission of The

Royal Society of Chemistry.



electrode surface. The redox-labeled probe strand was replaced by

the complementary/noncomplementary target strands. The slower

rate of mismatched strands discriminated them from the fast

complementary sequences.

Kwon et. al. [102] introduced a signal on/off sensor based on

enzymatic cleavage of the unhybridized Fc-labeled ss-DNA resulting

in lower electrochemical response for single mismatched strand and

no signal for non-complementary sequences. Another interesting

report by Panke et. al. [103] shows the electrochemical assay based

on competitive binding between the non-labeled target and the MB-

labeled reporter strand with a surface immobilized capture strand.

Sensitivity was reported up to 3 pmolar for nonlabeled binding

assay. Recent improvements include the use of locked-DNA (LNA),

[104] Scheme 7.4, and morpholino-oligomers, [105] to improve the

hybridization affinity.

Scheme 7.4. Example of a locked nucleic acid (LNA) which is significantly

more rigid compared to conventional nucleic acids.

LNAs contain a methylene bridge that connects the 2’-oxygen

atom with the 4’-carbon atom of the ribose ring of the ribonucleic

acid resulting in a locked 3’-endo conformation, which reduces the

conformational flexibility of the ribose and increases the degree of

local organization of the phosphate backbone. Presumably entropic

constraint improves the ability of hybridization affinity of the

capture strand. On the other hand, morpholino-oligomers are DNA

analogs in which the sugar phosphate backbone is replaced with


References 239

morpholine rings and bonded through phosphorodiamidate groups,

resulting in an uncharged nucleic acid. These structures are stable,

highly water soluble, and are cost-effective DNA analogs, which

exhibit improved base stacking compared to PNA analogues. The

utility of LNAs for electrochemical sensing of mismatches remains

to be explored but one can envision that the resulting duplex should

exhibit significantly different properties that can be exploited for

sensing.

7.4 Conclusion

Electrochemical detection of DNA mismatches continues to attract

significant attention of the research community. Numerous mis-

match detection schemes have been proposed, some of which even

have led to some limited commercial exploration and start-ups. The

reported detection methods vary widely from relatively simple ones

that exploit the intrinsic electrochemical properties of DNA and

electric properties of the DNA films, to more complex ones that

employ novel bioconjugates, nanopartiocles and DNA analogues.

This growing interest in electrochemical DNA biosensors is often

driven by the unique advantages offered by the electrochemical

detection methods. Application of electrochemical methods in

affinity DNA mismatch detection presents likely a promising

alternative for widely used optical methods, potentially allowing

miniaturization with the associated cost reduction, and potential

application in point-of-care assays. Clearly, the future is promising

for electrochemical DNA sensing and much can be expected in the

next few years.

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Chapter 8

Electrochemical Detection of DNAHybridization: Use of Latex to ConstructMetal-Nanoparticle Labels

Mithran Somasundruma and Werasak Surareungchaib

aBiochemical Engineering and Pilot Plant Research and Development Unit,National Center for Genetic Engineering and Biotechnology atKing Mongkut’s University of Technology, Thonburi,Bangkhuntien Campus, Bangkok 10150, ThailandbSchool of Bioresources and Technology, King Mongkut’s University of Technology,Thonburi, Bangkhuntien Campus, Bangkok 10150, [email protected]

8.1 Introduction

Of the detection schemes available for DNA biosensors [1], elec-

trochemistry has drawn increasing interest due to enabling high

sensitivities using equipment of relatively low cost. In addition,

electrochemical detection can be coupled readily with available

minaturization technologies [2]. The direct electro-oxidation of

guanine involves high background signals [3], while the use of

enzyme labels may involve deterioration of enzyme activity over

time. Redox compounds which can intercalate with the probe-target



246 Electrochemical Detection of DNA Hybridization

duplex can provide a more stable signal [4] but do not always

provide adequate sensitivity. This has led to interest in using

metal nanoparticles as labels for electrochemical detection of DNA

hybridization (reviewed in Refs. 5–7).

Where the nanoparticles are reacted directly (rather than being

used to catalyze reactions), the achievable sensitivity will depend

largely on the quantity of metal attached to each DNA sequence.

This has led to the development in label construction illustrated in

Fig. 8.1. From the binding of individual particles, researchers have

sought techniques to attach assemblies of nanoparticles to a given

DNA sequence. As will be described in this chapter, latex colloids

provide an ideal base for such assemblies, both as solid supports

and as templates for the construction of hollow capsules which can

take up nanomaterials. The point of importance is that the necessary

latex modifications have already been intensively researched for

other applications, and so the relevant physical chemistry theory

and experimental details are already available. Despite this fact, the

use of latex in constructing electrochemical DNA labels is relatively

unexplored.

8.2 Synthesis of Metal Nanoparticles

Colloidal gold was first prepared and studied by Faraday in 1857 [8].

In the early 1950s, the preparation of colloidal gold in homogeneous

solution was described by Turkevich et al. [9] using sodium citrate

to reduce a dilute solution of HAuCl4 under heating. This method

has become a standard for gold nanoparticle preparation and has

also been applied for the synthesis of platinum nanoparticles by

the reduction of PtCl62− [10]. Similar homogeneous synthesis can

be performed for silver nanoparticles, using NaBH4 as a reducing

agent for AgNO3 [11]. When the nanoparticles are formed there

needs to be a force resisting coagulation present for the particles to

remain stable in solution. This force can be provided by electrostatic

repulsion due to the adsorption of ions onto the metal surface and

in some cases the adsorption of the reducing agent (e.g., when

AuCl4− is reduced by citrate the citrate ions remain adsorbed on

the particles and impart a negative charge [12]). The electrostatic


Synthesis of Metal Nanoparticles 247

Figure 8.1. TEM images demonstrating strategies for nanoparticle

labeling of DNA. (a) Attachment of an individual gold nanoparticle.

(b) Attachment of a latex sphere bearing many gold nanoparticles. (c)

Attachment of a gold nanoparticle-latex sphere after gold enhancement by

autocatalytic deposition. (a) and (b) taken with permission from [131],

S. Pinijsuwan, P. Rijiravanich, M. Somasundrum, and W. Surareungchai,

Anal. Chem. 80, 6779–6784 (2008) c© American Chemical Society. (C)

Unpublished results.



stabilization can be added to by steric factors. If a polymer layer is

adsorbed on the particle or tethered to the particle at one end, then

this will further limit the inter-particle approach [13]. Examples of

polymeric stabilizers include polyvinyl alcohol (PVA) and sodium

polyacrylate. Typically the stabilizer is present during the metal-ion

reduction, and this means it can have an effect on the growth process

of the particle. Strong polymer adsorption will slow the growth

rate. Stabilizers can also have a catalytic effect on the reaction

[10]. In some cases a variation in the stabilizer concentration

can change the nanoparticle shape [14]. With regard to gold

nanoparticles, the synthesis can be performed and then long-chain

molecules get attached to the gold by a thiol terminus [15–16]. The

polymer stabilizer can also be provided by performing the reaction

in a water-in-oil (w/o) microemulsion. This is done by reacting

reverse micelles containing a metal salt with reverse micelles

containing reducing agent [17–20]. Mixing the two microemulsions

causes an exchange of material between the micelles. The reaction

occurs first at the edges of the micelle (the initial locus of the

reaction) and then moves into the centre, as demonstrated by

TEM studies [21]. Nanoparticles can also be synthesized from

a single microemulsion, usually containing the metal salt, while

adding the reducing agent directly to the mixture [22–23]. The

principle of microemulsion synthesis has been extended to water-

in-supercritical CO2 microemulsions, the rationale being that the

nanomaterial can be simply recovered by reducing the pressure and

releasing the resulting gas. Silver [24] and copper nanoparticles [25]

have been reported.

In general terms, if the rate of growth of the nanoparticles

is high relative to the rate of nucleation (i.e., the rate of new

particles forming), then the resulting materials will have a narrow

size distribution. This is highly desirable if the particles are to

be used as electrochemical labels, since the size distribution will

affect the precision of the resulting sensor. The rate of the reaction

can be influenced by the nature and concentration of the reducing

agent, with strong reducing agents favoring a faster reaction rate

and smaller nanoparticles [13]. Note however, that an overall fast

reaction does not necessarily imply a faster rate of growth relative

to nucleation.


Use of Metal Nanoparticles as Electrochemical Labels 249

8.3 Use of Metal Nanoparticles as Electrochemical Labels

Metal nanoparticles were first introduced as labels for DNA

sensing by Mirkin and coworkers [26] in 1996. The gold-labeled

ssDNA probes were used to detect complementary DNA targets

by a colorimetric method based on particle aggregation [26–28].

In 2000, Limoges and coworkers [29] became the first group

to use metal nanoparticle labels for electrochemical detection,

in an immunoassay. The group then extended this concept to

electrochemical DNA hybridization detection, based on labeling an

oligonucleotide with gold nanoparticles [30]. The assay, depicted

in Fig. 8.2, consisted of four steps: (a) passive adsorption of the

amplified target DNA on the walls of a polystyrene microwell, (b)

hybridization with an oligonucleotide probe conjugated to an Au-NP,

(c) oxidative gold metal dissolution in an acidic bromine-bromide

solution, and (d) anodic stripping voltammetry (ASV, see Sec. 8.4)

detection of the released Au3+ ions at a screen-printed microband

electrode (SPMBE) immersed in the microwell. The combination

of the sensitive Au3+ determination at a SPMBE with the large

number of Au3+ ions released from each gold nanoparticle allowed

detection down to 5 pM of an amplified human cytomegalovirus DNA

fragment.

In the same year (three months after Limoges’ work was

published), Wang’s group also reported a DNA hybridization assay

Figure 8.2. DNA detection scheme based on the capture and dissolution

of individual gold nanoparticles, followed by voltammetric detection at

a screen-printed microband electrode. Taken with permission from [30],

L. Authier, C. Grossiord, P. Brossier, and B. Limoges, Anal. Chem. 73, 4450–

4456 (2001). c© American Chemical Society.



based on gold nanoparticles [31]. The method differed slightly

from Ligomes’s work in that instead of microwells, magnetic beads

were used, and electrochemical detection was by potentiometric

stripping analysis. The protocol was based on the hybridization of

a target oligonucleotide to a magnetic bead-linked probe, followed

by binding of streptavidin-coated gold nanoparticles to the captured

DNA, then dissolution of the gold label and potentiometric stripping

measurement of the liberated Au3+ ions at a screen-printed carbon

electrode, as depicted in Fig. 8.3. Alternatively, direct oxidation,

using DPV, of the gold nanoparticle label contained in a duplex

Figure 8.3. DNA detection scheme based on immobilizing DNA probes

onto magnetic beads and attaching individual gold nanoparticles to the DNA

targets after hybridization. Following dissolution Au3+ ions are quantified

by PSA. Taken with permission from [31], J. Wang, D. Xu, A.-N. Kawde, and

R. Polsky, Anal. Chem. 73, 5576–5581 (2001). c© American Chemical Society.



attached to a graphite pencil electrode surface was used by Ozsoz

et al. [32]. Fang and coworkers [33] have reported the use of silver

nanoparticles (AgNP) as the oligonucleotide label. By oxidative

metal dissolution and the indirect determination of the solubilized

Ag+ by ASV at a carbon fiber microelectrode, detection down

to 0.5 pM DNA was reported. The same group has also labeled

oligonucleotide probes with an alloy of gold-coated copper core-

shell nanoparticles for a DNA sensing assay [34]. Hybridization

events between probe and target were monitored by the release

of the copper metal atoms anchored on the hybrids by oxidative

metal dissolution, and then indirect determination of the solubilized

Cu2+ ions by ASV. Despite the good sensitivity of all the above

reports, detection limits remained in the range of nanomolar to

subpicomolar (see Table 8.1). Further improvements are needed

to meet the challenge of detecting as low as hundreds of copies

of target DNA—required to avoid using pre-amplification schemes

such as the polymerase chain reaction.

Since the analytical signal in ASV comes from consumption

of the metal film deposited on the electrode (see Sec. 8.4), the

signal can be increased by increasing the size of the nanoparticle.

However, large diameters (e.g., for gold greater than about 20 nm)

are seldom used as electrochemical labels due to reasons such as

poor control of size distribution and poor stability in a solution

of the resulting bioconjugates, causing lower hybridization rates.

A preferable method has been to use smaller nanoparticles, and

then, after hybridization, increase the quantity of the metal by

forming shells of gold or silver on the original nanoparticle through

autocatalytic reduction. Silver deposition has been commonly used

in histochemical microscopy to visualize DNA-conjugated gold

nanoparticles. Based on this concept, Mirkin and coworkers [35]

developed a scanometric DNA array based on silver amplification

of the hybridization event. Wang and coworkers [36] extended this

form of amplification to electrochemical detection by measuring the

deposited silver by stripping analysis. Basically, after hybridization

gold nanoparticles function as catalytic sites for chemical reduction

of silver ions (from silver lactate or silver nitrate) in the presence of

the reducing agent, hydroquinone. Hence, metallic silver is formed

on the gold nanoparticles. This was detected at a screen-printed

March

19,201217:1

PSPBook

-9inx

6in08-O

zsoz-c08

252Electrochem

icalDetection

ofDN

AH

ybridization

Table 8.1. Metal nanoparticle-based hybridization detection methods and their limits of detection

Technique Detection Method Target Detection Limit Reference.

Au-NP label on DNA probe ASV detection of AuIII at SPE DNA fragment 406 bp 5 pM [30]

immersed in microwell

Au-NP label on DNA target Using probe immobilized DNA fragment 19 bp 0.1 μg mL−1 (15 nM) [31]

magnetic bead and PSV

detection of AuIII at SPE

Au-NP label on DNA probe DPV of Au oxidation at graphite 256 bp PCR amplicon 0.78 pM [32]

pencil electrode

Ag-NP label on DNA probe ASV of AgI at carbon fiber electrode DNA fragment 32 bp 0.5 pM [33]

Au coated Cu core-shell NP label on ASV of Cu2+ at GCE Colitoxin gene 24 bp 5.0 pM [34]

DNA probe

Au-NP label on DNA target/Ag dep. ASV of AgI at SPE DNA fragment 19 bp 0.2 ng mL−1 (32 pM) [36]

Au-NP label on DNA probe/Ag dep. DPV of Ag oxidation at GCE DNA fragment 32 bp 50 pM [38]

Au-NP label on DNA target/Ag dep. CP detection of Ag at SPE DNA fragment 19 bp 0.2 μg mL−1 (30 nM) [37]

Au-NP label on DNA probe/Ag dep. LSV of AgI at ITO electrode DNA fragment 16 bp Not given (report greater [39]

S = 20 N)

Au-NP label on DNA probe/Au dep. PEG + NaCl used in the catalytic DNA fragment 16 bp 0.6 fM [40]

process / ASV-CV at SPE

Au-NP label on DNA target/Au dep. SWSV at GCE Primer, wildtype and mutant 0.35 aM [41]

DNA of BRAF gene 23– 30 bp

Au-NP label on DNA target Catalyze reduction of p-nitrophenol BRCA 1 gene 31 bp 1 fM [42]

+ NaBH4 at Fc-modified ITO electrode

Pt-NP label on DNA target Pt-NP electrocatalyzed H2O2/LSV DNA fragment 27 bp 10 pM [43]



carbon electrode using potentiometric stripping voltammetry after

acid dissolution. However, excess silver ions are of major concern

in this method, as they can affect the reliability of the stripping-

based detection. This is because the polyanionic DNA backbone

itself can act as a nucleation site for silver deposition following

cation exchange with sodium for ion-pair complexation to the DNA

bases, which can lead to a high background. To obviate this problem,

sodium thiosulfate can be used as a fixer [it transfers the silver

cations to [Ag(S2O3)]5−]. Control of the silver precipitation time

is also needed. Silver-enhanced colloidal gold stripping led to a

dramatic (>100 fold) signal amplification. Instead of dissolving the

silver for stripping analysis, a direct assay of the silver metal can

also be performed by either constant-current chronopotentiometric

detection after magnetic collection of the duplex-linked particle

assembly [37], or a differential pulse voltammetry measurement

of the large number of silver atoms anchored on the duplexes,

using a glassy carbon electrode [38]. Lee et al. [39] reported the

catalytic effects of various gold nanoparticles for silver deposition

on indiumtin oxide (ITO)based electrodes.

The use of silver enhancement may cause a significant back-

ground signal due to non-specific silver deposition on the DNA

support (i.e., the magnetic bead or electrode surface) and/or on

the negatively charged DNA (as mentioned above). Hence, Rochelet-

Dequaire et al. [40] instead used gold ions for the catalytic enhance-

ment, since the gold autocatalytic process offers a lower background.

This is because there is minimal autonucleation from AuCl4− and

less interaction between the anionic AuCl4− and the negatively

charged DNA. Their work showed that classical gold enhancement

procedures based on incubation in a mixture of chloroauric acid

and hydroxylamine could not provide effective amplification, due to

loss of the enhanced gold labels during the post-enlargement rinsing

step. Therefore, the authors modified the enhancement procedure

to use polyethylene glycol and NaCl in the growth media, to act as

an aggregating agent during the catalytic process. This resulted in

retention of the enlarged labels on the bottom of the microwell,

providing a detection limit of 0.6 fM. Liao et al. [41] reported

a similar scheme by using a square wave stripping voltammetry

and were able to detect a mutated BRAF gene associated with



papillary thyroid carcinomas at a detection limit of 0.35 aM.

Selvaraju et al. [42] realized a drawback of the gold autocatalytic

process is that special care is required in the control of deposition

time and temperature to achieve a high signal-to-background ratio

Instead, they used DNA-labeled gold nanoparticles to catalyze the

reduction of p-nitrophenol to electroactive p-aminophenol. The

p-aminophenol can be catalytically cycled back to p-nitrophenol

at a ferrocene-modified indium-tin oxide (ITO) electrode, offering

large signal amplification. The high signal amplification and low

background current enabled the detection of 1 fM target DNA.

Willner’s group has used platinum-nanoparticle labeled DNA where

the nanoparticles catalyzed the reduction of H2O2 with a detection

limit of 10 pM for the hybridization [43].

8.4 Voltammetric Detection of Metal-Nanoparticle Labels

Voltammetry is of interest as a detection method for DNA due to

the fact that it provides high sensitivities and that the equipment

required is relatively cheap in comparison with techniques such

as fluorescence, surface plasmon resonance, and microfabricated

cantilevers, while safety issues exist with radioactive labels. Below

we summarize the voltammetry theory necessary to develop and

test DNA sensors.

8.4.1 Principles of Analytical Voltammetry

The principle of analytical voltammetry is that the current I from a

redox reaction is recorded under conditions of controlled potential

and is used to calibrate the concentration of the reacting species.

The electrode potential is set relative to a reference interface

which ideally does not change potential as the voltage applied

across the cell is changed. To exhibit this property the current

across the interface during equilibrium should be high. Common

reference interfaces are Ag/AgCl/KClsat and Hg/Hg2Cl2/KClsat. To

complete the current path for the reaction a third electrode is usually

incorporated, but two-electrode cells can also be used by passing

current through the reference, provided that current is in the order


Voltammetric Detection of Metal-Nanoparticle Labels 255

Figure 8.4. Screen-printed two electrode cell used for ssDNA immobiliza-

tion. The carbon track working electrode is held at 100 mV vs. the Ag/AgCl

track reference/counter electrode for 30 s in the presence of 20 μL ssDNA

solution. See also Color Insert.

of microamperes or less. An example of such a cell is the screen-

printed electrode strip shown in Fig. 8.4, which was used by us

to immobilize target DNA. The advantage of such a system is (a)

disposability, and (b) only a small volume of electrolyte is needed

to complete the cell, and therefore only a small quantity of DNA is

required.

The redox current is related to the charge Q passed during the

reaction by I = dQ /dt. That charge is connected to the quantity

of material reacting by Faraday’s law, Q = mnF, where m is

the number of moles converted, F is Faraday’s constant, and nis the stoichiometric number of electrons. Equating the material

consumption, the flux of electrons at the electrode must be equal

to the flux of the reacting species. This flux is described by Fick’s 1st

law. Hence, the electrode current is related to the concentration of



the reactant by

In F A

= D(

∂ c∂ x

)

electrode surface

(8.1)

Thus, to describe the current for particular experimental conditions

an expression is required for the concentration gradient at the

electrode. Often this is obtained by first deriving an expression for

c(x). If the experiment is performed in the presence of sufficient

electrolyte to disregard reactant transport by migration, then the

change of c(x) with time will be described wholly by Fick’s 2nd law,

∂ c/∂ t = D ∇2 c (8.2)

where the operator ∇ is dependent on the electrode geometry. Thus,

the expression of c(x) or c(x , t) can be found by solving Eq. (8.2)

under boundary conditions relevant to the experiment. Equation

(8.2) may have to be modified by preceding or following chemical

reactions.

8.4.2 Anodic Stripping Voltammetry (ASV)

One of the main reasons for interest in using metal nanoparticles

as electrochemical labels is that after acid-dissolution the resulting

ions are amenable to detection by ASV. The procedure consists of

two steps, as shown in Fig. 8.5: (1) Preconcentration of the analyte

M n+ by reduction to a film, or mercury amalgam, of M0 on the

electrode surface. (2) Re-oxidation of the metal M0 by scanning

the potential in a positive direction, causing the resulting ions

to be “stripped” back into the solution. Step (1) is performed at

a diffusion-limited potential, usually under stirring or electrode

rotation to maximize the amount of metal deposited. Step (2) is

performed in quiescent solution, resulting in a current peak which

can be used to calibrate M n+ concentration. The rest period between

(1) and (2) is to allow the solution to become quiescent. The

analytical importance of ASV is that while the analyte M n+ may be

present at a low concentration in solution, the analytical signal is

derived from a high concentration of M0 at the electrode surface.

Hence, metal ions can be detected down to 10−10 M to 10−11 M.

ASV was originally performed with a mercury electrode either

in the form of a hanging mercury drop (HMDE) or a thin mercury



Figure 8.5. Potential-time waveform used in ASV. (a) Deposition of metal

ions. (b) Rest period to allow solution to become quiscent. (c) Potential is

driven positive of the oxidation potential of the metal film.

film (MFE). The latter case is produced by reducing a layer of

mercury (thickness ∼1–1000 nm) onto a solid electrode. This can

be done conveniently by adding mercury ions (10−5 M–10−4 M)

to the analyte solution, so that the MFE forms during the analyte

preconcentration. Where the analyte has an oxidation potential

more positive than mercury (e.g., Ag or Au) a solid electrode must

be used. Screen-printed carbon electrodes have been successfully

applied to the ASV detection of metal-nanoparticle labels ([31],

[36]), although obviously a screen-printed electrode surface is less

reproducible than that of mercury. Where such electrodes are used

(and for that matter MFEs) the stripping step will remove virtually

all of the deposited material, resulting in a relatively sharp peak. This

characteristic, combined with the fact that E 0′is unique for each

metal, enables multianalyte detection from a single voltammogram.

Such voltammograms can thus be applied to the simultaneous

detection of more than one DNA sequence by using a different metal

label for each sequence [44].



8.4.3 Quantification

8.4.3.1 Linear sweep voltammetry

As noted above, quantification of the analyte in ASV comes from

the stripping step. Different methods of quantification are available,

based on different ways of scanning the potential. The simplest

method is linear sweep voltammetry (LSV), in which the potential

waveform is a linear increase as illustrated in Fig. 8.6a. For an MFE,

the concentration cM of metal inside a mercury film of thickness l is

Figure 8.6. Potential-time waveform for (a) linear sweep voltammetry

and (b) differential pulse voltammetry.



expressed by [45]

cM = D cb td

l δ(8.3)

where δ is the width of the diffusion layer (dependent on the stirrer

speed or electrode rotation rate), td is the deposition time, cb is the

bulk concentration of the metal ion, and D is its diffusion coefficient.

cM is related to the peak current IP by [46]

IP = n2 F 2ν l A cM

2.7RT(8.4)

where v is the scan rate, A is the film area, R is the gas constant,

and T is the temperature. The situation of a metal film on a screen-

printed electrode should approximate to the case of an extremely

thin mercury film, and therefore, Eq. (8.4) may approximate the LSV

stripping response at such an electrode.

As shown in Fig. 8.7a, a drawback to LSV stripping can be

the rising baseline, which limits the technique’s sensitivity. This

Figure 8.7. ASV detection in sea water containing 30 ppb Cd, 75 ppb

Pb, and 65 ppb Cu at a HMDE using (a) linear sweep voltammetry and

(b) differential pulse voltammetry. Taken with permission from [133], W.

Lund and D. Onshus, Anal. Chim. Acta 86, 109–122 (1976). c© Elsevier Ltd.



baseline represents the capacitive (i.e., non-Faradaic) current Ic

during the potential sweep. The capacitive current arises because

of the rearrangement of ions at the double layer in response to the

changing electrode potential. It is related to the scan rate by [46]

Ic = v C d [1 − exp (−t/RSC d)] (8.5)

where t is the time and RS is the solution resistance, taken as being

in series to the double-layer capacitance C d. Ic increases to reach a

constant value during the scan. Increasing the scan rate will increase

Ip, but will increase Ic by the same amount. In contrast to the

potential sweep, when a potential step to a value E is applied to

the same series resistor–capacitor combination, Ic can be shown to

decay exponentially with time according to [46]

Ic = ERS

exp (−t/RSC d) (8.6)

based on the equation for the charging of a capacitor. However, the

Faradaic current from the same potential step, as expressed by the

Cottrell equation [46], decays in proportion to 1/√

t. Therefore, the

capacitive current falls more quickly. This fact may be utilized in

pulse voltammetry to lower the baseline of the voltammogram, and

thus improve the sensitivity.

8.4.3.2 Differential pulse voltammetry

The potential waveform for differential pulse voltammetry (DPV) is

shown in Fig. 8.6b. The pulse height (�E in Fig. 8.6b) is typically

a few tens of mV, and the pulse width (�t in Fig. 8.6b) is typically

50 to 60 ms. The current is sampled immediately before the pulse is

applied (I1) and then at the end of the pulse (I2). The voltammogram

output is the difference I2–I1 plotted as a function of potential,

as shown in Fig. 8.7b. To understand the principle of DPV we can

consider the value of (I2–I1) at three different stages:

(1) Before the redox process begins. Here (I2–I1) represents the

difference in the capacitive currents at each set of the potentials

where I1 and I2 are measured. Because recording occurs after the

pauses shown in Fig. 8.6b, the currents will have decayed with time

according to Eq. (8.6). Therefore, the value of (I2–I1) will be very

small.



(2) Once the redox process begins. The value of (I2–I1) will

now represent almost wholly the Faradaic current from the

stripping reaction. Both currents will increase first linearly and then

exponentially with potential, according to the low and high field

approximations to the Butler-Volmer equation [47]. Once the region

of exponential increase has been reached the overall value of (I2–

I1) will increase, since the potential ramp is linear and therefore I2

becomes increasingly greater than I1.

(3) After the redox peak is reached. In any potential scan of an

immobilized redox material, a peak is observed due to the depletion

of that material as the voltage is increased. In the case of DPV, once

the current sampling for I2 reaches the peak potential, I2 reaches

its maximum value. However, since the sampling of I1 lags behind,

I1 continues to rise. Therefore the value of (I2–I1) goes down.

Eventually I1 will reach the peak potential also and then (I2–I1)

will be virtually zero (the difference between them will be the small

difference in the residual capacitive current). In this manner, DPV

provides a lower baseline than LSV, as shown in Fig. 8.7b. The DPV

detection limit for a species in bulk solution is estimated at 5 × 10−8

M (c.f. 5 × 10−6 M by LSV), and for stripping this lowers to 1 × 10−11

M due to the advantage of preconcentration (c.f. × 10−10 M for LSV)

[48]. For a species diffusing from bulk solution the DPV peak height

(I2–I1)max is given by [46]

(I2− I1)max = n F A√

D c√π

√�t

(1 − σ

1 + σ

)(8.7)

where

σ = exp

(n F �E2 R T

)(8.8)

The DPV response has also been derived for an HMDE where metal

ions are reduced at the mercury [48], but has not, to the best of

our knowledge been derived for mercury electrodes in conjunction

with ASV. This is probably because such systems are only used for

analytical calibrations and not for the determination of physical

parameters.



8.4.3.3 Potentiometric stripping analysis

In addition to LSV and DPV, chronopotentiometry has also been used

to detect metal nanoparticle labels ([31], [36]). In PSA, the metal

ions from the labels are reduced onto the electrode as described

previously and then the electrode is programmed to pass a constant

current, often in the order of microamps. To satisfy this current the

electrode potential moves to a value where M0 will be reoxidized.

Once M0 is depleted from the electrode the potential must shift

positively until a new redox reaction (possibly solvent electrolysis)

can provide the current. The time τ for this potential transition is the

analytical signal corresponding to current height in voltammetry. To

the best of our knowledge the expression of τ for an MFE used in

ASV has not been derived. For ASV using a HMDE of radius r passing

a current I , τ is related to cM by [49]

τ = nF ArcM

3I− r2

15 D(8.9)

assuming that all (or a considerable part) of M0 in the mercury drop

is oxidized and that the inequality r2 < 7Dt is fulfilled. When the

drop radius is small and the current is low, the second term becomes

negligible. Some of the chronopotentiometric stripping responses

for metal nanoparticle detection [31, 36] have been reported in the

form of peaks with heights measured in units of s V−1, which means

presumably some differential of the current was measured.

8.5 Latex as a Label Support

8.5.1 Introduction

The term “latex” originally referred to the milky sap of rubber

trees and certain plants. This sap was found to be an aqueous

medium containing colloids of natural rubber, stabilized by proteins.

Laboratory-synthesized polymer colloids were hence described as

“synthetic latexes,” and finally just “latexes.” That term will be used

here. Billions of pounds of latexes are synthesized worldwide each

year, for a large variety of applications. The reasons for interest in

their use in constructing electrochemical labels are


Latex as a Label Support 263

• The synthetic and physical chemistry of latexes has been

well-researched, including methods to control charge, size,

and hydrophobicity [50–52].

• It is possible to synthesize latexes with a very narrow

particle size distribution.

• A latex solution provides a large solid–liquid interfacial area

for modification.

• Many methods of chemical modification of latex are

available, and high surface concentrations of functional

groups can be achieved.

8.5.2 Latex Synthesis

Latexes can be synthesized by emulsion polymerization. Originally

this meant emulsifying an aqueous-insoluble monomer in water

with a surfactant and then using a water-soluble free radical initiator

to cause polymerization. The term emulsion polymerization is still

used, despite the fact that an emulsion is not always needed to

produce polymer colloids.

There are a huge number of methods available for latex

synthesis based on many industrial applications [53] and a thorough

review of that literature is beyond the scope of this chapter.

However, in general, the reaction mixture will contain one or more

monomers bearing double bonds capable of undergoing free radical

polymerization, water, emulsifier, (i.e., a surfactant), and an initiator

compound which will decompose to form free radicals. In batch

mode, all of the reactants are added together and heated to reaction

temperature. Hence, synthesis typically requires a heating bath

and a reaction flask with openings for a stirrer, reflux condenser

and an inlet and outlet for nitrogen (because oxygen is a free

radical inhibitor). A sampling device may also be useful to monitor

the reaction by extracting aliquots of the reactant over time. The

principle of the reaction is that the initiator compound decomposes

to form free radicals

I → 2R•

which then attack the monomer molecules to initiate chain growth

R + M → M•



Once a chain is started it propagates according to

Mj + M → Mj+1

As the length of the chain increases the molecule becomes

decreasingly water soluble, until it eventually comes out of the

solution and forms a primary particle. Thermal motion in the

solution causes collision between primary particles, leading to

coagulation and fusion into larger particles. These are spherical

because interfacial tension acts to minimize the interfacial area.

Since the initiator free radical is typically a water-soluble ionic group

such as SO3− or OSO3

−, it imparts a charge to the primary particle.

As primary particles coagulate, the surface charge density of the

growing sphere increases. This leads to electrostatic repulsion,

slowing and eventually stopping further coagulation. It is often easy

to produce latexes of a very narrow size distribution, described as

“monodisperse.” Synthetic methods appropriate to the construction

of micron and sub-micron sized latex electrochemical labels include

the synthesis of polystyrene (PS) latex colloids [54], which are then

present during the synthesis of polystyrenesulphonate (PSS) [55],

leading to a negative PSS shell around the PS core; the copoly-

merization of styrene and acrylic acid to produce a polystyrene-co-

acrylic acid (PSA) coploymer [56], which has a negative charge due

to acrylic acid deprotonation. Other than sulphonate and sulphates,

functional groups which can be introduced to the latex by the

initiator include alcohols, carboxylic acids, and =NH2+ [57].

8.5.3 Latex Solution Properties

In solution the latex spheres will experience van der Waals forces

of attraction, which at a separation r will be proportional to r−6.

For coagulation to not occur, these forces must be balanced by the

repulsive electrostatic force arising from either the ionic functional

groups on the latex, or adsorbed ionic surfactant. Hence, a latex

particle in an electrolyte will support a tightly bound layer of one ion

balanced by a diffuse layer of an oppositely charged ion. This diffuse

layer is equivalent to the diffuse layer at an electrode-solution

interface and so can be described by Gouy-Chapman theory [46].

Therefore, the width of the diffuse layer will be equal to the Debye



length κ−1, where [46]

κ =√

2n0z2e2

εε0kT(8.10)

in which n0 and z are the concentration and charge of ions in the

electrolyte, e is the charge on an electron, ε is the permittivity of the

medium, ε0 is the permittivity of free space, and k is Boltzmann’s

constant. The overall interaction between the latex particles is then

the sum of their attractive and repulsive forces, and is described

quantitatively by DVLO theory [58, 59]. The important experimental

parameter here is the electrolyte concentration, since this does not

effect van der Waals forces but when increased causes the diffuse

layer to shrink (e.g., from ∼300 A to 3 A going from 1 × 10−4 M to

1 M for a 1:1 electrolyte at 25◦C [46]). Thus, increasing electrolyte

concentration can cause coagulation. (This should also be noted

for solution phase nanoparticles since the physical principles are

exactly the same.)

8.5.4 Layer-by-Layer Deposition: Theory

In 1966, Iler demonstrated that films of alternating positively

charged alumina fibrils and negatively charged silica colloids could

be built up on hydrophilic glass [60]. In the early 1990s, Decher and

coworkers extended this procedure to the deposition from solution

of oppositely charged polyelectrolytes [61–63]. The technique,

known as “layer-by-layer” (l-b-l) deposition has since become widely

applied. The method is relatively simple, and as shown in Fig. 8.8,

consists of (1) derivatizing a substrate with a stable surface charge

excess, (2) immersing the substrate in a solution of an oppositely

charged polyelectrolyte (PE), (3) immersing in water to remove

weakly bound PE, and (4) immersing in a solution of a second PE,

oppositely charged to the first. Steps (2) to (4) can be repeated as

many times as necessary to give the required thickness. The reasons

for the popularity of the method are that, in addition to simplicity,

it allows us to control the resulting film thickness down to the level

of a few Angstroms, films of more than 1000 PE layers are possible,

the films are physically stable and are permeable to solution species,

enabling a film-confined catalyst to react with substrate. The l-b-l



Figure 8.8. Schematic of the layer-by-layer deposition process on a

substrate bearing an initial negative charge excess. Taken with permission

from [134], M. F. Castelnono and J.-F. Joanny, Langmuir 16, 7524–7532

(2000). c© American Chemical Society.

technique has been applied to the deposition of many different

charged species, including conducting polymers, DNA, and proteins.

Some recent reviews of the applications are given in Refs. 64 to 66.

The main driving force for the adsorption of, for example, a

positive PE onto a negative surface is electrostatic attraction. Zeta

potential measurements of such adsorption [67] have shown that

charge overcompensation occurs, that is, the PE/solid does not

become neutral, but is positive overall and so can then adsorb a

negative PE. As the layers are built up, the zeta potential oscillates

symmetrically around the zero value [67]. Neutron reflectommetry

experiments indicate the polymer layers are not flat, but penetrate

into each other [68]. Apart from Coulombic attraction, secondary

forces such as van der Waals, hydrogen bonding and hydrophobic

interactions also contribute, and these attractions give the process a

negative enthalpy. Also, small counterions and solvent shell water

molecules are liberated when the PEs come together and hence

entropy is increased. These two factors are responsible for the

negative free energy of l-b-l deposition according to �G = �H –

T�S .



Oppositely charged PEs can also form complexes in solution

[known as “interpolyelectrolyte complexes” (IPEC)] [69, 70]. Solid-

state nuclear magnetic resonance spectroscopy has shown that

these are structurally similar to l-b-l films, and hence l-b-l films

can be thought of as stacked layers of IPECs. Interestingly, IPEC

formation is almost entirely driven by the entropy increase [71],

which suggests that forces other than electrostatic attraction may

be used to form l-b-l films. This has been demonstrated for

hydrogen bonding [72–73], hydrophobic interactions [74], and DNA

hybridization [75, 76].

Where electrostatic attraction is used for l-b-l film formation,

increasing the ionic strength of the solution will generally increase

the film thickness [77, 78]. This is thought to be because a higher salt

concentration increases the shielding around the ionic groups of the

polyelectrolyte, causing it to adopt a more coiled, compact form [71,

79, 80].

8.5.5 Layer-by-Layer Modification of Latex

8.5.5.1 Latex surface charge excess

To modify latex spheres by the l-b-l method, there must be a

stable charge excess on the colloid surface (also required to prevent

coagulation). In the case of PS latex commercial samples are

available, from suppliers such as Sigma, bearing sulphate groups.

Otherwise, those groups can be imparted by synthesizing PSS in the

presence of PS [55]. PSS is a strong electrolyte and therefore can be

expected to be fully dissociated. In the case of PSA latex copolymers

[56] the negative charge will be dependent on the polyacrylic acid

(PAH) deprotonation and therefore on the contacting pH. This

deprotonation was studied in detail recently for sub-micron PSA

spheres by Li et al. [81]. The dissociation proceeded as

(latex) − COOH + OH− → (latex) − COO− + H2O

Based on UV absorbances measured after latex dissolution, it was

found that the PAH:PS ratio in the solid was 0.34, resulting in

latexes that were highly hydrophobic. This meant that deprotonation

only extended approx. 1.5 nm into the sphere, which for the

diameter of 0.265 μm means that of the total PAH in the particle



only approximately 3% was dissociated. The rate constant for

PAH dissociation on the latex decreased with increasing pH, as

expected, and was one order of magnitude lower than the value for

dissociation of PAH in solution at the same pH. This was attributed

to the hindering effect of neighboring −COO− groups on the latex

surface.

(a) Electrochemical labels by adsorption

Thus far, most of the studies of l-b-l material loading onto latex

spheres has focused on the layered deposition of biological macro-

molecules such as hemoglobin [82], DNA [75, 76], immunoglobulin

G [83], or enzymes such as glucose oxidase [84], horseradish

peroxidase [84], urease [85], and tyrosinase [55]. The layers were

deposited typically on PS latex stabilized by negative surface groups.

The available charge of the biological molecules, at any pH other

than the isoelectric point, meant that the deposited species could

replace one of the polyelectrolytes. To modify latex colloids the l-

b-l process is performed by adding the polyelectrolyte or biological

molecule to a colloidal suspension of the charge excess latex. After

20 min to 1 h (often under stirring), the suspension is centrifuged

down to a pellet and the solution decanted off to be replaced by

water. The colloids are then redispersed into the water by vortex

shaking. The centrifugation/water redispersion is performed twice

more. This provides the rinsing step noted previously to remove

weakly bound material. An oppositely charged material can then be

incubated with the latex in the same manner. Overall, the process is

simple and almost as reproducible as the modification of a planar

surface (allowing for a possible size distribution of the latex). Where

zeta potentials have been measured [83–85], the oscillating values

characteristic of l-b-l deposition have been found.

To the best of our knowledge, the previous l-b-l nanoparticle

modifications of latex have all been directed at incorporation into

the walls of hollow capsules, as described below. However, if metal

nanoparticles are stabilized by a surface charge then they can

be adsorbed to appropriately modified l-b-l latex by electrostatic

attraction, as shown in Fig. 8.1b. In this figure, the gold nanoparticles

were produced by citrate reduction and so had a negative charge



due to the adsorbed citrate. The particles were attached in a

manner analogous to polyelectrolytes: a dispersion of nanoparticles

was added to the latex suspension and then incubated at room

temperature for 30 min. The modified latex was then isolated by

filtration with a membrane (pore size 0.2 μm) which would admit

the unattached gold particles (mean diameter = 15.5 nm ± 1.6 nm),

but not the latex (mean diameter = 0.338 μm and 0.493 μm). We

found that the 0.493-μm latex had a higher gold coverage. Latex

particles have also been modified by coating with streptavidin and

then attached to biotin-coated gold nanoparticles via the strong

avidin-biotin bond [86]. However, the method gave a nanoparticle

coverage of 1 order of magnitude less than l-b-l deposition. This

corresponded to 2 orders of magnitude less metal ions released, due

to using gold nanoparticles of a smaller size.

(b) Electrochemical labels from hollow capsules

Capsule Formation: The l-b-l based construction of hollow capsules

was developed mainly as a technique for achieving localized drug

delivery, since the capsule can protect the drug from degradation

by the body. Recent thorough reviews of capsule formation and use

are available [87–91]. As shown in Fig. 8.9, there are three general

methods of constructing the capsules: (A) loading a preformed

capsule, (B) encapsulating crystals by l-b-l assembly, and (C)

incorporation into a porous sphere which is then coated by an l-b-l

process. The construction of nanoparticle electrochemical labels

from latex is based on method (A) and so only that will be

discussed further. The preformed capsule used in (A) is made by

l-b-l deposition onto an organic core which is then dissolved, as

shown in Fig. 8.10. Typical latexes which have been used for the

core are PSS and PSA, both of which have a negative surface charge,

as explained earlier, that can be utilized for l-b-l modification. Both

cores can be dissolved by THF. Other organic cores which have

been used for this method are melamine formaldehyde (MF), which

dissolves at low pH, and polylactic acid (PLA) or polylactic-co-

glycolic acid (PLGA), which can be dissolved in acetone/N -methyl-

2-pyrrolidinone mixtures. It should be noted that in this type

of capsule formation, the incomplete removal of core material is



Figure 8.9. Different methods of capsule construction and filling:

(a) Loading a preformed capsule by reversible pore formation. (b) Forming

a capsule by l-b-l deposition onto a crystalline material. (c) Loading a porous

sphere which is then coated by an l-b-l process. Taken with permission from

[88], A. P. R. Johnston, C. Cortez, A. S. Angelatos, and F. Caruso, Curr. Opin.Colloid Interface Sci. 11, 203–209 (2006). c© Elsevier Ltd. See also Color

Insert.

Figure 8.10. Schematic of capsule formation and loading with silver

nanoparticles by pH adjustment. Taken with permission from [105], P.

Rijiravanich, M. Somasundrum, and W. Surareungchai, Anal. Chem. 80,

3904–3909 (2008). c© American Chemical Society. See also Color Insert.

sometimes an issue. For example, MF-originated hollow capsules

have been found to contain MF at up to 30% of the total capsule mass

[92].



Capsule Permeability: The most widely characterized polyelec-

trolytes in the formation of hollow capsules have been alternating

layers of polystyrenesulphonate (PSS, negative) and polyallylamine

hydrochloride (PAH, positive), and therefore the following perme-

ability discussion will be based on that system. (PAH/PSS)n will refer

to a film of n bilayers.

The relatively loose, layered structure of the PEs renders

them porous to low-molecular-weight compounds. (When those

compounds are charged it has been suggested that their movement

through the shell is by “hopping” from oppositely charged sites [93].)

The shell porosity has been examined by entrapping fluorescein

microparticles at low pH and then measuring the fluorescence in

bulk after the microparticles are dissolved through a pH increase

[94]. It was found that the permeability to small molecules

decreased with increasing film thickness. For more than 8 layers,

the decrease was roughly linear with the film thickness increase

and corresponded to a diffusion coefficient of fluorescein through

the shell wall in the order of 10−12 cm2 s−1. For less than 8 layers,

the shell permeability decreased more quickly than described by

a linear relation, which is consistent with the finding that the first

eight layers have a more dense conformation than the subsequent

coatings [95].

Effect of ionic strength. As noted earlier, the initial structure of the

PE layers is affected by ionic strength. After hollow capsules are

formed from the PEs, they are also affected by the ionic strength

[96–98]. The exact reasons for the effect of ionic strength are

complex (see Ref. 91 for a detailed discussion), but in general

permeability increases nonlinearly with salt concentration. Human

serum albumin (HSA) has been incorporated into PSS/PAH capsules

by increasing the bulk NaCl concentration to 5 mM [99].

Effect of pH. If ionic strength changes cannot render PAH/PSS cap-

sules permeable to larger species (e.g., macromolecules, enzymes,

nanoparticles), then manipulation of pH or solvent polarity can

be used. The point about the (PAH/PSS)n system is that PSS is a

strong polyelectrolyte and remains fully ionized, whereas PAH is a

weak polyelectrolyte and so its dissociation is dependent on pH.



Therefore, the effect of pH on the capsule can be understood by

considering the effect of pH on the PAH layers. Those layers are

protonated according to the equilibrium

R − NH2 + H3O+ � R − NH3+ + OH−

The protonation has two effects: (1) Mutual repulsion from

neighboring −NH3+ sites causes a “stretching out” of the molecule.

Simulations on commercial software suggest the PAH length

increases by 7% from uncharged (pH 10.0) to fully charged (pH 3.0)

[91]. (2) The formation of −NH3+ requires charge compensation

by counterions. Each counterion is surrounded by a shell of H2O

molecules and their entry into the film causes osmotic pressure

between the PEs. These two factors combine to result in an opening

up of the film structure. The opening has been observed by scanning

force microscopy [93]. Capsules exposed to acidic solution exhibited

pores of up to 100 nm diameter, while capsules at pH 10.0 showed

no such effect. The same thing has also been observed for the

PAH/PSS system deposited on a planar surface [100]. Importantly,

when capsules from an acidic solution were transferred to a solution

at pH 10.0 the pores could not be observed [93]. Hence, the capsule

opening is reversible, and so pH may be manipulated to entrap large

molecules within the capsules. Such entrapment has been studied

by confocal microscopy using fluorescent-labeled dextran, and it

was demonstrated that in acidic conditions dextrans entered the

capsules [91]. Polyions and proteins have also been entrapped by

this method [101].

Effect of solvent polarity. Solvent polarity affects capsule perme-

ability by changing the solubility of the capsule walls. In the case

of PAH/PSS pairs, they are insoluble in water and soluble in ethanol.

Hence, varying the water-to-ethanol ratio of the suspension medium

can lead to a loosening of the film structure. At 20% ethanol content

a significant increase in the shell permeability was noted for a range

of high-molecular-weight materials including dextrans and proteins

[102]. As with pH-induced changes, the opening was found to be

reversible and could therefore be used for encapsulation.

Maximum loading. If the loading of the PE capsules is driven

solely by the concentration gradient across the capsule walls,



then we can expect the concentration of the loaded compound to

eventually become equal to its concentration in bulk solution. This

was found to be true for fluorescent-labeled HSA as quantified by

confocal microscopy [99]. However, in some cases enzymes have

been encapsulated at an internal concentration of over three orders

of magnitude greater than the bulk value [103, 104], and we have

incorporated silver nanoparticles at an effective concentration four

orders of magnitude greater than the bulk value [105]. This suggests

an additional driving force for encapsulation, possibly adsorption to

either the inner capsule walls or to undissolved core material.

Incorporation of nanoparticles: Metal nanoparticles were first

incorporated into the shells of hollow capsules [106] in order

to trigger light-assisted opening of the capsules [107–109]. This

technique was directed at the localized delivery of drugs at a

high dosage, the concept being that illumination in the near-IR

wavelength would cause heating of the nanoparticles and thus

degrade the shell walls. So far, hollow shells have been modified

by silver [108, 110, 111], gold [107, 109, 112, 113], and palladium

[110]. The modification has been performed by (a) depositing

(PAH/PSS)2 onto a latex core, reducing Ag+ onto the layers, then

depositing a further (PAH/PSS)2, followed by core dissolution [106,

108], or (b) forming (PAH/PSS)n shells by core dissolution and then

incubating with metal nanoparticles to allow adsorption, followed

by deposition of a further PSS layer [109].

We have found that Ag nanoparticles can be entrapped con-

veniently in (PAH/PSS)4 shells by pH manipulation, as shown in

Fig. 8.11 [105]. Because previous research was directed at light-

assisted capsule opening, there has not, to the best of our knowledge,

been any attempt to quantify the nanoparticle loading of the

capsules. However, as described below, this can be achieved to an

order of magnitude accuracy via voltammetry, UV absorbance, and

TEM. Using this process we estimated our loading as approximately

78 silver nanoparticles per capsule. From the mean size of the

nanoparticles (diameter = 15.8 nm), this corresponds to the release

of 9 × 106 Ag+ ions after acid dissolution. To determine the

distribution of the nanoparticles, we applied the same method

of quantification to nanoparticles adsorbed onto glass cover slips



Figure 8.11. TEM images of hollow (PAH/PSS)4 capsules obtained after

core dissolution by THF (a) and (PAH/PSS)4 capsules loaded with silver

nanoparticles (b). Taken with permission from [105], P. Rijiravanich, M.

Somasundrum, and W. Surareungchai, Anal. Chem. 80, 3904–3909 (2008).

c© American Chemical Society.

coated singly by PAH and by PSS, as well as by a (PAH/PSS)3PAH

coating. The results indicated the distribution was approximately

70% on the inner wall, 17% on the outer wall, and 13% intercalated

between. This is reasonable given that silver nanoparticles have an

isoelectric point of 2.7 [114], and therefore possess a negative zeta



potential at the pH used for encapsulation. Thus they adsorb most

strongly to the positive PAH surface.

Nanoparticle Quantification: To optimize the nanoparticle loading of

a particular electrochemical label it is necessary to have a means

of determining that loading. This can be done in a systematic way

based on UV absorbance, voltammetry, and TEM measurements, as

outlined below.

Nanoparticle recovery. The nanoparticle recovery is the proportion

of the initial metal ions that are converted into metal nanoparticles.

This can be calculated by first determining the mean nanoparticle

radius from TEM. Most methods of synthesis will produce a

distribution of radii, as shown for the silver particles in Fig. 8.12;

this distribution represents the main error in the determination.

Based on the mean radius and the bulk density value from literature,

we can calculate the mean mass of 1 nanoparticle. Since we know

Figure 8.12. TEM image of silver nanoparticles at pH 6, synthesized

by NaBH4 reduction of AgNO3 Inset: particle size histogram from >100

particles. Taken with permission from [105], P. Rijiravanich, M. Somasun-

drum, and W. Surareungchai, Anal. Chem. 80, 3904–3909 (2008). Supporting

Information. c© American Chemical Society.



the metal ion concentration used in the synthesis, we can calculate

the total mass of the nanoparticles assuming a 100% conversion,

[M ]100%. The actual mass of nanoparticles, [M ]exp, can then be

determined by acid-dissolution of a known aliquot followed by ASV

analysis, having plotted a calibration curve for that metal ion. The

recovery is obviously [M ]exp/ [M ]100%.

Nanoparticle stock concentration The total number of nanoparti-

cles is given by

mass of metal ion used in synthesis

x mass of 1 particle× recovery = no. of particles

From the volume used in the synthesis this can be converted to a

concentration of nanoparticles mL−1.

Capsule/latex concentration. A TEM image of some (PAH/PSS)4

capsules is shown in Fig. 8.10a, and a sub-micron PSA latex particle

shown in Fig. 8.13. The mass of one capsule or particle can be

calculated from the shell or particle dimensions, assuming a density

of 1.01 g cm−3 for the capsule and 1.05 g cm−3 for the particle [55].

Figure 8.13. TEM image of 493 nm diameter PSA particle. Taken with

permission from [131], S. Pinijsuwan, P. Rijiravanich, M. Somasundrum,

and W. Surareungchai, Anal. Chem. 80, 6779–6784 (2008). Supporting

Information. c© American Chemical Society.



By determining the dry weight of a known aliquot deposited onto a

glass slide, we can then calculate the concentration per mL.

Nanoparticle loading. The nanoparticle suspension should have a

UV/vis absorbance maxima. Since the nanoparticle stock concen-

tration is now known, this maxima can be used for calibration

as shown in Fig. 8.14 for 15.8-nm diameter silver nanoparticles

Figure 8.14. (a) Absorbance spectra of silver nanoparticles shown in

Fig. 8.12 (b) Calibration of silver nanoparticles from absorbance at 406 nm

following determination of stock concentration. Taken with permission

from [105], P. Rijiravanich, M. Somasundrum, and W. Surareungchai,

Anal. Chem. 80, 3904–3909 (2008). Supporting Information. c© American

Chemical Society.



(λmax = 406 nm). Hence, after take-up of silver particles by

the capsules/latex particles, the nanoparticles remaining in the

solution can be separated after centrifugation and the remaining

concentration determined. Knowledge of the initial concentration

used enables us to calculate the number of nanoparticles taken up.

Knowledge of the number of capsules/latex particles allows us to

calculate the capsule/latex loading.

8.6 DNA Measurement

The nanoparticle labels can be used to detect DNA following the

general stages: (1) Attachment of DNA probe or target to the

electrode, (2) attachment of DNA probe or target to the label, (3)

hybridization to form a duplex, (4) dissolution of the metal ions in

the label (50% HNO3 for Ag dissolution, 1 M HBr/0.1 mM Br2 for

Au), and (5) detection of the metal ions.

DNA probes are usually in the range 12–40 base pairs. Above 40

base pairs, folding of the probe on the electrode is likely to lower

hybridization efficiency by steric hindrance. Also, at such lengths

the degree of binding to partial mismatches may be significant. At

below 12 base pairs the probe is unlikely to be unique to a particular

sequence.

8.6.1 DNA Immobilization

DNA can be immobilized on the electrode by either covalent linking

or physical adsorption. DNA modified by a thiol group can be

chemically attached to gold electrodes [115–117] following the

formation of the sulphur-gold bond:

DNA − SH + Au → DNA − S − Au + e− + H+

Alternatively, the gold electrode can be modified with a thiol-based

self-assembled monolayer (SAM) bearing functional groups suitable

to bind DNA [118]. Often this binding is performed via a coupling

reagent such as 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide

(EDC), which enables aminated or carboxylated DNA to bond with

the appropriately carboxylated or aminated functional group on

the electrode [119], or on a polymer deposited on the electrode


DNA Measurement 279

[120]. Thiol-based linkage to the DNA means that at high coverages

the DNA is oriented normal to the electrode. If we approximate

the oligonucleotide molecule to a cylinder, then the maximum

possible loading is defined by the diameter of the cylinder. This

value is 20 A [121]. The coverage from such immobilization can be

calculated using the chronocoulometric method described by Steel

et al. [121]. The principle of the method is that [Ru(NH3)6]3+ is

used to compensate the negatively charged phosphate groups of

the DNA under conditions of low supporting electrolyte. Therefore,

when DNA is immobilized at the electrode surface the concentration

of [Ru(NH3)6]3+ is increased. The coverage Γ of DNA-bound

[Ru(NH3)6]3+ is determined by stepping the potential to a value

where [Ru(NH3)6]3+ is reduced at a diffusion-limited rate. From the

integrated form of the Cottrell equation [46], a plot of charge Qagainst

√t will have a y-intercept equal to Q dl + nF AΓ , where Q dl

is the double-layer charge, determined from the same potential step

in the absence of [Ru(NH3)6]3+. There are Γ NA molecules/cm2 of

[Ru(NH3)6]3+ on the electrode, where NA is Avagadro’s number, and

therefore, assuming each [Ru(NH3)6]3+ molecule is compensated

by three phosphate groups, and there are m phosphate groups

on one DNA probe, there are (3/m)Γ NA DNA probes/cm2 on the

electrode. The technique can be applied to ss and dsDNA and

thus the hybridization efficiency can be determined. Some studies

have suggested that the efficiency decreases with DNA coverage

[122]. When thiol-modified DNA is immobilized on gold, a “diluent”

alkanethiol is often also adsorbed to displace weakly bound DNA

bases. In these cases, the chronocoulometric method has indicated

that hybridization efficiency increases with DNA length above the

diluent layer [123].

A much simpler method of immobilization is direct adsorption,

in which case we would expect the DNA to be oriented horizontally

along the electrode. Therefore, the maximum coverage will be

determined by the number of layers it is possible to deposit.

Forces such as hydrogen bonding, base stacking, van der Waals and

hydrophobic interactions are expected to be involved [124]. Due to

the negative charges of the phosphate groups, the adsorption can be

assisted by electrostatic attraction. Glassy carbon electrodes have

been used to bind DNA after modification by the cationic polymer

chitosan [120]. Another way to assist electrostatic attraction is to



hold the electrode at a positive potential. AFM studies of ssDNA

immobilized on a pyrolytic graphite electrode showed that holding

the electrode at 300 mV (vs. Ag wire) increased the film thickness

of the adsorbed ssDNA film from 0.98 ± 0.40 nm (open-circuit

adsorption) to 2.37 ± 0.4 nm, which suggests that at a positive

potential more than a single monolayer was adsorbed [124]. The

electrode was almost completely covered, with very few holes.

In the case of screen-printed carbon electrodes, we have used a

mildly positive potential (100 mV vs. AgCl screen-printed track

for 30s) which produced a strong adsorption, such that the DNA

remained adsorbed after washing. This method is attractive since

the electrodes are disposable and, as noted earlier, it means only a

small solution volume is needed.

8.6.2 Probe Attachment

A convenient method of attaching latex-based labels to DNA is the

avidin (or streptavidin)-biotin system, which has been widely used

[125–130]. DNA sequences with a biotin tag at the 5’ end are

commercially available. Avidin and streptavidin are proteins which

possess a high binding affinity for biotin (K a = 1015 M−1) and can

be adsorbed onto labels by incubating the labels in an appropriate

solution (e.g., in 3 mg mL−1 of protein for at least 15 min). Uptake

of the protein can be monitored by centrifuging down the solid and

then decanting off the liquid. A reduction in protein absorbance

at 280 nm confirms uptake onto the label. The main difference

between the two proteins is in the value of the isoelectric point (5

for streptavidin and 10.5 for avidin), and in the fact that streptavidin

is much more expensive. In labeling hollow capsules and latex we

used pH values that would render avidin positive. This facilitated

adsorption to the negative PSS outer layer of the capsules. In the case

of adsorption to gold-modified latex, we expect the main location of

the avidin to be on the negatively charged gold particles, since the

PAH latex outer layer is positive.

8.6.3 Detection Sequence

The simplest detection scheme is to use a single probe to detect

the target, with either the target or the probe being labeled. This


DNA Measurement 281

Figure 8.15. Scheme of the DNA hybridization detection procedure using

the Au nanoparticle–coated latex labels shown in Fig. 8.1 [131], taken with

permission from S. Pinijsuwan, P. Rijiravanich, M. Somasundrum, and W.

Surareungchai, Anal. Chem. 80, 6779–84 (2008). c© American Chemical

Society. See also Color Insert.

form of detection was used by us to quantify latex-based labels,

following target immobilization, as shown in Fig. 8.15. While this

system is convenient if the experimental objective is to develop the

construction of the labels, it is not an ideal method for real samples.

As shown in Fig. 8.16 target sequences of a one base mismatch can

give a significant response. Since we would expect some mismatched

sequences in the sample to be immobilized also, this would provide

interference. A technique to minimize this form of interference is to

use two probes for one target, as shown in Fig. 8.17. A capture probe

is immobilized on the electrode, and then hybridized to one section

of the target. A signal probe, carrying the label, is then bound to a

remaining section.

Using the single-probe method, a 30-base sequence common to

five strains of E. coli could be detected using the latex-based labels,

with detection limits of ∼25 fM (silver nanoparticles on hollow

capsules) [105] and ∼0.5 fM (gold nanoparticles on latex) [131], as

shown in Fig. 8.18. The lower detection limit for the second method



Figure 8.16. LSV detection of DNA hybridization via the silver-loaded

capsules shown in Fig. 8.11, using 200 fM of target complementary to the

probe (a), 200 fM of target containing a single mismatch (b), and 60 pM

of a non-complementary target (c). LSV: Edep = −0.5 V, tdep = 500 s, scan

rate = 50 mV s−1. Taken with permission from [105], P. Rijiravanich, M.

Somasundrum, and W. Surareungchai, Anal. Chem. 80, 3904–3909 (2008)

Supporting Information. c© American Chemical Society.

Figure 8.17. Scheme of DNA detection by sandwhich assay. See also Color

Insert.


DNA Measurement 283

(

Figure 8.18. Detection of DNA hybridization via (A) the silver-loaded

capsules shown in Fig. 8.11 using LSV and (B) the gold-loaded latex

spheres shown in Fig. 8.1 using DPV. Taken with permission from [105],

P. Rijiravanich, M. Somasundrum, and W. Surareungchai, Anal. Chem. 80,

3904–09 (2008). American Chemical Society and [131], S. Pinijsuwan,

P. Rijiravanich, M. Somasundrum, and W. Surareungchai, Anal. Chem. 80,

6779–84 (2008). c©American Chemical Society.



is from a combination of a higher nanoparticle loading and the use

of DPV for quantitation, instead of LSV.

8.7 Areas for Further Research

It is hoped that this chapter has transmitted two general points:

that it is relatively straightforward to adapt latex colloids for use

as electrochemical labels, and that very little has been done in this

field up to now. Some possible further directions for research are as

follows:

1. Increasing the nanoparticle loading on the latex spheres via the

autocatalytic metal deposition previously described [35, 36].

2. Increasing the nanoparticle loading on the hollow capsules by

finding a way to load the central volume of the capsules, rather

than just the capsule walls.

3. Applying either latex or capsule labels to multianalyte detection

by preparing labels loaded with different metals.

4. Designing a cell arrangement to reduce the electrolyte volume

needed for ASV. This would increase sensitivity by increasing the

concentration of the liberated metal ions.

5. Extending the use of latex-based labels to the analysis of real

samples.

It should also be noted that many of the previously reported latex

l-b-l modifications have described the deposition of layers of redox

enzymes [55, 84, 85], and hence these structures could also be

used as labels. Hollow capsules have also been used to entrap

enzymes [101]. While enzyme stability can sometimes be an issue,

the sensitivity provided by enzymes is often very good. For example,

l-b-l deposition of alkaline phosphatase onto carbon nanotubes

resulted in electrochemical DNA sensing down to 5.4 aM [132]. In

comparison with a nanotube, a latex sphere of diameter ∼0.5 μm

presents a very much larger surface area for immobilization. Finally,

virtually everything stated in this chapter regarding DNA labeling

can equally be applied to the labeling of antibodies.


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Acknowledgments

The authors would like to thank Chatuporn Phantong for assistance

in preparing the figures in this chapter.

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Chapter 9

Screen-Printed Electrodes forElectrochemical DNA Detection

Graciela Martınez-Paredes, Marıa Begona Gonzalez-Garcıa,and Agustın Costa-GarcıaDepartamento de Quımica Fısica y Analıtica, Facultad de Quımica,Universidad de Oviedo, Julian Claverıa s/n, 33006 Oviedo, Asturias, [email protected]

The concept of DNA biosensors is sustained by the need for rapid

and highly sensitive analytical tools for genetic detection. Their

implementation is based on three steps: (i) immobilization of

single-stranded oligonucleotide (probe) onto a transducer surface;

(ii) hybridization with its complementary DNA sequence (target) in

order to form the DNA duplex called hybrid, and (iii) conversion of

the hybridization event into an analytical signal by the transducer

surface. A wide variety of measurement systems had been employed

[1], however, since Palecek discovered the electrochemical activity

of nucleic acids [2], the electrochemical studies on the behavior and

recognition of DNA have attracted considerable attention. In this

way, electrochemistry provides fast, simple, and low-cost detection

systems to produce biosensors promising a simple, accurate, and

inexpensive platform for patient diagnosis [3–6].



292 Screen-Printed Electrodes for Electrochemical DNA Detection

9.1 Introduction

Although numerous DNA hybridization assays have been routinely

used in diagnostic laboratories, there is a growing interest in

screen-printed DNA-hybridization sensors, because these can be

mass-produced by existing manufacturing processes at low cost.

Nowadays, screen-printed electrodes (SPEs) are being developed as

a suitable tool for electrochemical analysis because of their unique

properties such as small size, low detection limit, fast response time,

and high reproducibility. Furthermore, screen-printing technology

is a well-established technique for the fabrication of biosensors. It

has been exploited commercially in the production of these devices,

most notably, the personal glucose biosensor used by diabetics

[7]. In addition, many research laboratories in universities possess

screen-printing facilities for in-house production of sensors for

prototype devices.

9.2 Fabrication of Screen-Printed Electrodes

Summarizing, the process consists in forcing a conductive ink to pass

through a screen which is placed on a material that acts as support.

The screen only allows the pass across a few pores that define the

form and dimensions wished for the electrode, staying hereby an

image of the same one printed on the support.

The screen printing process uses a porous mesh stretched tightly

over a frame made of wood or metal. Fig. 9.1 The mesh is made

of porous fabric or stainless steel. A stencil is produced on the

screen either manually or photochemically defining the image to

be printed. Thus, the design of the stencil allows to obtain a

wide range of screen-printed electrodes in which the electrodic

configuration, as well as the size and form of these electrodes can be

controlled.

A great variety of inks are commercially available, but they can

also be made in order to attend to specific characteristics. The

ink generally contains a binder agent such as glass powder, resins,

cellulose acetate, or some solvents, and additives that provide the

wished functional characteristics. Screen printing ink is applied to

the substrate by placing the screen over the material and the ink


Fabrication of Screen-Printed Electrodes 293

Figure 9.1. Schematic representation of the screen-printed electrodes

production process.

onto the top of the screen. Ink is then forced through the fine mesh

openings using a squeegee that applies pressure. After every stage

of printing a series of drying stages to eliminate solvents, and a final

cured step to a certain temperature.

Finally, the support is covered with an insulating layer leaving

uncovered only the electrode area and the electrical contacts.

9.2.1 Types of Screen-Printed Electrodes

As it has been mentioned in the previous section, due to the

versatility of the production process of screen-printed electrodes,

a wide range of SPEs can be made, containing only the working

electrode, working and counter electrodes to work with an external

reference electrode, a complete electrochemical cell, or even with

multiple working electrodes, for applications where a disposable

electrode is desired to perform electrochemical measurements

Fig. 9.2.

The most employed inks for the fabrication of screen-printed

electrodes are made of carbon, gold, platinum, or silver. Neverthe-

less, other materials can be easily used.

Gold or platinum are used in SPEs fabrication, avoiding the

use of a great quantity of these expensive materials. In this sense,

sometimes a narrow single-electrode sensor is used to replace metal

electrodes.



Figure 9.2. Some commercially available screen-printed electrodes show-

ing different electrode configurations. See also Color Insert.

Carbon and gold have a wide use in the technology of disposable

sensors as electrodic materials [8]. Gold has been employed as

electrodic material for the genosensors construction for years, and

carbon is especially used due to its great superficial chemistry,

its low background current, the wide potential window at which

it is possible to be employed, its low cost, and its chemical

passivity. Nevertheless, the electronic-transfer rate obtained with

carbon-based electrodes is lower than that obtained with metallic

electrodes [9].

However, this disadvantage can be overcome by means of the

surface modification of these electrodes with nanostructures, as the

use of carbon nanotubes (CNTs) [10], or gold nanoparticles [11],

since they improve the electronic transfer of the surface of the

electrode, and improve the analytical characteristics offered by the

sensor. Carbon nanofibers can also be used to modify the electrodic

surface in order to improve the analytical characteristics of the

transducer.

In addition, SPEs surfaces have also been covered with a wide

variety of substances: bismuth oxide, Prussian Blue, ferrocyanide,

Meldola’s Blue, Co-phthalocyanine, or some enzymes, in order to

obtain suitable transducers for specific analytes.

9.3 Genosensors on Screen-Printed Electrodes

DNA detection is usually performed by hybridization. For designing

a genosensor, the crucial steps are the choice of the transducer

surface and the immobilization of the single-stranded (ssDNA)


Genosensors on Screen-Printed Electrodes 295

probes onto electrode surface, because the molecular recognition

event typically occurs directly on the surface of the signal transducer.

The immobilization method will determine the sensitivity

and reproducibility of the genosensor. Several strategies for the

immobilization of ssDNA have been carried out and will be discussed

in section 9.3.2. The ssDNA probe immobilized on the transducer

surface recognizes its complementary (target) DNA sequence via

hybridization. The DNA duplex is then converted into an analytical

signal by the transducer. Different strategies for electrochemical

detection have been performed and are mainly divided in two

groups: methods using direct detection (those in which the intrinsic

electroactivity of DNA is involved) or indirect detection methods

(those which imply the use of labels).

Electrochemical detection of hybridization is mainly based

on the differences in the electrochemical behavior of the labels

with or without double-stranded (dsDNA) or single-stranded DNA

(ssDNA). The labels for hybridization detection can be enzymes,

anticancer agents, organic dyes, colorants, metal complexes, or

metal nanoparticles among others.

9.3.1 Electrochemical Detection of Hybridization Reaction

As it has been mentioned previously, there are a wide range of

possibilities for the electrochemical detection of the hybridization

reaction, and they can be divided into two types, direct or indirect

methods.

9.3.1.1 Direct transduction methods

Direct transduction relies on the measurement of physico-chemical

changes occurring at the recognition layer induced by hybridization

event. These methods are generally based on the oxidation

processes of guanine or adenine that occur in an oligonucleotide

when the hybridization reaction takes place [1–16]. This is because

the nucleobases present in the double strand are oxidized in a lower

extension than when they are forming a part of ssDNA, making the

analytical signal decrease, but at the same time the target strand

adds new bases increasing in part the analytical signal. This fact



gives rise to non-linear calibration plots. The alternative is to use

probes in which guanine bases have been replaced by inosine.

Then the analytical signal appearing with the hybridization and

background signals are negligible.

Another strategy to differentiate the signal of the single strand

from that of the double strand is based on the use of a protein

that binds specifically to the ssDNA, preventing the oxidation of the

guanine in single strands of DNA [16].

The great advantage of this type of detection is to avoid the use

of marks or indicators of hybridization, simplifying the experimental

procedure. However, the detection based on the electroactivity of

bases gives rise to a lack of sensitivity. Various proposals based

on the use of oxidation products of adenine as catalysts of NADH

oxidation [17], or those based in the use of mediators for the

oxidation of bases, with ruthenium complex [18, 19] or osmium

complex [20] have been proposed in order to get an amplification

of the signal and thereby improve the sensitivity.

However, these methods induce an irreversible process prevent-

ing multiuse and are limited by the adenine and guanine content.

9.3.1.2 Indirect transduction methods

Indirect transduction relies on the use of indicators or labels.

The first ones are based on the differences in the electrochemical

behavior of indicators that interact in a different extension with

dsDNA and ssDNA. The indicators for hybridization detection can

be anticancer agents, organic dyes, or metal complexes, and are not

generally covalently joined to DNA. The latter strategies include the

use of labels covalently joined to DNA such as ferrocene, enzymes, or

metal nanoparticles.

Use of indicators Indicators are electroactive compounds that

present different affinity for ssDNA and dsDNA; they used to be

anticarcinogenic agents, organic dyes, or metallic complexes.

Some metallic complexes like Ru(NH3)3+6 , [Fe(CN)6]3−/4−,

Co(phen)3+3 , or Ru(bpy)2+

3 , and some organic compounds like

methylene blue (MB) recognize the hybridization reaction. The

union takes place via electrostatic interaction with the hollows of the



helix, or being inserted selectively and reversibly in the dsDNA. Their

use has been widely studied from the pioneering works of Millan and

Mikkelsen [21] in the early 90s. Most of them have been reviewed in

the work of Lucarelli [8]. Compounds that join the hollows of the

double helix have major affinity for dsDNA than for ssDNA, so the

signal due to the indicator oxidation increases when hybridization

takes place.

Other indicators, such as daunomycin or cobalt complexes, act

as intercalators. The changes in the area or peak potential of the

indicator oxidation process are used as analytical signal [22, 23].

Nevertheless, MB is another indicator that joins DNA by means

of intercalation, but generates minor reduction signals when it is

joined to dsADN than when joined to ssADN, because the specific

interaction of the MB with guanine bases is lower in the dsADN.

The hybridization indicators present the great advantage of

avoiding the processes of DNA labeling. Nevertheless, the discrim-

ination between single and double strand used to be not very good.

In addition, a general problem is the high backgrounds obtained,

due to unspecific adsorptions of indicators. However, if a negative

potential is applied to the electrodic surface once finished the assay,

these adsorptions can be repelled, diminishing the background

signals.

Use of labels There are two types of labels that join DNA covalently:

electroactive and non-electroactive labels.

The electroactive labels most used in genosensing design are

ferrocene and its derivates [24–27] (the reversible oxidation process

of ferrocene can be detected by means of several electrochemical

techniques), osmium complexes [28], platinum complexes [29],

gold complexes [30, 31], and metallic [32–36] or semiconductor

nanoparticles [37]. Among the last ones, gold nanoparticles are

the most used, their detection can be carried out by means of the

measurement of resistance or capacitance changes, usually after

an amplification procedure with silver, or by means of the anodic

stripping voltammetry of Au(III) obtained after the nanoparticle

oxidation Fig. 9.3.

An original approach consists in the use of ssDNA probes

labeled with an electroactive marker, the hybridization inducing



Figure 9.3. Particle-based protocols for electrochemical detection of DNA.

Reprinted with permission from Elsevier [33]. See also Color Insert.

the disappearance of the electroactivity of the probe, and the

appearance of a new signal characteristic of the resulting duplex.

The most used non-electroactive labels have been the enzymes

owed fundamentally to their capacity of amplification of the ana-

lytical signal, providing a great sensitivity. Generally, the analytical

signal is based on a redox process of some enzymatic reaction

product. Enzymes can be joined directly to the DNA strand [38–43],

or toward the interaction (strept)avidin-biotin [44–50] Figs. 9.4 and

9.5, digoxigenin-antidigoxigenin antibody [51–53], or FITC-antiFITC

antibody [54–56] among others.

The wide use of enzymes as labels in affinity assays is due to

their aptitude to turn the hybridization reaction into a wide range

of detectable molecules. The most usual enzymes are phosphatase

alkaline (AP), horseradish peroxidase (HRP), or glucose oxidase

(GOD). All of them are relatively stable, cheap, and generally have

high conversion speed.

9.3.2 Strategies for Immobilization of ssDNA over SPEs

The skill of immobilizing the probe onto the transducer in a

predictable way while keeping its inherent target affinity intact is

crucial for the development of the genosensor. In addition, if probe

strands are tidy and orientated, it can determine the sensibility

and reproducibility of the genosensor. Thus, independently of every

particular probe, some general aspects must be considered. The

immobilization of the probe must preserve the ability of target

recognition.



Figure 9.4. Schematic representation of a impedimetric genosensor

(sandwich hybridization assay). Unmodified PCR products (b) were

captured at the sensor interface (a) via sandwich hybridization with the

surface-tethered probe and a biotinylated signaling probe. The biotinylated

hybrid (c) was then coupled with a streptavidin–alkaline phosphatase

conjugate (d) and finally exposed to the substrate solution (e). The bio-

catalyzed precipitation of an insulating product (f) blocked the electrical

communication between the gold surface and the [Fe(CN)6]3−/4− redox

probe (published by Elsevier in Ref. 50).

Figure 9.5. Scheme of an assay in which enzyme is incorporated through

biotin streptavidin interaction. Reprinted with permission from Elsevier

[49].



It is obvious that the immobilization protocol depends on the

transducer characteristics, nevertheless it is preferred to use robust

immobilization methods in order to avoid the probe desorption from

the sensor [57]. Thus, the retention in polymeric matrix, covalent

bonds on a functionalized surface, SAMs, and immobilization

through affinity reactions are the most successful methods at the

moment, because these strategies give place to an immobilization

across the ends of the probes in a tidy and orientated way.

In addition, these strategies allow to control the conformational

freedom of the probes and the space between chains by means

of the control of the superficial covering obtaining hybridization

efficiencies up to 100%.

The most of screen-printed electrodes employed as transducers

of genosensors are made of carbon or gold inks. Further sections

detail the most used probe immobilization strategies in these types

of electrodes.

9.3.2.1 Immobilization of ssDNA over carbon electrodes

Several strategies of DNA immobilization have been described onto

screen-printed carbon electrodes (SPCEs).

Although it is frequently used [58, 59], direct ssDNA immobiliza-

tion over bare carbon surface happens in a random and untidy way

due to the multiple interactions between the carbon surface and the

phosphate structure of DNA. DNA strands immobilized by physical

adsorption are not orientated and present a limited mobility, so the

hybridization reaction is hampered by stearic impediments Fig. 9.6.

Adsorption at controlled potential is generally carried out on

pretreated SPCEs [13, 15, 60, 61]. Nevertheless, in these cases the

probe strands are not totally accessible for their hybridization,

diminishing the genosensor efficiency.

Avidin [24], neutravidin [51], and streptavidin [54] have been

used to immobilize biotinylated DNA strands onto carbon electrodes

Fig. 9.7, but before immobilizing the probe the surface must be

blocked to avoid unspecific adsorptions like that of the components

of the genosensor.

Other strategies are based on the formation of a polymer, by

means of the electropolymerization of the probe modified with the



Figure 9.6. Scheme of the electrochemical adsorption of probes and

detection by direct and indirect methods. Reprinted with permission from

Elsevier [57].

Figure 9.7. Scheme of the avidin-streptavidin immobilization method,

and detection by using an electroactive indicator. Reprinted with permission

from Elsevier [57].



Figure 9.8. Steps of the sandwich-type assay: (1) The redox polymer

and the oligonucleotide probe are electrodeposited on the screen-printed

electrode (SPE); (2) the capture probe and the target are hybridized; (3) the

electrode-bound target and the HRP-labeled oligonucleotide are hybridized,

the HRP labels are in electrical contact with the redox polymer; and (4) the

electrocatalytic reduction current of H2O2 to water is measured. Reprinted

with permission from ACS [38].

chosen monomer, the electropolymerization of a monomer, and the

further covalent bond of the probe strand or the copolymerization

of the monomer in presence of the DNA probe Fig. 9.8 [38,

57, 62].

There is also the possibility of forming self-assembled mono-

layers (SAMs) of oligos functionalizing these with hydrophobic

groups.

9.3.2.2 Immobilization of ssDNA over gold electrodes

Generally, the DNA immobilization onto screen-printed gold elec-

trodes (SPGEs) is carried out by means of SAMs formation of

oligo modified with thiol groups [4–48]. Covered surface and

spacing of oligos can be controlled through the addition of a

short-chain alcanothiol that acts as a solvent [63], blocks the

unspecific adsorptions, and at the same time orientates the

probe strands improving considerably the hybridization reaction

efficiency.

SAMs formation provides a high stability to the genosensor

Fig. 9.9: it is possible to avoid the oxidation and break of

the sulphur-gold link storing genosensors in a dark and dry

place, remaining unaltered for up to 2 months [64]. In addition,

thiolated oligonucleotides SAMs present a great thermal stability,

not being affected by gradients of temperature of up to 70◦C

[65].


Applications 303

Figure 9.9. Scheme of the DNA self-assembled monolayer formation on

gold electrodes. Reprinted with permission from Elsevier [57].

9.4 Applications

In this section several examples of genosensors based on hybridiza-

tion event, which have been constructed on screen-printed elec-

trodes, will be described. One of them has been designed to detect

a 30-mer SARS (severe acute respiratory syndrome) virus sequence

whilst the others have been designed to identify the nucleic acid

determinants exclusively present on the genome of the pathogen

Streptococcus pneumoniae.

Although in most of them alkaline phosphatase and 3-indoxyl

phosphate are used as label and enzymatic substrate, respectively,

other label, a platinum (II) complex, will be presented and its

detection discussed. In all cases, synthetic target oligonucleotides as

well as three-base mismatch and one-base mismatch strands of the

pathogen Streptococcus pneumoniae or SARS virus are tested using

these genosensor devices. In addition, in the last application of this

section, the versatility of the SPEs design is very useful to carry out

the simultaneous determination of two bacteria causing pneumonia.



9.4.1 Enzymatic Genosensors on Streptavidin-ModifiedScreen-Printed Carbon Electrode

This section outlines the development of genosensors on screen-

printed carbon electrodes (SPCEs) for the identification of nucleic

acid determinants exclusively present in the genome of the

pathogen Streptococcus pneumoniae. Orientation of the strands in

the sensing phase is achieved by modifying the surface of the

electrode with streptavidin by physical adsorption followed by the

immobilization of biotinylated oligo probes. The physical adsorption

of streptavidin must be performed at a constant temperature above

the room temperature. Moreover, the electrode surface must be

previously electrochemically pretreated at an anodic potential in

acidic medium to improve its adsorptive properties. In this way,

reproducible, sensitive, and stable sensing phases are obtained [66].

The biotinylated oligo nucleic acid probes used in this work target

the pneumolysin (ply) gene. This target is randomly labeled with

the Universal Linkage System (ULS). This labeling system consists

of the use of a platinum (II) complex that acts as a coupling agent

between DNA strands and a label molecule, usually fluorescent. This

platinum complex is a monofunctional derivate of cisplatin (a potent

anticancer agent used in the treatment of a variety of tumors) that

binds to DNA at the N7 position of guanine with release of one Cl

ion per molecule of the complex. The label molecule used in this

study was fluorescein (FITC). Electrochemical detection is achieved

using two strategies. One of them is carried out using an anti-FITC

alkaline phosphatase-labeled antibody and 3-indoxyl phosphate (3-

IP) as enzymatic substrate of AP. The resulting enzymatic product is

indigo blue, an aromatic heterocycle insoluble in aqueous solutions.

Its sulfonation in acidic medium gives rise to indigo carmine IC, an

aqueous soluble compound that shows an electrochemical behavior

similar to indigo blue. Both 3-IP and IC have already been studied on

SPCEs [67, 68]. However, although these genosensors are stable and

sensitive devices for the detection of specific nucleic acid fragments,

the need of two additional steps to obtain the analytical signal

resulted in a large time-consuming analysis. This fact can be avoided

using the second strategy for detection. In this case the analytical

signal is directly obtained from platinum (II) complex, which is


Applications 305

deposited on the electrode surface. In presence of the platinum

on the electrode surface and after fixing an adequate potential in

acidic medium, the protons are catalytically reduced to hydrogen.

The current generated by this catalytic reduction can be measured

and increases with platinum concentration and consequently with

labeled target concentration.

Data presented here demonstrate the potential applicability

of SPCEs genosensors in the diagnosis of a human infectious

pulmonary disease. These electrochemical genosensors are stable

and sensitive devices for the detection of specific nucleic acid

fragments. Moreover, these devices allow the detection of a one-

base mismatch on the targets if adequate experimental conditions

are used

9.4.1.1 Genosensor design

Electrode pretreatment: 50 μL of 0.1 M H2SO4 are dropped on the

SPCEs and an anodic current of + 3.0 μA is applied for 2 minutes.

Then, the electrodes are washed using 0.1 M Tris buffer pH 7.2.

Adsorption of streptavidin: an aliquot of 10 μL of a 1× 1−5 M

streptavidin solution is left on the electrode surface overnight at

4◦C. Then, the electrode is washed with 0.1 M Tris buffer pH 7.2 to

remove the excess of protein.

Blocking step: free surface sites are blocked by placing a drop of

40 μL of a 2% (w/v) solution of BSA for 15 minutes followed by a

washing step with 0.1 M Tris pH 7.2 buffer containing 1% of BSA.

Immobilization of oligonucleotide probes onto the electrode

surface: 40 μL of 3’-biotynilated oligonucleotide probes (0.5 ng/mL)

is left on the electrode surface for 15 minutes. Finally, the

electrodes are rinsed with 2 × SSC buffer pH 7.2 containing 1% of

BSA.

Hybridization is performed at room temperature placing 30 μL

of FITC-labeled oligonucleotide target solutions in 2 × SSC buffer

pH 7.2, containing 1% of BSA, on the surface of the genosensor for

45 minutes and then rinsing with 0.1 M Tris pH 7.2 buffer containing

1% of BSA. The methodology used to detect one-base mismatch

strands is similar, but in this case 25% formamide is included in the

hybridization buffer.



9.4.1.2 Analytical signal recording

Two strategies are performed to detect the hybridization event:

enzymatic detection and electrocatalytic detection. The following

steps are carried out:

Enzymatic detection: Reaction with antibody anti-FITC AP conju-

gate (Ab-AP): an aliquot of 40 μL of Ab-AP solution (1/100 dilution)

is dropped on the genosensor device for 60 minutes. Then a washing

step with 0.1 M Tris buffer pH 9.8, containing 1% BSA, is carried out.

Enzymatic reaction: An aliquot of 30 μL of 6 mM 3-IP is deposited

on the electrode surface for 20 minutes. After that, the reaction

is stopped by adding 4 μL of fuming sulphuric acid and 10 μL of

ultra-pure water. In this step, the corresponding indigo product is

converted to its parent hydrosoluble compound IC.

Analytical signal recording: The SPCEs are held at a potential of

−0.25V for 25 s, and then, a cyclic voltammogram is recorded from

0.25 to +0.20V at a scan rate of 50 mV/s. The anodic peak current is

measured in all experiments.

Electrocatalytic detection: A 50 μL portion of 0.2 M HCl solution

is dropped on the electrode surface and the electrode is held at a

potential of +1.35V for 1 minute. Then, the chronoamperometric

detection is performed at −1.40 V, recording the electric current

generated for 5 minutes.

Figure 9.10 shows the scheme of the genosensor device and

the analytical signals obtained with electrocatalytic detection (Fig.

9.10A) and enzymatic detection (Fig. 9.10B).

Moreover, the significance of the attachment of biotinylated

oligonucleotide probes through the streptavidin/biotin interaction

has been tested in a previous work [15]. When a double-labeled

(biotin and fluorescein) poly-T was attached to the electrode surface

through the streptavidin/biotin interaction, the peak currents were

much higher than those obtained when it was accumulated on

the electrode surface by physical adsorption. This fact means that

streptavidin/biotin interaction allows to attach and orient the

oligonucleotide strands on electrode surface, whereas the direct

adsorption of the oligonucleotide on the electrode surface results


Applications 307

Figure 9.10. Schematic representation of the analytical procedure fol-

lowed for the construction of the genosensor and the detection of a comple-

mentary target and a single-base mismatch target. (A) Electrocatalytic and

(B) enzymatic detection. Reproduced with permission from ACS [29, 54].

in very poor manner. Using this method of immobilization of the

oligonucleotide probes, the genosensor devices are stable for a year

if they are stored at 4◦C.

The ply (pneumolysin sequence) genosensor has been used for

detecting oligonucleotide sequences containing a one- or three-

base mismatch (plymism1 and plymism3, respectively). Three

different concentrations of complementary ply, plymism1, and

plymism3 targets were assayed and three genosensors were used

for each concentration. Figure 9.11 displays the results obtained

with both enzymatic and electrocatalytic detection. For the three

concentrations assayed, the analytical signal obtained for the three-



Figure 9.11. Ply genosensor response to the complementary target

(ply, white bars), the single-base mismatch target (plymism1, grey bars),

and the three-base mismatch target (plymism3, black bars) for different

concentrations. Data are given as average ±SD (n = 3). (a) Enzymatic and

(b) electrocatalytic detection. Reproduced with permission from ACS [29,

54].

base mismatch oligonucleotide sequence is almost the background

signal, indicating that three-base mismatch ply targets can be

perfectly discriminated from the complementary ply target. For the

one-base mismatch oligonucleotide sequence, the analytical signals

obtained only decrease about 30% with respect to those obtained

for the complementary target.

In the optimized experimental conditions the ply genosensor

has been tested for different concentrations of the complementary

oligonucleotide target. In the case of the enzymatic detection, a

linear relationship between peak current and concentrations of

complementary ply target has been obtained between 0.1 and

5 pg/μL, with a correlation coefficient of 0.9993. Thus, these

genosensors can detect 0.1 pg/μL, which is 0.49 fmol of ply target

in 30 μL.


Applications 309

In the case of the electrocatalytic detection, a linear relationship

between the recorded current and the logarithm of the concen-

tration of ply target is obtained for concentrations between 5 and

100 pg/μL. These genosensors can detect 5 pg/μL (24.5 fmol

in 30 μL) of complementary ply target, using the electrocatalytic

detection.

To improve the selectivity of the ply genosensor, more stringent

experimental conditions are tested. A concentration of 25% for-

mamide is added to the hybridization buffer. It is well known that

this molecule hampers the hybridization reaction. In these more

stringent conditions and using the enzymatic detection, a linear rela-

tionship between peak current and concentration of oligonucleotide

target is obtained for concentrations between 0.25 and 5 pg/μL.

Genosensors can detect about 1.2 fmols of complementary target in

30 μL in these more stringent experimental conditions.

In the case of electrocatalytic detection, a linear relationship

between the recorded current and the logarithm of the concen-

tration of oligonucleotide target is obtained for concentrations

between 50 and 1000 pg/μL.

Using this strategy of detection, the genosensors can detect about

245 fmol of complementary target in 30 μL in these more stringent

experimental conditions.

As expected, the sensitivity decreases in these stringent

experimental conditions for both enzymatic and electrocatalytic

detection but the detection of one-base mismatch on an oligonu-

cleotide sequence can be performed for any concentration assayed

(Fig. 9.12).

Although the sensitivity of the electrocatalytic detection is 50-

fold (under non-stringent conditions) and 200-fold (using 25%

formamide in the hybridization solution) lower than that obtained

with the enzymatic detection, the analysis time is considerably

shorter, because the analytical signal is achieved directly from

the platinum complex whereas in the enzymatic detection two

additional steps are necessary to obtain the analytical signal:

the reaction with antibody anti-fluorescein and the enzymatic

reaction. Thus, the overall analysis time of this chronoamperometric

method is about the half than that resulting from the enzymatic

method.



Figure 9.12. Ply genosensor responses for different concentrations of

complementary target (ply, white bars) and the single-base mismatch

target (plymisms1, grey bars) when 25% formamide is included in the

hybridization buffer. Data are given as average ±SD (n = 3). (a) Enzymatic

and (b) electrocatalytic detection. Reproduced with permission from ACS

[29, 54].

9.4.2 Alkaline Phosphatase-Catalyzed Silver Deposition forElectrochemical Detection

In this section a new substrate solution is described that combines

an indoxyl compound, 3-indoxyl phosphate (3-IP), and silver ions.

The resulting enzymatic product of 3-IP is indigo blue, an aromatic

heterocycle insoluble in aqueous solutions. Two strategies can be

carried out to detect the product: its sulfonation in acidic medium,

giving rise to indigo carmine (IC), or its solubilization in basic

medium and in the presence of dithionite salt, giving rise to

leucoindigo. The main drawback of these methodologies is that, in all

cases, it is necessary to add a step for detection after the enzymatic

reaction and the use of aggressive agents such as concentrated


Applications 311

sulfuric acid or sodium dithionite, respectively. The substrate

proposed here overcomes these drawbacks and, moreover, improves

the sensitivity of the methodology. To demonstrate the better

sensitivity obtained with this substrate, an enzymatic genosensor on

SPCEs for the identification of nucleic acid determinants exclusively

present on the genome of the pathogen Streptococcus pneumoniaehas been developed. The different steps of this genosensor have

been optimized in a previous work [54]. Orientation of the strands

in the sensing phase is achieved by modifying the surface of the

electrode with streptavidin by physical adsorption followed by

the immobilization of biotinylated oligo probe. The biotinylated

oligonucleic acid probe used in this work targets the autolysin

(lytA) gene. This target is randomly labeled with the Universal

Linkage System (ULS). This system binds to DNA at the N7 position

of guanine, resulting in the attachment of a label molecule to

the DNA. The label molecule used in this study was fluorescein

(FITC). Electrochemical detection is achieved with an anti-FITC

alkaline phosphatase-labeled antibody (Ab-AP) and using substrate

proposed here, 3-IP/Ag+.


The electrode pretreatment was carried out by applying an anodic

current of +5 μA for 2 minutes in a 40 μL aliquot of 0.1 M

H2SO4 Then, the electrodes were washed using 0.1 M Tris-HNO3

buffer pH 7.2. The adsorption of streptavidin onto the electrode

surface was performed leaving an aliquot of 10 μL of a 1× 10−5 M

streptavidin solution on the electrode surface between overnight at

4◦C. Then, the electrode was washed with 0.1 M Tris-HNO3 buffer pH

7.2 to remove the excess of protein.

Free surface sites were blocked placing a drop of 40 μL of a

2% (w/v) solution of BSA for 15 minutes followed by a washing

step with 0.1 M Tris-HNO3 pH 7.2 buffer containing 1% BSA.

Immobilization of the probe was performed dropping 40 μL of 3’-

biotinylated oligonucleotide probe (0.5 ng/μL) for 15 minutes. Then,

the electrodes were rinsed with 2 × SSC buffer pH 7.2 containing

1% BSA. After that, the hybridization was performed at room

temperature placing 30 μL of FITC-labeled oligonucleotide target



solutions in 2× SSC buffer pH 7.2, containing 1% BSA, on the surface

of the genosensor for 45 minutes and then rinsing with 0.1 M Tris-

HNO3 pH 7.2 buffer containing 1% BSA. Then, a reaction with Ab-

AP was performed dropping aliquots of 40 μL of Ab-AP solutions

(1/100 dilution) on the genosensor device for 60 minutes. After a

washing step with 0.1 M Tris-HNO3 buffer pH 9.8, containing 1%

BSA, the enzymatic reaction was carried out by dropping an aliquot

of 35 μL of a mixture of 5.6 mM 3-IP and 0.4 mM silver nitrate

solutions for 20 minutes, protected from light. Then, the SPCE was

held at −0.20 V for 5 s, and a cyclic voltammogram was recorded

(in the same enzymatic reaction medium) from −0.20 to 0.50 V at a

scan rate of 50 mV/s to obtain the analytical signal.

9.4.2.2 Results

Once the procedure was optimized, an enzymatic genosensor for the

identification of a nucleic acid determinant exclusively present on

the genome of the pathogen S. pneumoniae was developed. This DNA

sensor has been described and optimized by our research group in

the previous section. In this work, for the electrochemical detection

step, 3-IP was used as substrate and then sulfuric acid was added to

generate an electroactive compound termed indigo carmine, which

is quantified by cyclic voltammetry.

In this case, by combining the 3-IP with silver ions, the metallic

silver deposited on the electrode surface is detected directly without

the need of any more steps to obtain the analytical signal. Thus, the

use of sulfuric acid is avoided.

Using the optimized experimental conditions, the response of the

genosensor formed with 3’-biotinylated autolysin gene lytA probe

for different concentrations of the complementary oligonucleotide

target has been evaluated. Figure 9.13 shows the calibration

plot (Fig. 9.13A) and the voltammograms corresponding to each

concentration as well as the voltammogram corresponding to the

noncomplementary target for the highest concentration assayed

(Fig. 9.13B).

A linear relationship between peak current and concentration of

complementary lytA target is obtained between 7 and 700 fg/μL,

with a correlation coefficient of 0.9995. The reproducibility of the


Applications 313

Figure 9.13. (A) lytA genosensor responses for different concentrations

of complementary target. Data are given as average ±SD (n = 3). (B)

Cylic voltammograms corresponding to the background (700 fg/μL of

noncomplementary target) and to each concentration of complementary

target of the linear calibration curve. Reproduced with permission from ACS

[56].

analytical signal for the concentrations of complementary target

assayed is shown with error bars. It is composed between 4 and 10

in terms of percent RSD.

Also, comparing linear ranges obtained for target autolysin

through both methodologies, the sensitivity of the assay is improved

by at least 1 order of magnitude.

Thus, this genosensor can detect 7 fg/μL, approximately 14-

fold less than the concentration detected when the enzymatic

reaction was carried out only with 3-IP [54]. Also, the use of

3-IP as the enzymatic substrate allows a better control of the



silver deposition versus the use of another substrate such as p-

aminophenyl phosphate that is more unstable and produces higher

background signals.

Moreover, the hybridization reaction with noncomplementary

target does not occur for all concentrations assayed (see the

voltammogram in Fig. 9.13B for the highest concentration of

noncomplementary target assayed, 700 fg/μL). This fact shows that

non-specific adsorptions are not observed. Regarding the selectivity

of the genosensor, this system has been studied in the previous

section and this is able to discriminate one-base mismatched

strands.

9.4.3 Genosensor for SARS Virus Detection Based on GoldNanostructured Screen-Printed Carbon Electrode

In this section, a DNA hybridization assay with enzymatic electro-

chemical detection was carried out on a disposable gold nanos-

tructured screen-printed carbon electrode (SPCnAuE), which allows

working with small volumes. Gold nanoparticles (NPs) which are

formed in situ by applying a constant current intensity during a fixed

time act as an immobilization and transduction surface. Although

thick gold substrates are reported in the literature for enzymatic

DNA detection (screen-printed gold electrodes [46], 2 mm thick film

gold electrodes [66], or gold disk electrodes [69]), gold NPs have

been unusually used as electrochemical transducers, despite of their

widespread use as DNA labels due to the electrochemical properties

of gold NPs [70].

The sequence chosen as target is included in the 29 751-base

genome of the SARS (severe acute respiratory syndrome)-associated

coronavirus. A 30-mer oligonucleotide with bases comprised

between numbers 29 218 and 29 247, both included, was chosen.

This is the causative agent of an outbreak of atypical pneumonia,

first identified in Guangdong Province, China, that has spread to

several countries. The sequence corresponds to a gene that encodes

the nucleocapsid protein (422 amino acids), specifically a short

lysine-rich region that appears to be unique to SARS and suggestive

of a nuclear localization signal.


Applications 315

9.4.3.1 Gold nanostructuration of screen-printed carbonelectrodes

Gold nanostructures were in situ generated over SPCEs (SPCnAuEs)

applying a constant current intensity of −5 μA for 2 minutes in a

0.1 mM acidic solution of AuCl−4 . After that, and in the same medium,

a potential of +0.1 V was applied during 2 minutes, in order to

desorb hydrogen.


The formation of the sensing phase was performed by dropping

20 μL of 3’-thiolated oligonucleotide probe 10 nM for 20 minutes

and after rinsing with 0.1 M Tris-HNO3 pH 7.2, a blocking step

with casein (2%) was carried out. Then, the electrodes were rinsed

with 2 × SSC buffer pH 7.2 containing 1% BSA. After that, the

hybridization was performed at room temperature placing 40 μL

of 3’-biotinylated oligonucleotide target solutions in 2 × SSC buffer

pH 7.2, containing 1% BSA, on the surface of the genosensor for 1

hour and then rinsing with 0.1 M Tris-HNO3 pH 7.2 buffer containing

2 mM Mg(NO3)2. Then, a reaction with alkaline phosphatase labeled

streptavidin (S-AP) was performed dropping aliquots of 40 μL of

S-AP solutions (5 × 10−10 M) on the genosensor device for 60

minutes. Finally, after a washing step with 0.1 M Tris-HNO3 buffer

pH 9.8, containing 20 mM Mg(NO3)2, the enzymatic reaction of the

substrate, a mixture of 3-indoxyl phosphate (3-IP) and silver nitrate,

was performed. In this reaction, 3-IP produces a compound able to

reduce silver ions in solution into a metallic deposit. The deposited

silver is electrochemically stripped into solution and measured by

anodic stripping voltammetry giving place to the analytical signal

Fig. 9.14.

9.4.3.3 Results

Adsorption of thiolated probes was studied, in this sense adsorption

time and probe concentration were tested. Results obtained shown

that 20 minutes was enough time to reach a plateau in the analytical

signal, and probe concentration was fixed in 10 nM, because

higher concentrations resulted in a decrease in the analytical signal.



Figure 9.14. Schematic representation of genosensor design. Reproduced

with permission from Wiley InterScience [11].

This decrease could be because the amount of probe strand on

the electrode surface is too high so hampering the hybridization

reaction by stearic impediments and/or because high amounts of

strands on the electrode surface blocks the electrodic surface.

The effect of the thiol group was tested using non labeled probe

strands following a similar procedure. In this case signals obtained

were significantly lower than those obtained with the thiol group,

and due to the unspecific adsorption of the probes.

S-AP concentration was also tested, comparing the analytical

signal obtained with that obtained due to unspecific adsorption of S-

AP. In this case, 5 × 10−10 M of S-AP was the maximum concentration

where the unspecific adsorption was not observed.

Once the parameters that affect the procedure had been studied,

a calibration curve for the biotinylated target strand was performed.

The peak current was linear with the concentration of the target

strand in the range comprised between 5 and 100 pM. The detection

limit, calculated as the concentration corresponding to a signal that

is three times the standard deviation of the intercept, was found to

be 4.6 pM.

In addition, and in order to test the stability of the genosensor,

and minimize the analysis time, the sensing phase was formed and


Applications 317

Figure 9.15. Analytical signals obtained for 50 pM of biotinylated target,

using a genosensor fresh prepared (A), a genosensor in which thiolated

probes were immobilized and stored at 4◦C overnight (B), and a genosensor

stored at 4◦C where both, probe and blocking agent were immobilized.

Reproduced with permission from Wiley InterScience [11].

stored at 4◦C. With this aim, SPCnAuEs were modified following the

procedure previously described and stored at 4◦C overnight.

Results obtained are displayed in Fig. 9.15, and show that the

analytical signal due to a biotinylated target concentration of 50 pM

results incremented in about 15%. However, when the blocking

step is also carried out prior to the storage of the sensing phase,

the analytical signal due to the same concentration of biotinylated

target gives rise to a decrease of about 20% of the analytical

signal.

With SPCnAuEs modified with the probe strand and stored at

4◦C overnight, a calibration plot was recorded. A linear relationship

of the analytical signal with the concentration of the biotinylated

target strand in the range comprised between 2.5 and 50 pM

was obtained. The detection limit, calculated as the concentration

corresponding to a signal that is three times the standard deviation

of the intercept, was found to be 2.5 pM. The linear range obtained

with this methodology is closer than that obtained when the sensing

phase is freshly prepared, but its sensibility is around three times

that obtained with the former methodology. Moreover, storage of the

sensing phase permits to minimize the analysis time and increases

the possibility of storing the genosensors and using them when

necessary.



Figure 9.16. Analytical signals obtained for 50 pM of biotinylated target

with 25% formamide, using a genosensor in which thiolated probes were

immobilized and stored at 4◦C overnight. Reproduced with permission from

Wiley InterScience [11].

In order to study the selectivity of the genosensor developed,

hybridization was carried out with 1-, 2-, and 3-base mismatch

complementary strands. When hybridization was performed with-

out applying stringency conditions using a target strand con-

centration of 50 pM, there was no discrimination between the

analytical signals. However, when 25% formamide is added to

a biotinylated target concentration of 50 pM in order to apply

stringency conditions, it is possible to discriminate between the

complementary strand and the 1-, 2-, or 3-base mismatch strands,

as can be seen in Fig. 9.16.

9.4.4 Simultaneous Detection of Streptococcus andMycoplasma Pneumoniae Using Gold-ModifiedSPCEs

In this section, a genosensor for the simultaneous detection of two of

the principal causative bacteria of community acquired pneumonia

is developed using a dual screen-printed sensor. The genosensor

design is the same that the used in the previous section (see

Fig. 9.14).


Applications 319

Figure 9.17. Commercial dual screen-printed carbon electrode. See also

Color Insert.

The target sequences have been chosen so that the same primer

is able to generate the PCR products of both bacteria. Thus, with

a unique screen-printed strip it is possible to identify the causing

bacteria of the disease.

The dual screen-printed electrode used in this section is shown

in Fig. 9.17.


As it has been commented, the genosensor design is the same that

was used in section 9.4.3.2, with some variations

Gold nanostructuration is carried out by applying a constant

current of −5 μA for 2 minutes in an acidic medium containing

AuCl4− 1 mM. Then the electrode is generously rinsed with water.

A 4 μL aliquot of 50 nM thiolated probes is dropped in each

working electrode for 10 minutes. One working electrode supports

the probe corresponding to S. pneumoniae, and the other supports

the probe corresponding to M. pneumoniae. Then, the electrode is

rinsed with 0.1 M Tris-HNO3 buffer pH 7.2, and a blocking step is

carried out with a 40-μL aliquot of casein 2% for 20 minutes and

rinsed with 0.1 M Tris buffer pH 7.2

Hybridization step is carried out at room temperature in 2 × SSC

buffer by dropping a 40 μL aliquot of the biotinylated target for 1

hour and rinsing with Tris buffer pH 7.2. After that 40 μL of 5× 10−10

M S-AP are dropped on the electrode for 1 hour. Then the electrode

is rinsed with Tris buffer pH 9.8 and enzymatic reaction with 3-IP

and silver ions, and detection step is carried out as mentioned in

previous sections.



9.4.4.2 Results

After verifying that the presence of a certain quantity of not

complementary target strand does not concern the analytical signal

obtained by the complementary strand a simultaneous calibration

plot for both target sequences is carried out Fig. 9.18.

The S. pneumoniae target strand show a linear relationship of the

analytical signal with the concentration of the biotinylated target

strand in the range comprised between 50 pM and 1 nM. The

detection limit, calculated as the concentration corresponding to a

signal that is 3 times the standard deviation of the intercept, was

found to be 34 pM.

The M. pneumoniae target strand show a linear relationship

of the analytical signal with the concentration of the biotinylated

target strand in the range comprised between 10 pM and 1 nM. The

detection limit, calculated as the concentration corresponding to a

signal that is three times the standard deviation of the intercept, was

found to be 5 pM.

It has been seen that the presence of another bacteria in the

sample does not concern significantly the analytical signal obtained

for an individual bacteria (though the analytical signal diminishes a

bit), this indicates that simultaneous calibrations or identifications

of several bacteria can be done.

Later, identification of PCR products of these bacteria was carried

out. Dilution of the PCR product has been studied, and a 1:4

Figure 9.18. Simultaneous calibration plots for S. pneumoniae and M.pneumoniae obtained with gold nanostructured dual screen-printed carbon

electrodes.


References 321

dilution was determined as optimum for the bacteria identification.

Identification of PCR products has been realized successfully in 90%

of cases.

9.5 Conclusion

As it has been shown in previous sections, the use of screen-printed

electrodes as support for genosensor devices offers enormous

opportunities for their application in molecular diagnosis. The

technologies used in the fabrication of these electrodes allow the

mass production of reproducible, inexpensive and mechanically

robust strip solid electrodes. Other important advantages of these

electrodes are the possibility of miniaturization as well as their easy

manipulation in a disposable manner and therefore the use of small

volumes, diminishing the cost of the analysis. This is an important

issue that makes this methodology for the detection of DNA more

attractive.

Moreover, in addition, the versatility of design of screen-printed

electrodes allows to carry out a simultaneous detection of several

DNA sequences in the same analysis.

Very sensitive methods are always required for DNA sensing.

Although enough sensitivity to avoid PCR amplification has been

achieved by use of enzymatic labels or metal tags, most of the assays

routinely start with a PCR or other biochemical amplification. More-

over, although label-free formats are used, most of the strategies

followed to obtain the analytical signal involve several washing steps

and need the use of labeled reagents (or labeling procedures) or

indicators, which complicates the assay performance.

Sensitive methodologies can also be obtained through the

electrodic modification with a nanostructured material, taking

advantage of the special characteristics that nanostructuration

offers.

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Chapter 10

Synthetic Polymers for ElectrochemicalDNA Biosensors

Adriana Ferancovaa and Katarına Benıkovab

aProcess Chemistry Centre, Laboratory of Analytical Chemistry, AboAkademi University,FI-20500 Turku-Abo, FinlandbInstitute of Analytical Chemistry, Slovak University of Technology in Bratislava,81237 Bratislava, [email protected]; [email protected]; [email protected]

10.1 Introduction

In recent years, electrochemical DNA biosensors have been widely

used for many purposes, such as study of DNA hybridization as

well as investigation of interactions of DNA with other molecules,

including DNA association with low-molecular-weight compounds

or detection of damage to DNA. To make DNA biosensors powerful,

there is an increased interest in the use of different materials

which can be applied as the DNA–transducer interface. Among

them, conducting as well as nonconducting polymers have become

more and more popular. They offer an environment suitable for

direct simple adsorption of the DNA onto the polymeric matrix or

incorporation of the DNA into the polymeric network. Polymers

can also be mixed with the nanomaterials to form nanocomposites

providing many new interesting properties, including rapid electron



330 Synthetic Polymers for Electrochemical DNA Biosensors

transfer, enhanced DNA immobilization, and better stability and

sensitivity of resulting DNA biosensors.

The aim of this review is to describe the possibilities of modern

utilization of conducting as well as nonconducting polymers in the

preparation and application of electrochemical DNA biosensors and

to report their advantages and disadvantages. This chapter deals

mostly with the state of the art in the last few years.

10.2 Modification of Electrode Surface with Polymers

Polymeric films can be prepared at the surface of metal, glassy

carbon, as well as carbon paste electrodes. The preparation of

conducting polymers at the surface of carbon electrodes employed

in biosensors is already reviewed [1]. The methods mostly used are

solvent casting, spin coating, and electropolymerization.

10.2.1 Solvent Casting

In solvent casting method an already prepared polymer is first

dissolved in the appropriate solvent and then simply cast onto the

surface of the electrode. After solvent evaporation, the film of poly-

mer is formed. It is a very simple approach, but unfortunately two

disadvantages have to be considered, uniformity of the polymeric

film and reproducibility of its preparation [2]. This method is usually

used for the preparation of redox active or nonconducting polymers

[3]. Coatings of composites of nanomaterials with polymers are also

often prepared by this method [4].

10.2.2 Spin Coating

Problem with uniformity and reproducibility can be avoided using

the spin coating method. In this case, dissolved polymer is put onto

the electrode surface, which is then rotated at high speed. The

centrifugal force causes the spread of the solution, leading to a

more uniform coating than in the case of solvent casting. During

the rotation, the solvent is evaporated. Problem was reported with

control of the structure and thickness of polymer coatings [5].


Modification of Electrode Surface with Polymers 331

However, this method was successfully used for the preparation of

the film of poly(3,4-ethylenedioxythiophene) (PEDOT) doped with

poly(styrene sulfonic acid) (PSS) at the surface of ITO electrodes

[6] and for the preparation of immunosensors based on conjugated

poly(phenylene vinylene) derivatives of defined thickness [7].

10.2.3 Electropolymerization

Another method often used for the preparation of conducting

polymers, such as polypyrrole (Ppy), polyaniline (PANI), polythio-

phene, and their derivatives is deposition by electropolymerization

in the electrolyte-containing monomers. This method can be used

for the polymerization of compounds which possess a relatively

low anodic oxidation potential and are susceptible to electrophilic

substitution reaction. The electropolymerization is reported as a

simple as well as reproducible method, where the monomer is

first oxidized to a cation radical. Next, the molecule of monomer

is attached to form a dication. Repeated process lengthens out the

polymeric chain and the final polymer is formed. The advantage of

this method is that the rate of film deposition can be controlled by

varying the potential of the working electrode in the system. It is

a simple and reproducible method [8]. Electropolymerization can

be provided potentiostatically, galvanostatically, or by the potential

cycling method. In general, the potentiostatic method is used to

prepare thin films, while the galvanostatic method enables to

prepare thick films [9].

The properties of the polymeric films can be easily modified

by functionalization of the polymer. Two methods are reported for

these purposes:

(i) The functional groups are attached to the monomers through

covalent bonds and then electropolymerization is provided

[10]. The disadvantages of this method are loss of polymer

conductivity, steric hindrance, and cross-linking effects.

(ii) Another often used method is incorporation of a dopant into

the polymeric network electrostatically during the process of

electropolymerization [11].



Electropolymerized polymers are molecular composites containing

cationic polymer backbone counter anions for maintenance of

charge neutrality [12]. Anions from the electrolyte solution or other

negatively charged molecules present in the electrolyte solution

during electropolymerization can be employed as dopants. For

example, the polypyrrole/ferrocyanide-film-modified carbon paste

electrode was prepared by potentiostatic electropolymerization of

pyrrole in the presence of ferrocyanide ions [13]. Incorporated

ferrocyanide worked as a mediator of ascorbic acid oxidation.

10.3 Polymer-Assisted DNA Immobilization

Polymer-assisted immobilization of biomolecules, including DNA, is

widely reviewed [14–18]. DNA can be either immobilized at the sur-

face of polymer-modified electrode or incorporated in the polymer

layer. In the second case, the method of electropolymerization is

mostly used.

10.3.1 Immobilization of DNA onto Polymer-ModifiedElectrode Surface

DNA can be attached to the polymer-modified electrode surface

using several methods: simple adsorption, covalent bonds (first

appropriate functional groups are introduced to the polymer, then

DNA is covalently attached), or affinity binding (avidin–biotin).

Adsorption is the simplest method of DNA immobilization, and

it can be achieved by different ways. A polymer-modified electrode

can be simply dipped into the solution containing DNA [19] or a drop

of DNA solution is cast onto the polymer-modified electrode surface

and let to evaporate to dry [20]. It is also convenient to use negative

charge of DNA for its adsorption onto positively charged polymer

via electrostatic forces [21]. Electrostatic adsorption of DNA onto

conducting Ppy is well studied [22]. It was found that this process

is significantly pH dependent and is higher in acidic media as well as

at high ionic strength. Dielectric studies showed that DNA formed

an insulating layer at the surface, which significantly diminished

the ionic conductivity character and maintained the mobility of the


Polymer-Assisted DNA Immobilization 333

doping anions within the bulk Ppy [23]. Electrostatic adsorption

of calf thymus DNA onto the polypyrrole–polyvinyl sulphonate

(Ppy–PVS) film was also studied using cyclic voltammetry as well

as spectroscopic methods [24]. Maximum adsorption of DNA was

observed at the pH of 6.0. Time-dependent kinetics found in DNA

adsorption was explained by a gradual interchange of PVS with DNA.

Immobilization of the DNA onto polymer modified surface can

be realized by electrodeposition, which is a well-known method

[25]. Application of positive potential in this process can enhance

the DNA immobilization as well as the stability of immobilized

DNA. Diaz-Gonzalez et al. [26] studied the DNA immobilization onto

a polylysine-modified electrode at different potentials. The best

results were obtained using a potential of +0.5 V for 120 seconds.

DNA was also electrodeposited onto a poly( p-aminobenzensulfonic

acid)-modified glassy carbon electrode (GCE) at +1.5 V for 30

minutes [27] or onto overoxidized Ppy-modified electrode at +1.8 V

for 30 minutes [28].

Covalent immobilization of DNA onto polymer-modified surface

is also widely used. The advantage of this method is enhanced

stability and the possibility to control the orientation of DNA for

better accessibility to the substrate and to facilitate macromolecular

interactions [14]. This method needs functionalization of the

DNA or polymeric film, or both of them, with functional groups

appropriate for covalent linking. For these purposes, 1-ethyl-

3-(3-dimethylaminopropyl)carbodiimide (EDC) is often used for

electrode surface activation. The DNA immobilization is realized

by dipping a polymer-modified electrode into a solution containing

DNA or oligodeoxyribonucleotides (ODNs) and EDC [29–31]. EDC

can also be used in combination with N -hydroxysuccinimide

(NHS) [32–35]. Another possibility is covalent binding of DNA

onto functionalized polymeric film. For example, amino-labeled

ODN was grafted on the Ppy copolymer by a direct binding to

the activated ester groups [36] or on pyrrole–2-carboxyaldehyde-

Ppy/PVS, leaving —CHO groups [37].

Indirect immobilization of DNA using intermediate system

avidin–biotin is reported as a form of affinity binding [38]. Avidin

with an activated —COOH group was attached onto PANI film

electropolymerized at the surface of a Pt electrode, and then a



biotin-modified DNA probe was immobilized in order to prepare a

DNA hybridization biosensor [39]. Direct DNA immobilization via

EDC–NHS coupling was compared to indirect affinity immobilization

onto the Ppy–PVS modified Pt electrode [40]. It was found that

covalent DNA immobilization showed faster redox processes and led

to enhanced sensitivity, which was ascribed to increased interaction

of ODNs stationed near the Ppy–PVS surface.

10.3.2 Immobilization of DNA Within a Polymeric Matrixby Electropolymerization

Another widely used method of the DNA immobilization is incorpo-

ration of DNA into the polymer matrix during electropolymerization.

As it was described previously, negatively charged biomolecules,

such as DNA and oligonucleotides, can be advantageously employed

as dopants of a positively charged polymeric structure. The control

of the current density in the galvanostatic method or potential in

the potentiostatic method during the electropolymerization process

is very important to avoid loss of bioactivity or decomposition

of entrapped biomolecules. This method is widely used in the

case of conducting polymers, such as Ppy and PANI. Biomolecule

immobilization is realized in the solution containing monomer and

biomolecules. In this case DNA acts as solo dopant [41, 42]. In this

process, the supporting electrolyte (NaCl, LiClO4) can be used to

permit the growth of the film with low concentration of sample ODN

[43].

10.4 Application of Synthetic Polymers in DNABiosensors

10.4.1 Electronically (Intrinsically) Conducting Polymers

Conducting polymers (CPs) are very popular matrices suitable

for biomolecule immobilization in biosensors [44]. They show a

suitable flexibility and can be chemically modified as required. The

advantage of CPs is that their electrochemical synthesis allows

direct deposition of a polymer on the electrode surface while


Application of Synthetic Polymers in DNA Biosensors 335

(a) (b) (c)

Figure 10.1. The mostly used conducting polymers: polypyrrole (a),

polyaniline (b), and polythiophene (c).

simultaneously trapping the biomolecules [41]. It is also possible

to control the polymeric film thickness, the spatial distribution of

the immobilized biomolecule, and modulation of its activity [45].

They are mostly organic conjugated polymers with a conjugated

π -electron system. In general, conducting polymers are considered

those with the conductivity higher than 103 S cm−1, materials with

conductivity in the range from 103 to 10−8 S cm−1 are semiconduc-

tors, and materials with conductivity lower than 10−8 S cm−1 are

considered as insulators [46]. The conducting polymers mostly used

in DNA biosensors are polypyrroles, polyanilines, and polytiophenes

(Fig. 10.1).

10.4.1.1 Polypyrroles

Polypyrroles and their derivatives are one of the most extensively

used polymers for the preparation of biosensors. This group of

polymers has excellent properties which can be advantageously

used in enzyme (transducing the analytical signal generated by

redox enzyme reactions) as well as affinity biosensors (DNA biosen-

sors, immunosensors) [47]. Polypyrrole (Ppy) can be prepared by

chemical or electrochemical polymerization. For the preparation of

DNA biosensors, usually method of electropolymerization is used.

Cyclic voltammetry or deposition at constant potential is often

used for these purposes. Ramanavicius et al. [48] reported the

potential pulse technique as the most suitable method for the

preparation of nanostructured Ppy with entrapped biomolecules.

Ppy films prepared by cyclic voltammetry and normal pulse



voltammetry (NPV) in order to prepare electrochemical DNA biosen-

sors were compared [49]. The NPV method enabled to prepare Ppy

nanofiber films with higher electroactivity due to higher specific

surface area. The potentiostatic and potentiodynamic method of

electropolymerization was used to prepare Ppy nanofibers [50].

Electrodes prepared by the potentiostatic procedure showed higher

responses to the oxidation of dsDNA than the electrodes prepared

by potentiodynamic methods.

Ppy can be electropolymerized from both aqueous and nonaque-

ous solvents [12]. For DNA biosensors the biocompatability of Ppy is

important as well as the fact that it can also be electropolymerized

from neutral aqueous solutions. Different conditions affecting DNA

adsorption onto conducting Ppy, including pH, buffer nature, ionic

strength, and substrate, were studied [22]. Maximum amount of

DNA was adsorbed from a solution of pH 5.1 because of the

high density of positive charge of Ppy, and also positive effect

of ionic strength was reported. DNA adsorbed at the Ppy surface

decreases the ionic conductivity of the polymer, but on the other

hand maintains the mobility of the dopant anions within the bulk

Ppy [23]. Anions incorporated as dopants into the Ppy during the

process of electropolymerization have a positive effect on polymer

stability [12]. Anions from the supporting electrolyte incorporated

into the polymer achieve its electroneutrality. However, other

anions can also be used as counterions. Large polymeric anions,

such as polyvinyl sulfonate, were used as counterions in the

preparation of DNA biosensors [40, 51]. Such doped Ppy can

displace negative PVS with PO−4 of DNA [24]. It was found

that the adsorption of DNA onto electropolymerized Ppy–PVS

reached the maximum at pH 6.0, and FTIR studies showed the

electrostatic interaction between the DNA and polymeric film. A

Ppy–PVS film was prepared at the surface of ITO electrodes by

chronopotentiometrical electropolymerization from the solution

containing pyrrole and PVS [51]. DNA was then physisorbed onto

the polymer, and the resulting biosensor had improved sensitivity

to 3-chlorophenol (0.1–25 ppm) and 2-aminoanthracene (0.01–15

ppm). The response time was about 30 seconds. Incorporation

of the DNA into the polymeric layer during electropolymerization

led to increased sensitivity to both 3-chlorophenol (0.01–55 ppm)



and 2-aminoanthracene (0.001–6 ppm) [52]. A similar DNA biosen-

sor was also used for the detection of organophosphates such

as chlorpyrifos and malathion up to 0.0016 ppm and 0.17 ppm,

respectively [53].

Because of the negative charge, ODN can also serve as a dopant of

Ppy. A Ppy–ODN film was prepared at the surface of gold electrode

for genoelectronic application [54]. It was found that the redox

activity of the biosensor was affected by presence of ODN molecules

and it was able to discriminate between synthetic oligonucleotides

and chromosomal DNA. Komarova et al. [55] prepared the DNA

biosensor at the surface of ITO electrodes electrochemically from the

solution containing Py and ODN. ODN served as a sole dopant, and

a prepared biosensor was used for chronoamperometric detection

of the target ODN with the detection limit of 1.6 fmol in 0.1

ml. An ssDNA/polypyrrole-modified electrode for the detection

of specific bovine leukemia virus provirus DNA sequences was

prepared [56]. In this case, Ppy was electrochemically doped with

ssDNA in the presence of KCl, which eliminated a nonspecific

contribution. A Ppy film doped with oligonucleotide probe was

also formed at the surface of microelectrodes in the presence

of LiClO4 in order to prepare an impedance DNA hybridization

biosensor [57]. The biosensor was applied for the detection of

nanomolar concentrations of target ODN at the silicon array chip

containing four gold microelectrodes. The Ppy–ODN film was

also electropolymerized at the surface of Au electrode from the

solution containing pyrrole, ODN, and NaCl by continuous cyclic

voltammetry [43, 58]. An electrode was used for the detection of

DNA hybridization. A thin film of Ppy doped with an ODN probe was

electropolymerized at the surface of gold microelectrodes integrated

on the chip and used for sensing electrical potential-assisted DNA

hybridization and pathogen target DNA detection [59]. Detection of

0.34 pmol and 0.072 fmol of complementary ODN target in 0.1 mL

within a time of seconds were achieved on unpolished and polished

electrodes, respectively.

Another approach was described by Livache et al. [60]. Pyrrole

was first functionalized by ODN using pyrrole–phosphoramidite

building blocks. Next, Ppy copolymer was prepared by electropoly-

merization in the solution containing pyrrole and pyrrole–ODN. This



procedure led to the synthesis of a Ppy film bearing covalently

linked ODN. The hybridization event was detected using a quartz

crystal microbalance method (QCM) [60–63]. This method was also

applied to prepare silicon DNA chip containing 48 or 128 gold

microelectrodes, where the hybridization reaction was evaluated

using fluorescence [60, 62, 64, 65]. A similar procedure was

described for the preparation of biotinylated Ppy film at the

surface of gold quartz crystals as well as silicon chip containing

48 gold microelectrodes [66–68], where biotin was used for the

immobilization of avidin. Then biotinylated ODN was immobilized

via the biotin/avidin affinity bond.

Polypyrrole can also be functionalized with the electrochemical

indicator of DNA, such as ferrocenyl groups bearing an active ester

group used for the covalent binding of amino-labeled ODN probe

[69]. Hybridization with complementary ODN caused a decrease in

the current density and a shift of the oxidation wave of the ferrocenyl

group because of the decrease of polymer permeability. This was

explained by the change of the conformation along the conjugated

backbone of the polymer. The prepared DNA biosensor was able to

detect less than 1 pmol of target ODN.

The gold electrode was modified with a copolymer using

the monomers 3-acetic acid pyrrole and 3-N -hydroxyphthalimide

pyrrole [36, 70]. This copolymer contained activated ester groups

used for covalent grafting of an ssDNA probe bearing a terminal

amino group. It was found that porous Ppy led to a higher density

of immobilized DNA probes and improved the detection of the

hybridization reaction. The same copolymer was used for the

preparation of a multiplot DNA biosensor based on microelec-

trodes deposited on the chip [71]. The polymer offered direct

transduction of the recognition process into an electrochemical

signal because its signature varied according to a hybridization

event. Another copolymer-based DNA biosensor was prepared

by electropolymerization of Py in the presence of 4-(-3-pyrrolyl)

butanoic acid [29]. Ppy was also used as an electrostatic adsorption

matrix, which allowed immobilization of DNA onto the porous

silicon substrate without using covalent bonds [72]. Polypyrroles

are reported as a convenient matrix for the immobilization of

nanomaterials at the surface of an electrode [73]. In this case the



combination of the unique properties of conducting polymers and

those of nanomaterials exhibits a synergic effect, which positively

affects the stability, electron transfer, and performance of the

final biosensors. The Ppy film possessed the uniform surface

for the immobilization of Au–Pt hybrid nanoparticles [74]. Ppy

was also prepared by electropolymerization in the presence of

multiwalled carbon nanotubes (MWCNTs) ended with carboxylic

groups [75]. MWCNTs served as the nano-sized backbone for Ppy

polymerization, which allowed the formation of porous Ppy film

covered around the MWCNTs in a cylindrical structure and offered

stable surface for DNA immobilization. The activity of MWCNT

surface can lower the nucleation energy required for the beginning

of electropolymerization of the Ppy/DNA film [41]. Therefore, the

growth of the polymer film occurred at potential +0.4 V vs. Ag/AgCl

in contrast to +0.6 V observed at bare GCE. The high surface area

of MWCNTs also allowed the deposition of greater volume of the

polymer without increasing the thickness of the film.

10.4.1.2 Polyaniline

Polyaniline is widely used for the preparation of the electrochemical

enzyme biosensors and immunosensors [76]. However, several

applications in DNA biosensors can also be found. PANI can be pre-

pared by electropolymerization using the galvanostatic method or

the potentiostatic method, leading to a polymer adhered weakly at

the electrode surface or potential cycling, which produces polymer

well adhered at the electrode surface [77]. The electropolymerizaton

of PANI is usually provided from acidic media [78]. The properties

of the PANI synthesized from different acids were investigated

[79]. The authors showed that polymer synthesized with perchloric

acid had the highest conductivity in neutral solutions (pH of 6.6),

which is environmentally convenient for biomolecules. Abdullin

et al. [80] studied the redox properties of the DNA–polyaniline film

over a wide range of pH. Authors found that the well-reproducible

and reversible voltammetric signals of the DNA–PANI film were

observed at physiological pH values. Screen-printed carbon elec-

trodes modified with electropolymerized PANI, electropolymerized

polydiaminobenzene (PDAB), and polyethyleneimine (PEI) were



compared in order to prepare ssDNA biosensors for detection of the

hybridization event [81]. The best results were obtained using PANI

and PEI-modified DNA biosensors. PDAB-modified DNA biosensors

showed some unselective binding. Moreover, PANI allowed finer

control and monitoring of the deposition process. Similarly to Ppy,

PANI can also be doped by anionic dopants, which improve the

conductivity and stability of the resulting polymer. PANI fibers were

used as electrodes to study the influence of electrolyte counterions

and pH on the electrochemical behavior of PANI fibers [82]. The

highest currents were observed in a solution of HCl and HNO3, and

the authors concluded that the size of counterions is less important

than the anion charge. Moreover, only fully protonated PANI fibers

showed the same electrochemical properties as the PANI film. DNA

was covalently attached onto PANI nanotubes synthesized on the

graphite electrode [83]. The collective effect of PANI nanotubes as

well as enhanced conductivity led to an extremely high sensitivity

and fast hybridization kinetics. Biotinylated ODN specific to E.coli was immobilized onto an avidin–PANI-modified Pt electrode

[84]. The bioelectrode enabled faster, ultrasensitive, and direct

reagentless detection of E. coli. A PANI–PVS film was prepared at

the surface of the ITO electrode by electropolymerization of aniline

in the presence of PVS, LiClO4, and DNA [85]. The increase in the

conductivity with the increased concentration of PVS was attributed

to an acidic microenvironment for PANI formation. The DNA

biosensor was prepared using the copolymer of PANI and chitosan

[86]. The biosensor showed enhanced electron-transfer properties

toward [Fe(CN)6]3−/4−, which was attributed to the combination

of the excellent conductivity of PANI and the cationic character of

chitosan. PANI nanowires were synthesized electrochemically on the

surface of GCE [87]. Then phosphate-ended ODNs were covalently

attached onto the amino groups of PANI nanowires. The biosensor

effectively discriminated complementary and noncomplementary

DNA sequences. The positive effect of nanomaterials on the PANI

properties was reported. Due to the synergistic effect of MWCNTs

and PANI, a high amount of the DNA probe was immobilized on the

surface of the electrode [88]. Enhanced stability of the PANI film was

observed when it was electropolymerized in the presence of ssDNA-

wrapped single-walled carbon nanotubes (ssDNA-SWCNTs) [89].



ssDNA-SWCNTs served as conductive polyanionic doping agents

and, therefore, enhanced the conductivity and redox activity of the

resulting film.

10.4.1.3 Polythiophene and its derivatives

Polythiophenes and their derivatives are also widely used for the

preparation of DNA biosensors. The disadvantage of these polymers

is difficult electropolymerization of polymers with functional groups

suitable for the immobilization of biomolecules (amino or carboxylic

groups) [90]. Another reported disadvantage is very positive

oxidation potential of monomers [91]. Electropolymerization by

several cycles between 0.0 and +1.1 V was used for the preparation

of terthiophene with an activated ester-terminated side chain [92].

Then the polymer-bearing electrode was incubated in a solution

of aminoalkyl-terminated ODNs. After immobilization of the ODNs,

the authors observed a decrease in the oxidation current as well

as a slight shift of the peak potential. The authors concluded that

immobilized ODNs could cause distortion of the polythiophene

polymer and loss of conjugation. A modified electrode was used

for detecting the presence of mRNA in biological samples. A

poly(cyclopentadithiophene) matrix was tested for electrochemi-

cally controlled DNA delivery [93]. DNA was covalently immobilized

at the surface of the polymer-modified electrode. Quartz crystal

microbalance was used to detect the amount of delivered DNA. The

redox and ion exchange properties of poly(cyclopentadithiophene)

matrix covalently modified with ODNs were investigated using

electrochemical impedance spectroscopy [94]. It was shown that

the ODNs immobilized at the surface of a quartz crystal caused the

blocking of the surface. After hybridization with long target ODNs

a Warburg behavior was restored. DNA was employed as dopant of

PEDOT [95]. Electropolymerized poly(4-hydroxyphenyl thiophene-

3-carboxylate) as cationic polymer was advantageously used for

the electrostatic binding of polyanionic ODNs [96]. Moreover,

interaction between PEDOT and specific ODNs was studied using

electophoresis and spectroscopic methods [97]. It was shown

that together with nonspecific electrostatic interactions, specific

hydrogen binding interactions between polymer and methylated



ODNs appeared, and stable complexes were formed. PEDOT was first

prepared by electropolymerization at the surface of GCE, and then

a DNA solution was spread over the polymer-modified electrode

[98]. DNA was available for the electrostatic binding of Nile blue

as redox indicator. The composite electrode showed electrocatalytic

properties toward the reduction of hydrogen peroxide.

10.4.2 Redox Polymers

Redox-active polymers are conducting polymers containing specific

electrostatically isolated but electrochemically active sites which

can be oxidized or reduced [99]. Redox centers are either organic

molecules or redox-active transition metals covalently bound to

polymer backbone.

10.4.2.1 Quinone-containing polymers

Quinone-containing polymers, namely poly(5-hydroxy-1,4-naphtho-

quinone-co-5-hydroxy-3-thioacetic acid-1,4-naphthoqinone), also

known as poly(JUG-co-JUGA), are also popular for the preparation

of DNA biosensors. In contrast to classical conducting polymers,

such as Ppy or PANI where signal transduction is performed via

redox process of the polymer exchanging anion, in the case of

poly(JUG-co-JUGA) the signal is transduced by the quinone group

in the polymer [33]. The carboxylic group in such copolymers

allows the binding of amino-terminated ODNs, and it shows a very

stable electroactivity in neutral aqueous solutions and can also

work as a hybridization indicator [100]. The copolymer poly(JUG-

co-JUGA) was used for the preparation of DNA biosensors. ODN

was immobilized covalently onto polymeric film from a solution

containing ODN, EDC, and NHS [33–35]. It was shown that, due

to the redox characteristics of the quinone group, the poly(JUG-

co-JUGA) film can be used as an enhanced transducer in ODN

hybridization detection. Interaction and steric effects between DNA

and poly(JUG-co-JUGA) were studied [101]. The authors observed

that only a very short DNA was adsorbed onto the polymeric film

and that the surface concentration of hybrids depended on the

target length. Poly(JUG-co-JUGA)-modified electrodes were tested as

label-free DNA hybridization electrochemical sensors, which used



the electrochemical activity of quinone in the polymer for the

detection of the hybridization event [34, 102]. By the electropoly-

merization of poly(JUG-co-JUGA) onto an MWNT-modified electrode

in nonaqueous media, an interpenetrated conductive network

electroactive in both aqueous and nonaqueous media was produced

[103]. An electropolymerized polyquinone film was derivatized

with glutathione [104]. Glutathione was used as a precursor

for subsequent biomolecule linkage via carboxylic groups. Free

carboxylic groups were first transformed into ester groups using

EDC, and then amino-terminated DNA was immobilized. Because

the polymeric film is a cation exchanger, the negatively charged

DNA cannot be nonspecifically adsorbed at the surface. A solution

of poly(1,4-benzoquinone) prepared by enzymatic synthesis was

cast at the surface of carbon fiber electrodes, and then DNA was

immobilized [105]. The polymer film allowed the hybridization

detection by scanning electrochemical microscopy in the positive-

feedback mode.

10.4.2.2 Redox-active polymers containing organometalicredox center

Redox-active polymers containing ferrocene as redox center were

employed in DNA biosensors. Poly(vinylferrocene) is a soluble

polymer which can be easily deposited at the surface of Pt [106,

107] or graphite working electrode [108] by its electrooxidation

resulting in a less soluble polymer, poly(vinylferrocenium). Such an

electrode can then be advantageously used for the immobilization of

negatively charged DNA. Low nonspecific immobilization of DNA on

this polymer was reported [108]. The electrochemical signal of such

polymer can be used for the detection of the hybridization event

[107]. Another approach was used by Cui et al. [109]. First GCE was

modified with DNA. After drying, the layer of poly(ferrocenylsilane)

was cast at DNA/GCE. It was shown that the DNA at the surface of

GCE enhanced the adsorption of the polymer as well as the electron-

transfer properties. Therefore, the prepared biosensor showed good

electrocatalytic activity toward oxidation of ascorbic acid.

Osmium bipyridyl complexes are known to catalyze the elec-

trooxidation of the guanine base in DNA and also enhance the



detection of DNA [110]. It is convenient to immobilize these

complexes at the electrode surface using a polymer matrix. DNA

can be immobilized either with polymer [111] or redox polymer

electrodeposited onto a DNA layer [112, 113]. Such a polymer forms

a stable and reproducible surface and works as an electron-transfer

mediator.

10.4.3 Nonconducting Polymers

Nonconducting polymers are not so frequently used in DNA

biosensors as conducting ones. They have high resistivity, but

their permselectivity is very useful in preventing interferences

in electrochemical biosensors [114]. In this group of polymers,

polyethyleneimine (PEI) and chitosan (CHIT) are very often used for

the preparation of DNA biosensors.

CHIT is a pseudonatural polymer formed from chitin when

the degree of its deacetylation reaches 50% [115]. Both PEI and

CHIT are cationic polymers with good biocompatibility and high

positive charge density, which allows for easy electrostatic DNA

immobilization. Study of interaction between the DNA molecule and

PEI–copper(II) complexes showed that together with electrostatic

interaction, van der Waals interactions and hydrogen binding are

also employed probably due to the presence of multiple copper(II)

complex molecular units and free amine groups of the polymer

[116, 117]. Electron-transfer kinetics at the PEI–DNA-modified

electrode was studied [118]. It was shown that the surface of

modified electrodes was homogeneous and electron transfer was

slower when PEI formed an external layer. Moreover, further

modification with PEI–gold nanoparticles enhanced the electron

transfer. PEI was used to disperse the MWCNTs, and the screen-

printed electrode (SPCE) was modified with the resulting composite

[119]. MWCNT–PEI formed a layer suitable for the electrostatic

adsorption of negatively charged DNA. DNA/MWCNT–PEI/SPCE was

used for the detection of DNA damage by quinazolines. Interaction

of PEI and CHIT with plasmid DNA (pDNA) on a hanging mercury

drop electrode was compared [120]. Voltammetric studies showed

that each polymer interacts with pDNA by different mechanisms

and that a higher amount of PEI interacts with pDNA than was



observed in the case of CHIT. However, DNA and CHIT can form

stable complexes of specific sizes, influenced by the molecular

weight and pH of CHIT [121]. An assembled film composed of DNA

and CHIT was prepared using the layer-by-layer technique at the

surface of the pyrolytic graphite electrode [122]. CHIT enabled the

effective intercalation of 9,10-anthraquinone-2,6-disulfonate into

the double helix of DNA. A biosensor was successfully applied for

the detection of DNA damage caused by the Fenton reagent. Cu(II)

ions were successfully immobilized in the DNA/CHIT layer due to

the formation of Cu(II)–DNA complexes [123]. This amperometric

biosensor showed excellent electroactivity toward hydrogen per-

oxide with the detection limit of 3 μmol/l. CHIT was also used

to disperse MWCNTs [124–126]. CHIT as GCE modifier partially

blocked the electrochemical response of electroactive species [124].

Introduction of MWCNTs enhanced the electron-transfer properties

of the electrode surface, although values obtained at the bare GCE

were still better. It was shown that CHIT strongly enhanced the

homogeneity of MWCNT deposition onto the electrode surface in

comparison to dispersion in dimethylformamide and MWCNT–CHIT

formed a suitable interface for the immobilization of the DNA

layer in order to study the DNA damage [125]. Electrochemical

properties of screen-printed electrodes modified with composites

of SWCNT–CHIT, MWCNT–CHIT, and (SWCNT-COOH)–CHIT were

studied [127]. It was shown that CHIT alone was able to decrease the

charge-transfer resistance (RCT) of the electrode surface. However,

the decrease in RCT was much more significant when the carbon

nanotube–CHIT composite was used as an electrode modifier. The

best results were obtained in the case of (SWCNT-COOH)–CHIT

because of electrostatic interaction of the negative charge of the

carboxylic group of SWCNTs and the positive charge of CHIT.

Moreover, the (SWCNT-COOH)–CHIT composite was shown as the

best environment for DNA immobilization and was successfully used

for the study of DNA damage caused by lipid peroxidation products.

Overoxidized Ppy(Ppyox) is another example of nonconducting

polymers. It is known that Ppy irreversibly loses electroactivity at

potentials more positive than +1.0 V vs. Ag/AgCl yielding into the

formation of an insulating layer with the porous structure [128]

and the nanoporous diffusion activity [28], both convenient for DNA



immobilization [129]. The Ppy layer was overoxidized potentiosta-

tically at the potential of +1.8 V vs. Ag/AgCl. The prepared DNA–

Ppyox-modified carbon fiber electrode showed excellent sensitivity

and selectivity toward neurotransmitters. In order to increase the

permeability of the Ppy film, an electrochemical overoxidation was

also performed by cycling the potential between 0.0 and 1.3 V vs.

Ag/AgCl until the reversible peak, indicating the Ppy conductivity

disappeared [130, 131].

10.5 Conclusions

Today, there is an increasing interest in the construction and

utilization of DNA biosensors. Successful DNA immobilization plays

a key role in the final efficiency of biosensors. Using polymers

seems to be an elegant way for immobilization of biomolecules.

Moreover, conducting as well as nonconducting polymers not

only represent a matrix suitable for DNA immobilization but also

increase the sensitivity and selectivity of the final biosensor by

avoiding interferences and enhance the stability of the modifier

layer. The thickness of the electropolymerized polymers can be

easily controlled selecting the electropolymerization conditions, and

redox properties can be modified by choosing a suitable dopant

molecule. In recent years, various nanomaterials have been used in

the construction of DNA biosensors. They are usually insoluble in

most solvents, but they can be advantageously entrapped within the

polymer at the electrode surface. Moreover, composites of polymers

and nanomaterials offer a range of new properties, such as enhanced

electron transfer, biocompatibility, and small dimensions with large

surface area, attractive for the miniaturization of DNA biosensors.

List of abbreviations

CHIT chitosan

CP conducting polymer

DNA deoxyribonucleic acid


References 347

EDC 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide

GCE glassy carbon electrode

ITO indium tin oxide

MWCNT multiwalled carbon nanotube

NHS N -hydroxysuccinimide

NPV normal pulse voltammetry

ODN oligodeoxyribonucleotide

PANI polyaniline

PDAB polydiaminobenzene

PEDOT poly(3,4-ethylenedioxythiophene)

PEI polyethyleneimine

Ppy polypyrrole

PSS poly(styrene sulfonic acid)

PVS polyvinylsulfonate

QCM quartz crystal microbalance

SPCE screen-printed carbon electrode

SWCNT single-walled carbon nanotube

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Chapter 11

Electrochemical Transducer forOligonucleotide Biosensor Basedon the Elimination and AdsorptiveTransfer Techniques

Libuse Trnkova,a Frantisek Jelen,b and Mehmet Ozsozc

aDepartment of Chemistry, Faculty of Science, Masaryk University,Kotlarska 2, CZ-611 37 Brno, Czech RepublicbInstitute of Biophysics, v.v.i., Academy of Sciences of the Czech Republic,Kralovopolska 135, CZ-612 65 Brno, Czech RepubliccAnalytical Chemistry Department, Faculty of Pharmacy,Ege University, 35100 Bornova, Izmir, [email protected]

11.1 Introduction

Electrochemical biosensors are usually based on redox reactions

that consume or produce electrons. Such a device can be re-

presented by an indication electrode, which integrates receptor–

transducer element providing selective quantitative analytical infor-

mation; recorded signals are proportional to analyte concentrations.

There are several types of electrochemical transducers, from which



356 Electrochemical Transducer for Oligonucleotide Biosensor

amperometric transducers are most used in biosensors due to

their high sensitivity and selectivity. Except an indication electrode,

the electrochemical system contains other two electrodes, that

is, a reference electrode and an auxiliary electrode [1–4]. Our

approach in electrochemical oligonucleotide (ODN) transducer is

built on the adsorptive stripping voltammetric (AdSV) technique

in connection with elimination voltammetry with linear scan

(EVLS). Generally, EVLS enables the elimination of selected partial

voltammetric currents and the conservation of the other one

contributing to the increase of current sensitivity, the expansion of

electrode potential range (potential window) and the separation of

overlapped voltammetric signals. The basic idea of EVLS procedure

lies in the different dependencies of various voltammetric current

components on the scan rate. The elimination result can be achieved

by a function obtained by linear combination of total voltammetric

currents measured at different scan rates [5, 6].

11.2 Theoretical Fundamentals of EliminationVoltammetry with Linear Scan (EVLS)

11.2.1 Elimination Functions

Fourteen years ago, the theory of elimination voltammetry with

linear scan (EVLS) was published and experimentally verified for

selected electrode systems [5, 6]. To this date, the method has

been applied not only in electroanalytical chemistry, but also in the

study of electrode processes of inorganic and organic electroactive

substances at mercury, silver, or graphite electrodes [7–20]. EVLS

can be considered as a mathematical model of the transformation

of current–potential curves capable of eliminating certain selected

current components while securing the conservation of others by

means of elimination functions. For the calculation of the elimina-

tion functions, two or three voltammetric curves at different scan

rates should be recorded under identical experimental conditions.

It means that the linear sweep voltammetric (LSV) curves have to

be recorded with the same potential step, so that the I –E data sets

obtained for the same number of points on the potential axis, and


Theoretical Fundamentals of Elimination Voltammetry with Linear Scan (EVLS) 357

for the same potential range and the data are not influenced by the

current offset. One scan rate is taken as a reference, while more

scan rates are chosen as selected multiples of the reference scan

rate.

For elimination procedure, two necessary assumptions must be

fulfilled:

1. The total current resulting from different individual processes

such as diffusion, adsorption, and kinetics is formed by the sum

of these particular currents:

I = Id + Ic + Ik,

where Id, Ic, and Ik are the diffusion, charging, and kinetic

currents, respectively.

2. The particular currents eliminated are expressed as the product

of two independent functions:

I j = Y j (E )Wj (ν),

where Y j (E ) is the electrode potential function and Wj (ν) is the

scan rate function.

The scan rate function has the form of a certain power of x of the

scan rate. For example, for a substance transported to an electrode

only by diffusion, the rate power coefficient of 1/2 corresponds to

the diffusion current Id, while x = 1 or 0 holds for the charging

current Ic, or the kinetic current Ik, respectively [5–7, 21]. According

to the second condition of the elimination procedure, the particular

currents take the form

Id = Yd(E )v1/2, Ik = Yk(E )v0, and Ic = Yc(E )v1,

where Y j (E ) of the individual current characterizes a proportion-

ality which is independent of scan rate at the selected potential

value. It has been proved that for the elimination function f (I ) in

addition to the total current at a reference scan rate I , the total

currents for half and double of its value, I1/2 and I2, are suitable

[5, 7, 13, 21]. EVLS functions have been used for the different

combinations with the same scan rate ratio (integer 2) for more than

13 years. The types of six elimination functions are presented in

Table 11.1.



Table 11.1. Types of elimination functions

EVLS Function Characteristics Equation f (I) EVLS Equations

E1 Id �= 0; Ik = 0

a1 I1/2 + a2 If (I ) = − 3.4142I1/2 + 3.4142I

E2 Id �= 0; Ic = 0 f (I ) = 4.8284I1/2 − 2.4142I

E3 Id = 0; Ik �= 0 f (I ) = 3.4142I1/2 − 2.4142I

E4 Id �= 0; Ik = 0; f (I ) = − 11.657I1/2 + 17.485I

Ic = 0 − 5.8284I2

E5 Id = 0; Ik �= 0; a1 I1/2 + a2 I f (I ) = 6.8284I1/2 − 8.2426I

Ic = 0 +a3 I2 + 2.4142I2

E6 Id = 0; Ik = 0; f (I ) = 4.8284I1/2 − 8.2426I

Ic �= 0 + 3.4142I2

Id, Ik, and Ic are the diffusion, kinetic, and charging currents, respectively; a1, a2, and a3 are the

elimination coefficients; and I1, I2, and I3 are the total currents measured at three different scan

rates (v1, v2 = vref and v3).

Generally,

(i) EVLS functions can be set up for the different selected ratios

of scan rates. Then the new coefficients a1,a2 for E1, E2, E3 or

a1,a2, a3 for E4, E5, E6 EVLS functions must be calculated [22].

(ii) The elimination procedure is not limited to the number of

particular currents. It can choose currents with different

dependences on scan rate and calculate the corresponding

elimination functions.

When the above two conditions are not fulfilled, the elimination

function obtained from the experimental voltammograms does not

correspond to the theoretical elimination function, and usually a

distortion of elimination curves may be observed. This distortion

can be used for the electroanalytical determination of some

depolarizators. A large increase in the sensitivity and resolution was

found in the case of the simultaneous elimination of charging and

kinetic currents (Ic, Ik), while conserving the diffusion current (Id)

— the EVLS function E4 (Table 11.1). According to the behavior of

electroactive species, there are two types of transformation of I –Ecurves for an irreversible redox process:

(a) The transport of the electroactive species to the electrode is

controlled only by diffusion and in comparison to the measured



voltammetric signal, EVLS E4 provides a higher and narrower

signal (confirmed by the theory in Ref. [5]).

(b) The electroactive species are adsorbed on the electrode surface

before the electron transfer and EVLS E4 gives the special signal

which is important for the sensitive ODN detection.

11.2.2 EVLS of Adsorbed Species

As already mentioned above, the best EVLS E4 signal was observed

for the electroactive particle, which is pre-adsorbed on the electrode

surface and undergoes an irreversible electron transfer. This

elimination signal corresponds to a well-developed and well-

readable peak–counterpeak. The theoretical curve (Fig. 11.1) has

been calculated according to the equations for the irreversible I –Ecurve of totally adsorbed electroactive species [4]. The theoretical

form of the peak–counterpeak was experimentally verified by the

means of homo-ODN (adenine nonamer), which is strongly adsorbed

on a mercury electrode.

From the analytical point of view, two aspects are important

for an ODN transducer. First, the EVLS signal obtained for totally

Figure 11.1. (a) Theory: LSV and EVLS voltammograms. (b) Experiment:

LSV and EVLS voltammograms of homo-ODN (dA9) in acetate buffer (pH

5.3). f (I ): elimination function E4 for simultaneous elimination of kinetic

and charging currents, and conserving the diffusion current. I p and I p + Icp

are peak and peak–counterpeak heights, respectively. Scan rates for EVLS:

100, 200, 400 mV/s, reference scan 200 mV/s, time of accumulation 90 s,

and potential of accumulation −100 mV vs. Ag/AgCl/3M KCl.



adsorbed electroactive species is seven to ten times higher than

the original voltammetric signal. Second, the shape of the signal

allows the subtraction as the distance between the current

minimum and maximum and does not require any other baseline

correction.

11.2.3 Single and Double Mode of EVLS

The above-mentioned EVLS procedure corresponds to the sin-

gle mode, and functions eliminating two currents require three

voltammetric curves measured at three different scan rates. When

this elimination procedure is repeated three times using LSV

curves measured at five different scan rates (v1/4, v1/2, v, v2, v4),

for example, 25, 50, 100, 200, and 400 mV/s, respectively, the

double EVLS function E4 is obtained, where Id �= 0, Ik = 0, and

Ic = 0.

I1/4

I1/2

I

⎫⎬

⎭f (I ) = a1 I1/4 + a2 I1/2 + a3 I

I2

⎫⎬

⎭f (I ) = a1 I1/2 + a2 I + a3 I2

I4

⎫⎬

⎭f (I ) = a1 I + a2 I2 + a3 I4

⎫⎪⎪⎪⎪⎪⎪⎪⎪⎪⎪⎬

⎪⎪⎪⎪⎪⎪⎪⎪⎪⎪⎭

double f (I ) =a1(a1 I1/4 + a2 I1/2 + a3 I ) +a2(a1 I1/2 + a2 I + a3 I2) +a3(a1 I + a2 I2 + a3 I4)

⇓double f (I ) = ad1 I1/4 + ad2 I1/2 + ad3 I + ad4 I2 + ad5 I4

Equations corresponding to the double elimination functions,

eliminating two current components and conserving one current

component, are shown in Table 11.2. It should be noted that error of

double EVLS is relatively high, and therefore it is necessary to work

carefully with it. On the other hand, a voltammetric signal increases

by more than one order (Fig. 11.2). Moreover, the separation of

overlapped voltammetric signals in the double EVLS mode is much

more successful [23].



Table 11.2. Double EVLS functions eliminating two current compo-

nents

EVLS Function Characteristics Double EVLS Equations

E4 Id �= 0; Ik = 0; Ic = 0 d f (I ) = 135.9I1/4 − 407.7I1/2 + 441.6I

− 203.8I2 + 33.97I4

E5 Id = 0; Ik �= 0; Ic = 0 d f (I ) = 46.63I1/4 − 112.6I1/2 + 100.9I

−39.80I2 + 5.830I4

E6 Id = 0; Ik = 0; Ic �= 0 d f (I ) = 23.31I1/4 − 79.60I1/2 + 100.9I

−56.28I2 + 11.66I4

double EVLS

I (μA) EVLS

df(I)

E (mV)

or f(I)or df(I)

Figure 11.2. Linear sweep (black), EVLS (blue), and double EVLS (red)

voltammograms of homo-ODN (dA9) in acetate buffer (pH 5.3). d f (I ) is the

double elimination function E4 for simultaneous elimination of kinetic and

charging currents, and conserving the diffusion current. Scan rates for EVLS:

50, 100, 200, 400, and 800 mV/s, potential step 2 mV, reference scan 200

mV/s, time of accumulation 90 s, and potential of accumulation –100 mV

vs. Ag/AgCl/3M KCl. Reproduced with permission from Mikelova, R., et al.,Double elimination voltammetry of short oligonucleotides, Electroanalysis19, 1807 (2007). Copyright Wiley-VCH Verlag GmbH & Co. KGaA. See also

Color Insert.



11.3 EVLS Increasing the Transducer Potential Range

It has been known that the reduction signals of nucleobases are

overlapped in a wide interval of pH by catalytic hydrogen evolution.

Mixtures of adenine (A) and cytosine (C) have been analyzed at low

concentrations by different methods, for example, by differential

pulse polarography [24] or sinusoidal voltammetry [25]. However,

these methods were not fully successful in resolution of individual

signals. The problem of mixed signals of A and C interfering with

hydrogen evolution has also been evaluated by artificial neural

networks, using linear sweep voltammetry and differential pulse

polarography results [26].

For the resolution of reduction signals of A and C in mixtures,

the EVLS functions eliminating the kinetic current component

and conserving diffusion current component were applied [27].

This approach enables extending a potential range (window) and

monitoring voltammetric signals hidden in the discharge current of

the supporting electrolyte. The essential requirements are fulfilled

by two functions: (i) the EVLS function eliminating the kinetic

current Ik and conserving the diffusion current Id (E1) and (ii)

the function eliminating the kinetic and charging currents (Ik and

Id) simultaneously and conserving the diffusion current Id (E4)

(Table. 11.1). Our results proved that EVLS is an electrochemical

method suitable for the analysis of purine and pyrimidine bases,

providing the reduction signals in the close vicinity of background

electrolyte discharge [27].

11.4 EVLS in Connection with Adsorptive StrippingTechnique

From the definition of AdSV it follows that this method is character-

ized by the nature of the accumulation process, where adsorption

plays an important role [28, 29]. In AdSV, the pre-concentration

step is not controlled by electrolysis, but it is accomplished

by analyte adsorption on the working electrode surface or by

reactions with chemically modified electrodes. From the early

1960s, this technique (in connection with dc polarography and


EVLS in Connection with Adsorptive Stripping Technique 363

oscillographic polarography without at controlled ac method and

mercury electrodes) was successfully applied to biomacromolecules

analysis, especially to DNA and synthetic polynucleotides (reviewed

in [30] and [31]). Later, it was found that AdSV (in connection with

CV and pulse methods [32–34]) represents a sensitive method for

electrochemical analysis of DNA.

Adsorptive transfer stripping voltammetry (AdTSV) was intro-

duced in 1986 as a new analytical procedure based on the

adsorptive pre-concentration of biomacromolecules on an electrode,

the transfer of the adsorbed layer into a background electrolyte

and subsequent voltammetric analysis [35]. The advantages of

AdTSV were summarized as follows: (i) the method utilizes

differences in adsorbability of substances to their separation, (ii)

due to their strong adsorption, analytes (oligonucleotides) can

be separated from complex media, which are not suitable for

voltammetric analysis of the conventional type, (iii) the interaction

of biomacromolecules immobilized on the surface of the electrode

with substances contained in the solution is possible, and (iv)

all mentioned points can be affected by electrode potential

[35].

An even higher difference was found in stirred solution when the

anodic peak of guanine was measured instead of the cathodic one

[33]. AdSV measuremets of nucleic acids or oligonucleotides were

also performed by square wave voltammetry and ac voltammetry

[36, 37]. Details about AdSV of nucleic acids were summarized in

several reviews [34, 38–43].

As the first EVLS application to adsorbed electroactive species,

the adsorptive stripping voltammetry of thermally denatured DNA

(ssDNA) on a hanging mercury drop electrode (HMDE) was

performed. While the LSV signal of ssDNA at low concentrations

gives a slight indication of the cathodic peak (due to the reduction

of adenine and cytosine residues), the elimination function (elimi-

nating Ic, Ik, and conserving Id) provides a clear peak–counterpeak

signal (Fig. 11.1) [7]. Using this EVLS function E4 it is possible to

determine DNA at concentrations below micrograms per milliliter.

In comparison to the SWV (square-wave voltammetric) signal,

the EVLS signal of ssDNA is one and a half times higher. It was

demonstrated that EVLS, in relation to the accumulation of adsorbed



DNA, can considerably contribute to the electrochemical analysis of

nucleic acids [7].

11.4.1 AdS EVLS of Homo- and Hetero-oligonucleotides

The EVLS has been frequently utilized in the electrochemical

research of short synthetic homo- and hetero-ODNs [10, 15, 44].

Similar to the nucleobases on HMDE, EVLS has been able to resolve

the overlapped reduction signals of adenine (A) and cytosine (C)

in mixtures of dA9 and dC9 [10]. On the other hand, while EVLS

function E4 provides for the nucleobases only enhanced signals due

to the transport of electroactive species to the electrode surface

controlled only by diffusion, in case of ODNs a typical peak–

counterpeak-shaped signal is observed, indicating the electrode

process of completely adsorbed species (Fig. 11.1). The height and

potential of LSV and EVLS signals were affected by the dA9/dC9

ratio, the time of accumulation, the stirring during the adsorption,

and pH. The best results were obtained when the adsorption of

ODNs was carried out at −100 mV for accumulation time of 120 s

under stirring. While on LSV curves the only one reduction peak of A

and C residues was observed in all ODNs, EVLS yielded two separate

peaks in dependence on A–C representation and pH. Subsequently,

our effort was aimed at the separation of A and C reduction signals

of hetero-ODNs containing nine nucleotides with different A–C

sequences, but with the same C/A ratio. We found that (i) EVLS can

be used for the resolution of reduction signals of A and C located on

the same ODN chain, and (ii) the EVLS signal is influenced by the A–

C sequence in ODN chain and pH [15]. The best resolution of both

A and C signals was observed for ODN with triple adenines in the

central part of the nonamer (Fig. 11.3).

The resolution of reduction signals of C and A residues in hetero-

ODNs (9-mers and 20-mers) adsorbed from a small volume on

a HMDE was performed by EVLS in combination with the AdTS

procedure [45]. The suggested connection represents a new, original

detection method for ODN biosensors and provides the possibility to

distinguish between neighboring and non-neighboring bases in the

ODN chain.



Figure 11.3. LSV and EVLS voltammograms of hetero-ODN

5’-CCCAAACCC-3’ in phosphate buffer (pH 6.2). f (I ): elimination function

E4 for simultaneous elimination of kinetic and charging currents, and

conserving the diffusion current. Scan rates for EVLS: 100, 200, 400 mV/s,

reference scan 200 mV/s, potential step 2 mV, and time of accumulation

90s at –100 mV vs. Ag/AgCl/3M KCl. Reproduced with permission

from Trnkova, L., et al., Application of elimination voltammetry to the

resolution of adenine and cytosine signals in oligonucleotides II. Hetero-

oligodeoxynucleotides with different sequences of adenine and cytosine

nucleotides, Electroanalysis 18, 662 (2006). Copyright Wiley-VCH Verlag

GmbH & Co. KGaA.

It was found that the AdS EVLS is capable of reflecting

small differences in the sequences and of distinguishing adjacent

and nonadjacent bases in the ODN chain. Depending on pH the

substantial changes in EVLS signals were observed in the case of

ODN containing a triplet of As and Cs. Alternating A and C in

ODN chains has resulted in weakening of noncovalent interactions

(i-motif) and in decreasing of efforts to form a chain of ODN

multiplexes. The worse separation of A and C signals can indicate

that ODN chain contains A at its end (Fig. 11.4).

As shown in Fig. 11.4, EVLS sensitively reflects the change in

sequence of the ODN chain. Moreover, the EVLS peak–counterpeak

signal is about 5 times higher than the original LSV signal.



Figure 11.4. LSV and EVLS voltammograms of three hetero-ODN: 5’-

CCCAAACCC-3’ (red), 5’-CACCACCAC-3’ (blue), and 5’-ACCCACCCA-3’(green)

in phosphate buffer (pH 6.2). f (I ): EVLS E4 for simultaneous elimination

of kinetic and charging currents, and conserving the diffusion current.

Scan rates for EVLS: 100, 200, 400 mV/s, reference scan 200 mV/s,

potential step 2 mV, time of accumulation 90 s at −100 mV vs. Ag/AgCl/3M

KCl. Reproduce with permission from Trnkova, L., et al., Application of

elimination voltammetry to the resolution of adenine and cytosine signals in

oligonucleotides II. Hetero-oligodeoxynucleotides with different sequences

of adenine and cytosine nucleotides, Electroanalysis 18, 662 (2006).

Copyright Wiley-VCH Verlag GmbH & Co. KGaA. See also Color Insert.

Our results showed that EVLS in connection with the adsorption

procedure (adsorptive or adsorptive transfer stripping, i.e., AdS or

AdTS) is a useful tool for qualitative and quantitative studies of

short oligonucleotides and can be used as a proposed transducer

for the electrochemical sensor. EVLS sensitively reflects not only the

sequence of nucleobases in the ODN chain, but also the structure of

ODN, which can be changed on electrode surfaces.

11.4.2 AdS EVLS of Hairpins

Hairpin structures in DNA and RNA consisting of stem and loop

regions occur naturally not only in ssDNAs and RNAs but also in

double-stranded DNAs (dsDNAs), and they have an important role

in many biological processes. They play a major role in expansion

events, mainly in the case of triplet-repeated expansion diseases

(X syndrome, Huntington disease, Friedreich ataxia). The short

fragment d(GCGAAGC) has been found in the replication origins of

phage φX 174 and herpes simplex virus, in a promoter region of an



Escherichia coli heat-shock gene, and in rRNA genes. Except spectral

and thermodynamic analysis (CD, NMR, and calorimetry), this

heptamer was studied electrochemically (CV, LSV, EVLS) [46]. On

mercury electrodes the hairpin d(GCGAAGC) provides voltammetric

reduction signals of A and C, and oxidation signals of G. Both signals

have been studied in dependence on pH, accumulation time, scan

rate, and loop sequences. The AdS EVLS was employed for the

determination of the detection limit (2 nM), which was verified

by multidimensional voltammetric analysis using Fourier transform

in combination with the confidence ellipse statistic method. Our

results showed the difference in electrochemical behavior of DNA

and RNA heptamers (Fig. 11.5).

While RNA hairpin (Fig. 11.5b) provides one anodic G peak, DNA

hairpin gives two G peaks (Fig. 11.5a) whose heights depend on pH.

This phenomenon is very interesting because guanine-containing

compounds on mercury electrodes provide a single anodic peak G,

which corresponds to the oxidation of reduction product generated

(a) (b)

Figure 11.5. Application of AdS EVLS in the research of DNA and

RNA hairpins (5’-GCGAAGC-3’). LSV and EVLS of anodic signal of G in

(a) heptamer DNA and (b) heptamer RNA in a concentration of 1 μM

(phosphate–acetate buffer, pH 5.3). LSV parameters: scan rate 200 mV/s,

potential step 2 mV, accumulation time 90 s, and time of accumulation 90

s at −100 mV vs. Ag/AgCl/3M KCl. EVLS E4 utilized three scan rates: 100,

200, and 400 mV/s.



at negative potentials [47–49]. The difference in electrochemical

behavior between DNA and RNA mini-hairpins may be explained by

the conformational difference (DNA B form and RNA A form) in the

stem structures [50].

11.5 EVLS of Nucleobases and Oligonucleotides in thePresence of Copper Ions

The purine ring is reducible on mercury electrodes in slightly

acidic medium in a wide pH scale. The electrode redox mechanism

is known and was reviewed [51, 52]. Electrochemical analysis

based on adsorptive properties of long ODNs containing purine

nucleobases, where the transfer technique involves an electrode

transfer step, cannot be applied to monomeric units of nucleotides

or nucleobases because these substances have much less absorba-

bility on electrode surfaces compared to long ODNs. New analyt-

ical approaches were developed to overcome this disadvantage.

One possibility is the interaction of purine nucleobases or their

derivatives with metal ions for example, Cu(II) ions resulting

under suitable conditions in the formation of the complex Cu(I)–

purine. In this reaction, the required monovalent copper ions are

generated electrochemically in the vicinity of electrode surface. The

formed complex is adsorbed on an electrode surface and in the

following reaction step is stripped from the surface by changing

the potential either cathodically (mercury electrodes) or anodically

(carbon electrodes). In both cases, the stripping process resulted in

the formation of a new peak on the voltammetric curve and in the

enhancement of the corresponding redox signal. Using EVLS to the

Cu(I)–purine complex analysis, the more sensitive determination of

purine derivatives was achieved.

11.5.1 Mercury and Mercury-Modified Electrodes

Under specific conditions, adenine forms an intermediate Cu(I)–

adenine species which is sparingly soluble and adsorbs strongly on

the mercury surface [53–55]. The reaction involves electrochemical

reduction of Cu(II) to Cu(I) at a suitable potential and the reaction


EVLS of Nucleobases and Oligonucleotides in the Presence of Copper Ions 369

of Cu(I) with purine bases forming sparingly soluble compounds

that are accumulated on the surface of mercury electrodes. Using

cathodic stripping voltammetry (CSV), the Cu(I) in Cu(I)–purine

complex is reduced to Cu(0), and this very sensitive reaction

is monitored. Under optimal conditions (accumulation potential,

accumulation time, scan rate, copper concentration, and pH), the

ultra-trace CSV determination of adenine and guanine was done

by Farias [56–58]. DosSantos et al. [59] showed that the Cu(I)–

purine complex on HMDE can also be oxidized to Cu(II) using anodic

stripping voltammetry (ASV). There are other examples where a

combination of Cu(II) and purine derivatives was used for sensitive

AdSV determination, for example, xanthine and its derivatives [60],

guanine [61], or methylated guanines [62].

In the case of ODN, the formation of the corresponding Cu(I)–

purine complex is suppressed and its determination is possible after

the release of nucleobases from its chain by acid hydrolysis. Purine

nucleobases from an oligonucleotide chain can be released under

acid hydrolysis, for example, 0.5 M perchloric acid, at a temperature

of 75◦C for 30 min. Under these conditions, only purine bases

are released from the oligonucleotide chain. Then the sample is

cooled and neutralized, and aliquots are mixed with the background

electrolyte for voltammetric measurements.

In our recent experiments, we have studied the determination

of adenine (A), adenosine (Ado), and hydrolyzed adenosine (hAdo)

in the presence of Cu(II) ions using LSV and EVLS in connection

with the adsorptive stripping technique [63]. The differences in

the electrochemical behavior of A and Ado were found to be

dependent not only on the presence of copper ions, scan rate,

adenine concentration, and pH, but also on the accumulation

time and potential where a Cu(I)–adenine complex is formed. A

deeper evaluation of voltammetric responses was carried out by

EVLS using function E4, eliminating charging and kinetic current

components and conserving the diffusion current component. This

function was capable of enhancing the current sensitivity of LSV

peaks and of detecting electron transfer in adsorbed state. The

irreversible electrode process of a totally adsorbed electroactive

species is indicated by means of a peak–counterpeak signal

(Fig. 11.6).



Figure 11.6. AdS LSV and EVLS E4 of hydrolyzed adenosine on HMDE

in the presence of 20 μM Cu(II). Scan rates of 125, 250, and 500 mV/s

and potential step 5 mV. Reference current (black line) at a scan rate of

250 mV/s, accumulation time 120 s, accumulation potential –0.3 V, 0.1 M

acetate buffer, pH 5.1. Reproduced with permission from Jelen, F., et al.,Voltammetric study of adenine complex with copper on mercury electrode,

Electroanalysis 21, 439 (2009). Copyright Wiley-VCH Verlag GmbH & Co.

KGaA. See also Color Insert.

Results show that EVLS is a useful and sensitive tool not only for

both qualitative and quantitative microanalysis of adenine by means

of Cu(I) ions but also for revealing details in corresponding electrode

processes.

Voltammetric measurements confirm that Hg-modified carbon

electrodes are suitable for sensitive electrochemical detection of

ODN compared to mercury electrodes. In the presence of the copper

ions, these electrodes modified by a mercury layer were used for the

detection of a picomolar quantity of ODN. The electrochemical step

includes a potential-controlled reduction of the copper ions Cu(II)

and accumulation of the Cu(I)–purine base residue complex on the

Hg-modified carbon surface. The proposed electrochemical method

can be used for the determination of different ODN lengths because

the stripping current peak of the electrochemically accumulated

Cu(I)–purine complex increased linearly with the length of ODN. The

optical microscope images were used for the visualization of the

surface morphology of the bare and Hg-modified carbon electrodes

[64].


EVLS of Nucleobases and Oligonucleotides in the Presence of Copper Ions 371

11.5.2 Solid Electrodes

There are many types of solid electrodes which are used for the

determination of purine nucleobases, their derivatives, and ODNs

containing purines in the presence of Cu(II). The glassy carbon

electrode (GCE) was used for the electrochemical anodic stripping of

adenine and guanine in Cu(II) solution [65]. It was found that Cu(II)

can be reduced to Cu(I) and the generated Cu(I) reacts with A and G

to accumulate on the GCE as an insoluble compound. Reoxidation of

Cu(I) to Cu(II) at positive potentials gives a large oxidation current

for the base. The same electrode was used for an ultra-trace assay of

some derivatives of nucleic acid bases in Cu(II) solution. Promising

results were obtained also for xanthine determination [66]. The

copper solid amalgam electrode is suitable for a sensitive analysis of

A at very low concentrations. Compared to HMDE, the voltammetric

peak resulting from reduction of the Cu(I)–adenine complex with

the increasing concentration of A shifted to more negative potentials,

indicating the adsorption of this complex on the electrode surface

[67].

Using a paraffin-impregnated graphite electrode (PIGE) and

mercury-modified pyrolytic graphite electrode with basal orien-

tation (Hg-PGEb) Cu(I)–purine complex was studied by LSV in

connection with EVLS [68]. According to the elimination function

E4, the first cathodic peak corresponds to the reduction Cu(II) +e− → Cu(I) with the possibility of fast disproportionation 2Cu(I)

→ Cu(II)+ Cu(0). Anodic stripping voltammetry (ASV) on PIGE and

cathodic stripping voltammetry (CSV) on Hg-PGEb were carried out

at potentials where the reduction of copper ions took place and

Cu(I)–purine complexes were formed.

Electrochemical oxidations of aminopurines (adenine,

2-aminopurine, 2,6-diaminopurine) and their complexes with Cu(I)

were investigated on a pencil graphite electrode (PeGE) by LSV

and EVLS [69]. The anodic process of the sparingly soluble Cu(I)–

aminopurine complex, corresponding to the oxidation of Cu(I) to

Cu(II), takes place in the potential range between 0.4 and 0.5 V. At

more positive potentials, the aminopurines provide voltammetric

peaks resulting from the oxidation of the purine ring. The appro-

priate complex of Cu(I)–aminopurine has a synergic effect on the



scan

Figure 11.7. LSV and EVLS curves of adenine (Ade), 2-aminopurine (2-

AP), and 2,6-diaminopurine (2,6-DAP) (10 μM ) with 20 μM Cu(II) on

PeGE (pencil graphite electrode) in 0.1 M BR buffer, pH 5.1. Reference scan

rate 500 mV/s. Peak OxCom is the anodic signal of Cu (I)–Ade complex,

peak OxAde is the anodic signal of Ade, accumulation potential Ea −0.15 V,

accumulation time ta120 s. Reproduced with permission from Aladag, N., etal., Voltammetric study of aminopurines on pencil graphite electrode in the

presence of copper ions, Electroanalysis 22, 1675 (2010). Copyright Wiley-

VCH Verlag GmbH & Co. KGaA. See also Color Insert.

heights of these peaks. The stability of the accumulated complex

layer was investigated by the AdTS technique. EVLS analysis

using the elimination function E4, eliminating kinetic and charging

current components and conserving the diffuse current component,

provides the possibility of increasing current sensitivity and of

changing peaks into well-readable peak–counterpeaks (Fig. 11.7).

Fadrna et al. [70] has proved that a polished silver solid amalgam

electrode, free from liquid mercury, is a suitable substitute for the

HMDE in CSV analyses of purine bases and of acid-treated ODNs.

The analysis was done at nanomol level in alkaline medium in the

presence of Cu(II). Similarly, the application of gold amalgam–alloy

electrode for a sensitive voltammetric detection of ODNs containing

the purine units within the ODN chains in the presence of Cu(II)


Conclusions 373

ions was described [71]. The proposed electrochemical method was

used either for the detection of different ODN lengths containing

only adenine units (with the number of adenine units within 10 and

80) or for the determination of the number of purine units within

the 30-mer ODNs containing a random sequence segments involving

both the purine and pyrimidine units. A good correlation between

the content of purine units with the whole length of different

30-mer ODNs and the current intensity of the electrochemically

accumulated complexes was found. The sensitive detection of

different ODNs containing the purine units in their chains in the

presence of copper can also be performed at other amalgam

alloys, for example, the platinum amalgam–alloy electrode, copper

amalgam–alloy electrode, and silver amalgam–alloy electrode [71].

Copper-enhanced label-free anodic stripping detection of guanine

and adenine bases in acid-hydrolyzed DNA at anodically oxidized

boron-doped diamond electrode (BDDE) has been published [72].

The BDDE was successfully applied in a three-electrode micro-cell

in which a 50 μL drop of the analyte solution can be efficiently

stirred during the accumulation step by the streaming of an inert

gas. Accelerated mass transport due to the solution motion in

the presence of copper resulted in enhancement of the guanine

oxidation signal, allowing easy detection of 25 fmol of ODNs. It

was also shown that the edge-plane pyrolytic graphite electrode,

whose surface was mechanically roughened, enables voltammetric

analysis of purine nucleobases, acid-hydrolyzed synthetic ODNs, and

a nonhydrolyzed plasmid DNA [73]. In the presence of copper ions,

they caused a strong enhancement of the purine oxidation responses

at fine-polished carbon electrodes.

11.6 Conclusions

EVLS is an unconventional, perspective electrochemical method

capable of eliminating or conserving selected partial currents

(diffusion current, charging current, kinetic current, etc.) from the

total voltammetric current and thereby enhancing the sensitivity

and improving the resolution of the measured voltammetric signals

[5, 6]. EVLS in combination with the adsorptive stripping or

adsorptive transfer stripping (AdS or AdTS) techniques has been



designed for electrochemically fast, sensitive, selective, and low-

cost detection and characterization of surface active compounds

[6, 7, 15, 21]. To obtain sensitive EVLS signals, the elimination

method analyzes voltammetric curves measured at different scan

rates and in connection with adsorptive techniques works with

small amount of samples utilizing strong adsorption of analyte on

the electrode surface. EVLS was applied mostly to the resolution of

reduction signals of adenine (A) and cytosine (C) in short synthetic

homo- and hetero-ODNs [10, 15, 44, 45], but preliminary results

showed that the chosen elimination functions would be useful for

the study of anodic processes, especially for the anodic processes

of guanine [47–49]. For an adsorbed electroactive substance, the

elimination function E4 (the simultaneous elimination of charging

and kinetic currents, and conservation of diffusion current) gives

a well-readable peak–counterpeak, which has been successfully

utilized in the analysis of overlapped reduction signals of A and

C on HMDE [10, 15, 44, 45]. Using the AdTS procedure, ODNs

were immobilized at the HMDE surface from a small drop of the

analyzed solution (5 μL); then the ODN-modified electrode was

washed and immersed into buffer solutions (not containing ODN)

to perform voltammetric measurements [35]. Our new analytical

approach contributed to the transformation of LSV data (overlapped

signal) to EVLS data (resolved signal).

The sparingly soluble complex of Cu(I)–purine reduced at

mercury electrodes and oxidized at carbon electrodes (carbon paste

or carbon pencil electrodes) was recently utilized for the sensitive

detection of purine derivatives [69]. It was found that this complex

has a synergic effect for reduction or oxidation of corresponding

nucleobases because it brings more electroactive materials to the

electrode surfaces. Purine signals processed AdS and AdTS EVLS

and were 15 times more enhanced than the original signal. The

advantage of mercury and carbon electrodes is a good adsorption

capability of this complex on their surface.

In summary, our research showed that the EVLS is not limited to

mercury electrodes, to reduction processes, or to the elimination of

one current component only. Voltammetric signals of purine deriva-

tives at carbon electrodes are amplified using our approaches. Com-

pared to elimination in various combinations (different functions


References 375

elimination, Table 11.1), the electrochemical process can be eval-

uated in detail. Using kinetic current component elimination, the

extending potential window can be achieved. Generally, AdS EVLS or

AdTS EVLS E4 peak–counterpeaks are an order of magnitude higher

than their corresponding LSV signals and, moreover, do not require

baseline correction.

It was found that EVLS is capable to detect (i) minor signals

hidden in major ones, (ii) small changes in ODN structure and the

interaction between ODN and electrode surface, and (iii) potentially

closed signals (resolution of overlapped peaks). On the basis of

the above-mentioned advantages, EVLS in connection with the

adsorption procedure fulfills the requirements for a perspective and

promising tool for qualitative and quantitative studies in bioanalysis

in bio- and nanotechnologies. Therefore, the implementation of

EVLS in electrochemical analyzers should be of great interest.

Acknowledgement

This work was supported by the Ministry of Education, Youth and

Sports of the Czech Republic (INCHEMBIOL MSM0021622412 and

BIO-ANAL-MED LC06035), the Academy of Sciences of the Czech

Republic (grant A400040804), the Czech Grant Foundation GACR

(P205/10/2378), and institutional research plans of the Institute of

Biophysics (AV0Z50040507, AV0Z50040702).

References

1. R. N. Adams, Electrochemistry at solid electrodes Marcel Dekker, New

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Chapter 12

Electrochemical DNA Biosensors forDetection of Compound-DNAInteractions

D. Ozkan-Ariksoysal, P. Kara, and M. OzsozDepartment of Analytical Chemistry, Faculty of Pharmacy, Ege University,35100, Bornova, Izmir, [email protected]

The interactions of some compounds such as anticancer drugs with

DNA have been performed by a variety of techniques. In recent

times electrochemical DNA biosensor systems have been taking an

increasing interest in the analysis of compound-DNA interactions for

understanding the action mechanism of many chemical molecules

due to their high sensitivity, portability, low-cost structure, single-

use property, and compatibility with microfabrication technology.

Based on these electrochemical methods, binding of compounds

onto DNA and/or general DNA damage occurred by these com-

pounds, have been identified by using the voltammetric signals

of guanine, adenine, or related compound molecules. In most of

these applications for the detection of compound-DNA interactions,

anticancer drugs have been studied because of their known effects

on DNA molecule.



380 Electrochemical DNA Biosensors for Detection of Compound-DNA Interactions

12.1 Introduction

12.1.1 Aim of Electrochemical DNA Biosensors

After the first biosensor was described by Clark and Lyons in 1962,

scientists design electrochemical DNA biosensors based on analyti-

cal methodologies for a variety of reasons. They may be interested in

monitoring of DNA hybridization event for the detection of genetic

disease, genetically modified organism, biological warfare agent,

etc. The goal might be the analysis of a solution which contains

trace amounts of hazardous compound that may interact with DNA.

In these examples, electrochemical DNA biosensors (genosensors)

are employed as tools for the identification of DNA sequences

based on the hybridization event and DNA-compound interactions.

In this chapter, the terms and concepts employed in describing

DNA-compound interactions are introduced. Additionally, before

embarking on a detailed consideration of detection techniques and

mathematical equations that gave an idea for the mechanism of the

interaction between compound and DNA, we will mention about the

structure of DNA and possible binding sites of DNA for compounds.

12.2 The Structure of DNA

Deoxyribonucleic acid (DNA) is the most biologically significant

target for electrochemical biosensors for testing of hazardous com-

pounds. Binding of different molecules on DNA and the detection of

DNA damage have been monitored based on both electrochemical

signals of DNA and related compounds. Before the identification

of these interactions, we prefer to give a brief information about

DNA structure due to the importance of its binding sites for

compounds.

The individual DNA molcule which localized in eukaryotic

chromosomes are large polymers and they contain a linear back-

bone of alternating sugar and phosphate residues. DNA molecule

includes the five carbon sugar “deoxyribose,” and consecutive sugar

structures are linked by covalent phosphodiester bridge. Covalently

bonded to carbon atom number 1′ (one prime) of each sugar is a


The Structure of DNA 381

nitrogenous base, for example adenine (A), thymine (T), guanine (G),

or cytosine (C). Adenine and guanine are the member of purines

which consist of two heterocyclic rings of carbon and nitrogen atoms

while cytosine and thymine have a single such ring. A sugar and a

base are composed of a “nucleoside” and if a phosphate group is

attached on it (carbon atom at 5′ or 3′ position), then the main unit

of DNA which is called a “nucleotide” occurs. Phosphate groups have

negative charges [1].

The stable double-stranded DNA structures are held together by

the strong covalent and noncovalent bonds (i.e., hydrogen bonds,

ionic bonds, Van der Waals and hydrophobic forces) which are

theoretically 10 times weaker than covalent bonds. In aqueous

media, the strength of these bonds increase because of the hydrogen

bonds formed between the partially negative oxygen atom and the

partially positive hydrogen atom of water.

While covalent bonds don’t get affected from heat, noncovalent

bonds can be broken reversibly by a high temperature. For molecular

interactions in living cells, this situation is desired because it

plays an essential role in biological functioning. This reversible

interactions are also used in the development of bisensor systems.

The stable duplex DNA molecule is also protected via weak

hydrogen bonds, which occurs between A-T and G-C bases, when

a hydrogen atom is sandwiched between two elecron-attracting

atoms, usually oxgen or nitrogen. It should not be forgotten that

hydrogen bonds can also form between bases within a single-

stranded DNA or RNA molecule dependent on the sequence of

molecules and the distance of its complementary region on the

same strand. As a result of this bonding, hairpin DNA structures

or loops occur which are called as “the secondary structure of

DNA” [1]. Some compounds which have planar aromatic ring in

their chemical structure bind DNA between adjacent base pairs (or

between hydrogen bonds) via intercalation such as daunomycin [2]

and bleomycin.

Most of the DNAs have a B-DNA in living cells. DNAs also have

different helical structures such as A-DNA or Z-DNA. A and B forms of

DNAs are both right-handed helices (clockwise direction) and their

one turns contains 11 (A form) and 10 (B form) base pairs. Left-

handed Z-DNA form has 12 base pairs per turn.



Figure 12.1. The double-stranded DNA structure.

The distance between turns of the helix is called a “pitch” which

is 3.4 nm long.

1 pitch (3.4 nm) = minor groove length + major groove length

The double-stranded DNA molecule also has an antiparallel

nature because the two strands have opposite directions for the

linking of a 3′ carbon atom with a 5′ carbon atom. According to


Natural Electronalytical Characterictics of DNA 383

2 nm

minor groove

major groove

B DNA form

3.4 nm(1 pitch)

Figure 12.2. The double helical structure of B-DNA. One pitch represents

10 nucleotides which are composed of a single turn of DNA. See also Color

Insert.

the Watson-Crick model, base composition of DNA is not random,

total amount of G equals to the total amount of C, and similarly total

amount of A and T are equal based on the complementary rule [1].

12.3 Natural Electronalytical Characterictics of DNA

The electroactivity of purine and pyrimidine bases were found by

Emil Palecek in 1958. While bases have electroactive properties

and they are able to receive reduction and/or oxidation, other

components of nucleic acids such as sugar and phosphate groups are

electroinactive (reviewed in Refs. 3–6). In these reviews, oxidation

parts of A and G [6] and reduction parts of A, C, and G [3–6] were

shown besides the effect of secondary structure of DNA on A and C

reduction signals at mercury electrode.

Carbon-based electrodes are less sensitive to changes in DNA

structure [4, 7, 8]. It was shown that G and A can be detected at



Figure 12.3. Grooves in DNA structure (S: sugar).


Types of DNA Immobilization Methodologies onto Sensor Surfaces 385

Table 12.1. Electroactivity of DNA bases and their detection conditions

Method Base Ox/red Electrode Peak Potential (V) vs. SCE pH

DPV G Ox carbon +1.0 4.8

DPV G Ox carbon +0.8, +0.9 7.4

CV G Ox for reduced HMDE −0.3

product

DPP A Red DME −1.5 Acid/neutral

DPV A Ox carbon +1.2 4.8

DPP C Red HMDE −1.5 Acid/neutral

Abbrevations: DPV is differential pulse voltammetry; CV, cyclic voltammetry; and DPP, differential

pulse polarography. Source: Ref. 6

carbon transducers and C and A at mercury electrode by Trnkova

et al. [8].

The electrochemical signals of nucleic acid bases were shown to

have insufficient sensitivity for DNA analysis in the 1960s, because

of the poorly developed detection devices without software systems.

However, recent advancements in this field started with digital

potentiostats and sophisticated baseline correction techniques in

connection with differential pulse voltammetry (DPV) [9] and

square wave voltammetry (SWV) [10–12]. Therefore, well-defined

voltammetric peaks have been obtained from DNA or RNA at carbon

electrodes in the last decade [13].

DNA adsorption at carbon electrodes reflected by DPV signals

is sensitive to single/double-stranded DNA structure at electrodes.

When compared with the sensitivity of mercury electrodes, carbon

electrodes are less sensitive for conformational changes in DNA [6].

12.4 Types of DNA Immobilization Methodologies ontoSensor Surfaces

Earlier DNA biosensor applications were performed in a solution

phase (DNA solution) [3, 4]. However, in the last decade, researchers

focused on the ordered structure of DNA onto the sensor surface

because of its high sensitivity for detection of target DNA. For this

reason, scientists prefer synthetic and short DNA fragments with



known base sequences related to genetic diseases or microorgan-

isms such as viruses, bacteria, etc.

Typical DNA probes take 15 to 25 base pair long that are able to

detect their target sequences. Besides probe, calf thymus double(ds)

or single-stranded DNA (ssDNA) molecules also immobilized onto

the recognition element of a biosensor.

If we look from the viewpoint of compound-DNA interactions,

dsDNA has been used in numerous sensor applications [14] for the

detection of DNA damage based on electrochemical signal of nucleic

acids especially guanine base.

DNA immobilization step plays the most important role in

determining the performance of an electrochemical genosensor

(DNA-based biosensor) [15]. Control of the DNA binding surface

in terms of surface orientation and coverage is essential for the

sensitive monitoring of DNA–DNA and compound-DNA interactions

by electrochemistry.

12.4.1 Adsorption (Wet Adsorption/ElectrostaticAccumulation)

The adsorption method at controlled potential or without potential

application called “wet adsorption” [16, 17] is the easiest way to

immobilize DNA (or probes) onto carbon transducers [2, 18, 19].

There is no need of special reagents, expensive labeled nucleic acids,

or long experimental steps in adsorption-based immobilization

technique. Hovewer, random immobilization of DNA were obtained

with this technique and nucleic acids bound weakly to the surface as

parallel layers. Additionally, it is possible to aglomerate DNA onto the

surface and when the electrode is rinsed stringently, noncovalently

bound DNA can be removed from the transducer surface.

12.4.2 Covalent Binding to Activated/NonactivatedSurfaces

DNA was first bound to a pretreated electrode via covalent

attachment using carbodiimide molecules by Millan et al. [20] in

1992. After the carbodiimide reaction, DNA was bound to the

surface from its guanine bases. This method was later improved


DNA-Compound Interactions 387

by additional reagent N-hydroxysulfosuccinimide (NHS) in order to

activate carboxyl groups on the carbon electrode.

Single-stranded amino-linked DNA or label-free short DNA

sequences are bound to these groups by their amino tags [21] and

deoxyguanosine residues, respectively [20].

On the other hand, covalent agents can also be applied to the

unpreated carbon surface directly before DNA immobilization onto

activated sites of carbodiimide compounds [21].

12.4.3 DNA Immobilization onto Transducer Surfaces viaAvidin-Biotin Interaction

Biotin binds very tightly to the tetrameric protein avidin (also

streptavidin and neutravidin), with a dissociation constant K d in

the order of 10−15, which is one of the strongest known protein-

ligand interactions, the strength being approximately due to the the

covalent bond [22].

12.5 DNA-Compound Interactions

Voltammetric methods can be used for (1) the identification of

DNA strand breakage and damage, and (2) the determination of

electroactive compounds that specifically bind to DNA (covalently

and/or noncovalently) [23]. For these purposes, electrochemical

DNA biosensors based on the investigation of DNA-compound

interactions has been extensively studied with a number of different

techniques in the past 15 years and this subject has attracted

increasing attention due to its important roles in living organisms

toward the aim of inexpensive and rapid analysis in molecular

biology.

Electrochemical DNA biosensors offer sensitivity, selectivity, and

low-cost detection in this field; therefore, numerous voltammetric

approaches have been developed containing direct electrochemistry

of DNA bases and electrochemistry of DNA-specific electroactive

mediators (reporters) [24, 25].



12.5.1 Types of Molecular Binding to DNA

There are several modes of interactions related to compound-

DNA binding such as covalent binding, noncovalent groove binding,

intercalation, non-specific external association, cross-linking, etc.

However, some of the well-known examples are presented in this

chapter.

12.5.1.1 Electrostatic interactions

Some of the metal ions interact with DNA via electrostatic

binding that are also called as non-specific external association

[26]. Compounds can bind to the negatively charged phosphate

backbone (by covalent or noncovalent binding) or interact with

the electron donor parts of the bases. The strength of these types

of interactions is affected by the charge of the compound, the

hydrophilic–hydrophobic structure of the molecule, and the total

size of the ions. After the interaction between the compound and

DNA, the double helix structure of DNA can be seen damaged as a

separation.

12.5.1.2 Groove binding interactions

Minor grooves in DNA structure are highly attractive regions for

some of the small, flat, and positively charged molecules especially

metal complexes because of their electrostatic and flexible struc-

tures [27]. After this interaction, hydrogen bonding and electrostatic

interactions occur between minor groove bases/phosphate groups

and compounds like Mitramycin [28]. It was also reported that

minor-groove binders have a special chemical structure, usually

containing aromatic heterocycles linked by amide or vinyl groups

with positively charged sections at either ends [29]. Because of

these steric hindrances, only part of metal complexes generally

slot into the minor groove [30]. After a minor groove binding

between the compound and DNA, this formed structure on DNA

is also held together in a stable position by van der Waals forces.

Minor groove interactions do not cause an important and harmful

effect on DNA, according to the reports of Marrington et al. [31].

One of the sample redox active molecule [Co(bpy)3]3+ has been



reported by Mikkelsen’s group as a minor groove binder in biosensor

applications for the determination of the cystic fibrosis �F508

deletion sequence [32].

There are classes of small compounds that bind to DNA from

its major groove via hydrogen bonds. One of the famous anticancer

compound is cis-platin that was found by Rosenberg et al. [33].

This compound was used in many biosensor applications for the

detection of DNA damage [34]. The compound covalently binds

to the DNA from its purine bases (N7 of guanine base, major

groove side) [35] and references within. Although the interaction

mechanism of Ruthenium with DNA is not yet known, it does form

cross-links and groove binder [36]. Two chelates of Ruthenium

complexes are bound to the minor groove of DNA, one chelate of it

is inserted into the major groove.Other metal complexes are cobalt

amines, most of which interact with the major groove of the helix

[37].

12.5.1.3 Intercalation mode

The term “intercalation” was first described in 1982 and it was found

that intercalators shows a high affinity to double-stranded DNA

structures because they prefer to locate between two adjacent pairs

of bases [38]. Intercalator molecules usually have planar aromatic

rings, for example, some antibiotics such as daunomycin destroy

deoxyribose-phosphate structure. These molecules are stabilized by

π -bonds with bases [39].

Intercalators have generally high DNA-binding constants (par-

tition coefficients), and therefore after the interaction between

intercalator compound and double helix, a conformational change

occurs onto DNA that gives a very favorable free energy of

complex formation [14]. On the other hand, in bis-intercalators, for

example, Echinomycin, two intercalative interactions perform via

covalent bonds between aromatic rings of the molecule and DNA

[14].

7-dimethyl-amino-1,2-benzophenoxazinium salt (Meldola’s blue

[MDB]) is also used as an electrochemical hybridization mediator

[40–43] and an analysis of its intercalation mechanism has been

reported by Reid et al. [44].



12.5.1.4 Specific binding for single-stranded DNA

Some of the organic dyes, for example, methylene blue (MB), bind to

DNA from its guanine bases. Ozsoz’s group [45] used MB molecule,

that belongs to the phenothiazine family, as a redox-active indicator

for the electrochemical detection of hybridization based on the

interaction of MB with guanine. Yang et al. [46] also reported

this interaction between guanine and MB by using carbon paste

electrodes (CPEs). A model study was performed for MB binding to

guanine–cytosine base sequences of DNA by Rohs et al. [47]. Enescu

et al. [48] found the MB–guanine complexes with three different

conformations via simulation.

However, Kelley et al. [49] investigated the intercalation of MB

into the thiol-labeled self-assembled monolayer (SAM) containing

dsDNA on the gold electrodes in different experimental conditions.

Tani et al. [50] reported a shift in the peak potentials of MB with

square wave voltammetry by using AuE. MB signal at thiol-labeled

probe-modified AuE was found to be 10 to 15 mV more positive than

the one obtained at thiol-terminated dsDNA-modified electrode.

12.5.2 Detection Techniques for Compound-DNA BindingReactions Using Electrochemical DNA Biosensors

The oxidation/reduction of a compound which shows an affinity to

DNA or intrinsic oxidation signals of guanine/adenine can be used

for detecting the interaction mechanism of related compounds with

DNA at the sensor surface or in the solution [28].

12.5.2.1 Label-free detection based on intrinsic DNA signals(direct detection)

DNA changes by a chemical or its metabolites are of importance for

the carcinogenic processes [51]. The interaction of environmental

carcinogens, drugs, chemical, or the metabolized chemical with

cellular DNA is the first step in the induction of mutations and

carcinogenesis. DNA damage can cause the genetic mutations

which may cause several effects on living functions. Therefore, the

quantification and detection of the compound-DNA interactions and

adducts have major importance in cancer research.



Figure 12.4. Electrochemical detection of substance-DNA interactions

based on sensor surface.

The decrease or increase of the intrinsic guanine oxidation

signal enables the monitoring of the DNA-molecule interactions

electrochemically; these events especially give an idea about the

DNA damage. Additionally, if it is obtained as a new peak in the

voltammogram, then this situation reflects the extent of an adduct

formation [52]. All this qualitative work related to measurements

of the difference in the peak heights of the electrochemical signals

were examined with dsDNA- or ssDNA-modified sensor before and

after the interaction with a compound.

In order to prove that one compound specifically interacts with

guanine and adenine bases, some experiments can be performed

by using synthetic polynucleotides of guanine (poly[G] and adenine

(poly[A]) [53].

In compound-DNA interaction studies, three different assump-

tions could be put forward to explain the decrease in the guanine

oxidation signal: (a) the decrease in the peak height of guanine

could be explained by the covering of oxidizable groups of guanine

while a molecule interacts with DNA, (b) the binding of a chemical

compound to guanine bases, and thus, forming a damage in guanine,

reviewed in Refs. 28 and 54–56, and (c) after the interaction with

the compound, a change in the charge-transfer properties of DNA

[57, 58] could decrease the signal observed from the oxidation of

guanine at CGE surface.



12.5.2.2 Compound-based detection (indirect redoxindicator-based detection)

The compound-based electrochemical detection studies for

chemical-DNA interactions start with the identification of redox

potentials of related compounds by using cyclic voltammetry in

general. The redox peak potential of guanine (+1.0V) [9] is also

evaluated by obtaining compound peaks if both DNA and compound

signals don’t lie in the same peak position in the voltammogram.

Total evaluations are performed with the results of bare and DNA-

modified surfaces together based on both DNA and compound

signals.

In some promising applications about compound-DNA inter-

actions, these molecules can be found as a “DNA hybridization

indicator” because of their different binding behaviors to dsDNA

or ssDNA [59, 60]. This knowledge provides the development of

new drugs and DNA sensors which will further become microchip

devices. Indicator-based electrochemical DNA biosensors contains

electroactive compounds such as methylene blue (MB) [61],

ferrocenylnapthalene diimide [62], several metal complexes such

as cobalt phenanthroline [20], osmium, and ruthenium [63]. In

other applications, Kelley et al. [64] and Boon et al. [57] used

electroactive intercalators which noncovalently bound to DNA for

the detection of different kinds of single-base nucleotide changes.

Some redox-active DNA markers such as ferrocene [65], amino

and nitro-phenyl tags [66], tris-bipyridine complexes of osmium

or ruthenium were applied by Fojta et al. [67] for the detection

of SNPs (single nucleotide polymorphisms). Furthermore, carbon-

based transducers have also been used with several noncovalent or

covalent binding labels on DNA [2, 61, 68].

Panke et al. [69] performed a different approach related to a

competitive binding protocol for the determination of DNA single

base mismatches by using methylene blue in combination with

differential pulse voltammetry technique. Duwensee et al. [70]

reported a strategy for sequence-specific DNA detection by means

of a competitive hybridization assay with osmium tetroxide-labeled

signaling probes.



Marrazza et al. [71] investigated a electrochemical hybridization

indicator “daunomycin” for detecting Apo E polymorphisms in real

PCR. Wang et al. [72] performed the detection of interaction between

daunomycin and DNA in the solution phase and at the sensor sur-

face. Erdem and Ozsoz [60] was showed the other electrochemical

redox-active indicator drug “Epirubicin” which was used for the

detection of mismatched sequences. Hashimoto et al. [73] obtained

that the anodic signals of daunomycin and doxorubicin shifted to

more positive values after DNA immobilization onto basal plane

pyrolytic graphite transducer.

The changes monitored in the electroactive signals of DNA bases

indicate the behavior of compounds toward DNA [74, 75].

For the investigation of interaction mechanism of some com-

pounds as metal coordination complexes with DNA, Bard etal. [76] reported comprehensive electrochemical studies using

cobalt/ferrum phenanthroline or cobalt/ferrum bipyridine. In that

paper, they evaluated limiting shifts and binding constants of

mediator compounds by cyclic voltammetry in the absence and

presence of DNA in solution phase experiments. As a result of

their report, they found those forms (oxidized or reduced form) of

mediator compounds which bind to the DNA molecule with a high

affinity.

Carter et al. [77] also investigated cobalt phenanthroline and

DNA interactions in their previous paper which contained explana-

tions about the dependence of the redox behavior on the nature of

the ligands coordinated to the metal center.

Some other examples about drug-DNA interactions have been

seen in the literature. The antibiotic mitomycin C (MC) and its

interactions with DNA were investgated based on guanine oxidation

signal by Ozkan et al. [78]. Meric et al. [53] described a biosensor

for the detection of interaction between a compound synthesized

as an alkylating anticancer agent and DNA. Jelen et al. [79] found a

redox active bis-intercalator anticancer drug, Echinomycin, and they

showed its interactions with DNA. The intercalator “Adriamycin” and

its in situ interaction with DNA was reported by Brett et al. [80].

These types of interactions have been reviewed by Palecek and

Fojta [54, 55].



12.6 Calculations About Compound-DNA Interactions

In order to investigate the interaction mechanism of a compound

with DNA, different approaches have been presented which can

be used in practical applications such as guanine signal-based

measurement, compound signal-based detection.

The change in the guanine peak is generally used for the

calculations of electrochemical DNA biosensors because guanine is

more easily oxidized than other DNA bases and it can be evaluated

as one of the key criteria for the voltammetric detection of DNA-

drug interactions. The decrease in the guanine signal is estimated

with interactions between compounds and DNA, the current ratio of

guanine (S%) is calculated according to the Bagni et al. [34] equation

which shown below:

S% = (Ss/Sb ) × 100

According to the equation, Ss is the signal ratio of the peak

height of guanine after the interaction with a sample compound,

and Sb is the magnitude of guanine signal after the interaction

with the buffer which is used for the preparation of the related

compound. The guanine oxidation signal obtained with differential

pulse voltammetry (DPV) in absence of a compound served as 100%.

After the interaction between a compound and DNA, if it is obtained

at S > 85% of value, the molecule is considered nontoxic, if it

is obtained that S% value is between 50 and 85, compound is

evaluated moderately toxic, and if the calculation of S% values are

obtained as S < 50%, compound is accepted as toxic.

In order to find an idea about interaction mechanism of a

compound with DNA, the other important value is “partition

coefficient” which was investigated by Millan and Mikkelsen [20]

in 1993. The partition coefficient value is calculated for DNA

biosensors using current signals obtained from probe modified,

hybrid modified, and bare electrodes according to the equation:

Partition coefficient = Compoundbound/Compoundfree

= |(ibound − ifree)/ ifree|Here ifree is the electrochemical peak height of a compound obtained

at bare electrode, and ibound is the oxidation peak current of a com-


Conclusions 395

pound obtained from probe(ssDNA)-modified or hybrid(dsDNA)-

modified electrodes after their interaction with DNA. After the

calculations, if it is seen that a higher value with ssDNA-modified

transducer is obtained than the one with dsDNA-modified electrode,

the related molecule is accepted to show a high affinity to single-

strand DNA structure. In other words, the compound partitions

more into the ssDNA microenvironment than the one of dsDNA as

a result of these calculations.

Carter et al. [76] showed important calculations by using

voltammetric methods for the detection of interaction (electrostatic

or intercalative) of metal complexes with calf thymus DNA. In that

report, binding constant (K n+) and binding region size(s) were

detected from voltammetric data, that is, shifts in potential and

changes in limiting current with the addition of DNA.

The shift in E1/2 value can be used to estimate the ratio of

equilibrium constants for the binding of the oxidized and reduced

forms of ions to DNA molecule. Similarly, for the detection of small

molecules and micelles interactions this value was used [81].

Considering the Nernstian electron-transfer rate for the

reversible redox reactions of the free and bound forms of com-

pounds and the corresponding equilibrium constants for binding of

each oxidation state to DNA yields, for a 1-e− redox process,

E o′b − E o′

f = 0.059 log(K red/K ox)

E o′f and E o′

b are the formal potentials of the oxidized and

reduced forms of a compound couple, in the free and bound forms,

respectively. K ox and K red are the corresponding binding constants

for the oxidized and reduced species to DNA.

As a result, according to limiting shift the ratio of K red/K ox is

calculated and which form of a compound binds to DNA strongly is

determined.

12.7 Conclusions

Electrochemical DNA biosensors (genosensors) developing for the

detection of compound-DNA interactions are very competitive

devices for the aim of detection time and cost, with the possibility



of user-friendly analysis of interaction of various substances such

as carcinogens, mutagens, or drugs with DNA, according to the

requirements of point-of-care analysis.

The specific determination of interaction between DNA and

related molecules is of impotance in the design of the electrochem-

ical genosensors for application in diagnosis tests and in the design

of new drugs, especially for chemotherapy.

In this chapter, the usage of voltammetric techniques for

compound-DNA interactions were shown with detailed information

which contains some key ways to discover unknown molecule-DNA

interaction mechanisms as electrostatic interactions with the DNA

backbone, covalent or groove binding of the double strand of helix,

and intercalation of aromatic compounds between adjacent base

pairs.

When compared to other analysis methodologies such as surface

plasmon resonance (SPR), quartz crystal microbalance (QCM)

or impedance (EIS), voltammetry-based sensors provide short

response time, less-expensive analysis about immobilization of

molecules, and in many analyses they allow real-time measure-

ments.

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Chapter 13

Electrochemical Nucleic Acid BiosensorsBased on Hybridization Detection forClinical Analysis

P. Kara, D. Ariksoysal, and M. OzsozDepartment of Analytical Chemistry, Faculty of Pharmacy, Ege University,35100, Bornova, Izmir, [email protected]

13.1 Introduction

Identification of nucleic acid sequences especially in biological

samples led to early diagnosis of many mutations, microbiological,

and inherited diseases [1]. The detection of specific base sequences

in human, viral, or bacterial DNA holds great importance in

diagnosis of several diseases. Detection of infectious and inherited

diseases at molecular levels provides reliable and early diagnosis.

Traditional diagnostic methods for clinical analysis based on

coupling of electrophoretic separations, radioisotope or fluorescent

labeling are toxic and time consuming. Due to these labor-intensive

methods these are not well suited for routine and rapid clinical

analysis [2]. Recently, there have been major advances in DNA

sequencing technologies [3]. Several methods including various



404 Electrochemical Nucleic Acid Biosensors Based on Hybridization Detection

approaches have been available in genotyping processes such as

polymerase extension [4], oligonucleotide ligation [5], enzymatic

cleavage [6], and flap endonuclease discrimination [7].

An optimum detection method should be compact, highly

sensitive and selective, rapid, high throughput, and cost effective.

Many fast and sensitive methods have been designed to specify only

one or a few target sequences simultaneously. While thousands of

genotypes can be analyzed by using several methods, these devices

are still very expensive and time consuming [8].

Nucleic acid–based biosensors have gained a broad acceptance in

diagnostic testing, sequence specific analysis, DNA drug interactions,

detection of transgenic foods, and microbiological and inherited

diseases in clinical analysis. The growing number of nucleic acid–

based biosensors has stimulated a demand for automated, cost-

effective testing devices that also afford miniaturization of the test

platform [9].

Recently, some reports have indicated that electrochemical

techniques in nucleic acid biosensors are well suited for measuring

hybridization event [10].

Electrochemical DNA biosensor techniques for the detection of

microbiological and inherited diseases devoted to clinical analysis

are presented dealing with past and novel developments in this

chapter. For this purpose; particular emphasis will be given to the

most important approaches for electrochemical genosensing.

13.2 Biosensors

A biosensor is an analytical device that has a recognition capability

for biochemical reactions. It consists of a biological material incor-

porated into a recognition interface connected with a physicochem-

ical transducer [11]. The recognition interface is based on specific

biochemical reactions such as enzyme/cofactor, antigen/antibody,

cell/receptor, and nucleic acids. The physicochemical transducer

recognizes this reaction and converts it to quantitative or semi-

quantitative measurable signal [12]. The aim of the biosensor

techniques is monitoring the biological analytes for in vivo and in


Biosensors 405

vitro applications. Most popular biosensor transductions are optical

[13], piezoelectrical [14], and electrochemical [15] techniques.

The basic scheme of a biosensor device is based on a biochemical

recognition surface, a physicochemical transducer, and a data

analyse equipment. When performing an analysis, biological sam-

ples that specifically interact with its substrates on the surface are

detected by the recognition surface. The results of the interaction

should form changes which can be physical or chemical. After

recognition, the detection signals are converted to another signal by

the transducer that can be analyzed easily. The transformed signal is

amplified and processed for user analysis.

The first biosensor was based on an enzyme electrode and

developed for glucose analysis in 1962 [16]. Then many researches

have focused on biosensing systems. This is mostly due to the

biosensor’s high selectivity and sensitivity [17]. In 1975, Divis

proposed microorganism electrode for determining the alcohol level

in a solution [18]. Also, same year the first glucose biosensor was

produced commercially by Yellow Springs Instruments.

13.2.1 Nucleic Acid Hybridization Biosensors

Nucleic acid is a biosensor which integrates nucleic acid hybridiza-

tion recognition with a signal transducer. Figure 13.1 is a schematic

representation of a nucleic acid biosensor.

The nucleic acid recognition part selectively detects a specific

gene sequence of DNA. A DNA hybridization biosensor uses a DNA

strand of known sequence as a probe of a target DNA sample.

In the last decade there has been a considerable interest in DNA

biosensors due to its significant analytical properties. The most

popular application of DNA biosensors is based on nucleic acid

hybridization detection of specific DNA sequences [19].

Such biosensors have many potential applications — for exam-

ple, identification of genes that are implicated in inherited diseases,

single nucleotide polymorphisms (SNP), and some mutations that

play a major role in causing diseases [20–21], identification of

pathogenic microorganisms which are responsible for infectious dis-

eases [22–23], transgenic organisms for food quality [24], detection



Figure 13.1. Schematic representation of a nucleic acid biosensor. See

also Color Insert.

of DNA damage caused by drugs, toxins, or radiation [25–26], and

many more clinical applications.

Nucleic acid biosensors can be classified on the basis of

their transduction technology. The transducer converts the nucleic

acid hybridization recognition into a measurable analytical signal

[27–28]. Electrochemical, optical, piezoelectrical, acoustical, and

mechanical transducers are among the many types found in DNA

biosensors.

Optical sensors employ optical fibers or planar waveguides

to direct light to the sensing film. The measured optical signals

often include absorbance, fluorescence, chemiluminescence, surface

plasmon resonance (to probe refractive index), or changes in light

reflectivity. Many studies on SPR as an optical method for biosensing

have been carried out because this method allows the measurement

of the kinetics of biomolecular interactions in real time with a high

degree of sensitivity without labeling of the biomolecules [29–30],

however, they cannot be easily miniaturized for insertion into the

bloodstream. Most optical methods of transduction still require a

spectrophotometer to detect any changes in signal [31].

Piezoelectric biosensors are mass-sensitive biosensors which

can produce a signal based on the mass of chemicals that interact

with the sensing film. Quartz Crystal Microbalance (QCM) sensorsa


Electrochemical Nucleic Acid Biosensors 407

re piezoelectrical based sensors that are operated by applying an

oscillating voltage at the resonant frequency of the crystal, and

measuring the change in resonant frequency when the target analyte

interacts with the sensing surface [32]. The QCM method has

been adopted by several groups to detect the DNA hybridization

reaction because of its great sensitivity as a mass sensor capable of

measuring subnanogram mass changes [33–34].

Electrochemical biosensors measure the electrochemical

changes that occur when biochemical element interacts with a

sensing surface of the detecting electrode. The electrical changes

can be based on a change in the measured voltage between the

electrodes (potentiometric), a change in the measured current at a

given applied voltage (amperometric), or a change in the ability of

the sensing material to transport charge (conductometric) [35].

13.3 Electrochemical Nucleic Acid Biosensors

Electrochemical nucleic acid biosensors are based on electro-

chemical transduction of the hybridization event and show great

promise for detection of specific gene sequences related to inherited

and infectious diseases. Electrochemical detection of specific DNA

sequences has an advantage in reducing the size of the total

detection system [36]. The advantages of electrochemical nucleic

acid biosensors include potential of miniaturization, short response

time, ease of use, low cost, and compatibility with microfabrication

techniques [37].

The aim of electrochemical genosensing techniques is to design

DNA systems allowing early diagnosis of microorganisms and poly-

morphisms in clinical analysis. For this purpose several techniques

have been investigated based on recognition of DNA hybridization,

by using electroactive labels, dye molecules, nanoparticles, or label-

free methods.

Electrochemical DNA biosensors are divided into two main

groups:

1. Label-based DNA hybridization detection method

2. Label-free DNA hybridization detection method



13.3.1 Label-Based Electrochemical Nucleic AcidBiosensors

Label-based electrochemical nucleic acid hybridization biosensors

work on the principle of the following groups:

1. Using redox active hybridization indicator which has an affinity

for ss or ds DNA

2. Using labeled signaling probes or labeled target DNA

13.3.1.1 Electrochemical genosensing by using hybridizationindicator

This approach is based on the electrochemical response of a redox

active label changes upon DNA hybridization, when the hybridiza-

tion process occurs due to change of the indicator concentration at

the electrode surface [38]. These redox active labels can be called as

“hybridization indicators” and have high affinity for either ssDNA or

dsDNA to transduce hybridization.

Hybridization indicators have various interaction properties of

dsDNA and ssDNA. Some metal complexes or dyes are intercalator

molecules which interact with hydrogen bonds of dsDNA [39], and

some indicators have selective binding processes onto DNA bases

such as guanine [40].

Intercalator hybridization labels are complex molecules that

have a planar aromatic group. Several methods for indicator-based

electrochemical sequence specific to DNA detection have been

reported. Wang et al. [41] described the hybridization detection

of short DNA sequences related to HIV virus genome due to the

chronopotentiometric transduction of Co(phen) as an hybridization

label. Electrochemiluminescense assays have also been reported by

Carter et al. [42] for specific DNA sequence detection.

Early studies on electrochemical nucleic acid biosensors were

based on electrochemical transduction of redox labels (indicators)

that have significant different behaviors between dsDNA and ssDNA.

These intercalator molecules have higher binding affinity to dsDNA

than ssDNA. Mikkelsen and coworkers investigated this approach

by using Co(phen) as a hybridization indicator. The intercalator

molecule was accumulated at ss and dsDNA at covalently attached



on to glassy carbon electrode surface [43]. Millan, also used

Co(phen) and osminum complexes as hybridization indicator to

detect hybridization. Short oligonucleotides related to cystic fibrosis

diseases were used as probe and target sequences [10]. Unmodified

probe sequence was immobilized onto carbon paste electrode (CPE)

and voltammetric transduction of metal intercalators was moni-

tored after hybridization occurred. Same year Millan’s group studied

on covalently attachment scheme. Oligonucleotides including Poly A,

Poly T, Poly C, and Poly G were used as model case. Carbodiimide

chemistry was first used onto glassy carbon electrode (GCE) surface

for covalently bounding of DNA [2].

This intercalator molecule has been investigated by many

workers, such as Mascini [44–45], Wang [46], and our group [38–

39] in detailed. In 1999, J. Barton’s group first worked on single–base

mismatch detection [19]. Thiol-modified oligonucleotide sequences

were attached on to gold electrode surface and hybridization

occurred with both full-match and mismatch target sequences.

The cyclic voltammetric transductions of intercalator molecules

including ruthenium complexes were monitored.

Mascini and coworkers were focused on detection of real sam-

ples, and they used PCR products related to human Apolipoprotein

E genotypes in 2000 [47]. Graphite screen-printed electrodes

were firstly used for clinical detection as sensor surface. Probe

sequences were adsorbed at SPE and hybridization was determined

by using daunomycin as indicator. Kobayashi et al. [48] investigated

a microelectrode array for simultaneous and multiple analysis. They

designed a sensor which had 32 arrays, and therefore it was possible

to work with several hybridization detection events at the same

time. Hybridization and mismatch detection was performed by using

lineer sweep voltammetric transduction of a commercial redox

active dye molecule as an intercalator. Yang et al. [49] developed

a genosensor for detection of PCR products by using 7-deaza

analogues of guanine and adenine. Cyclic voltammetry was used for

transduction of ruthenium complexes for the detection of E. coli PCR

product. Barton’s group used Rhodium derivates as intercalating

agent for rapid mismatch detection [50].

Our group is also focused on the detection of clinical analysis

based on intercalator molecules and on voltammetric transduction



of intercalator molecules related to specific gene sequences.

Figure 13.2 represents the voltammetric hybridization and mis-

match detection in a PCR amplicon by using intercalator hybridiza-

tion indicator.

A new intercalator dye molecule Meldola’s blue (MDB) was first

used by our group for hybridization detection in a PCR amplicon

Figure 13.2. A schematic representation of voltammetric hybridization

and mismatch detection.



[51]. A PCR sequence was used related to Hepatitis B (HBV) virus

genome. Optimization of hybridization detection was performed

with 17-mer short oligonucleotides. Carbon paste electrodes (CPE)

and hanging mercury drop electrodes (HMDE) were used as

sensor surface, CV and DPV transduction of MDB accumulated

after hybridization between 23-mer capture probe and HBV–PCR

amplicon was monitored for the detection. By using MDB indicator,

Herpes simplex (HSV) virus genome detection and discrimination

of HSV type I and type II viruses were performed in PCR amplicon

[52]. HSV type I and type II have very similar pathogenesis

mechanisms and have a homogeneous genome sequence. Two

types of PCR products related to type I and II which had 12

base differences in 179 base long amplicon were used as target

genomes. 22-mer capture probes related to type I and II had four

base differences between each other, were attached onto disposable

graphite electrode surfaces and hybridization occurred with both

types of PCR amplicons. The detection and the discrimination of

genotyping were accomplished by DPV transduction of accumulated

MDB. Consequently four base differences were detected by using

long PCR amplicon devoted to clinical analysis.

One base mismatch detections in real samples based on MDB

indicator were also performed in our following researches. In 2007,

we developed a genosensor for detection of toll-like receptor 2 (TLR

-2) gene polymorphisms [53]. In this study, one base mismatch

detection was performed in a 267 base long PCR product. Two types

of capture probes were used representing wild-type and mutant-

type genomes. DPV reduction signals of MDB were monitored after

hybridization with denatured amplicon at PGE surfaces. Heterozy-

gous and homozygous discrimination was also performed by using

two types of capture probes. Biosensor selectivity was achieved with

HBV non-complementary (NC) amplicon. Consequently an allele

specific genosensor was developed for SNP detection in this study.

Another polymorphism detection related to Apa I vitamin D receptor

gene was also performed in 2010 [54]. DPV signals of accumulated

MDB indicated hybridization and mismatch detection in 247-mer

PCR sequence.

Some hybridization indicators have chemical affinity to DNA

bases. Our group used another dye molecule, methylene blue



(MB) as hybridization indicator in many researches. MB is an

aromatic heterocycle; although MB is an organic dye molecule and

an intercalating agent, it has a higher affinity to guanine bases

[55]. Enescu et al. [56] investigated the conformation of MB–

guanine complex by molecular dynamics simulation. The position

and orientation of MB–guanine complexes were found to be in three

modes: T-shaped, non-stacked and face to face. Due to this affinity,

we used MB as a hybridization indicator in several works.

Early studies with MB were performed by Barton [57] as

intercalator molecule. In 2000, our group used MB for the first time

as a hybridization indicator which has a strong affinity to guanine

[58]. In this study, 21-mer oligonucletotides related to Hepatitis

B virus (HBV) were immobilized onto CPE and hybridization was

detected after MB accumulation. Voltammetric transduction of MB

reduction was monitored. The comparison of indicator behaviors

between intecalator molecule ruthenium complex and MB was

performed in 2001 [59]. Calf-thymus dsDNA and ssDNA were

immobilized onto CPE electrostatically. CV and DPV transduction of

hybridization indicators were monitored after accumulation.

Figure 13.3 represents the voltammetric detection of hybridiza-

tion in the presence of MB hybridization indicator. Due to strong MB

affinity, voltammetric peak of MB after accumulation with (a) ssDNA

is significantly higher to (b) bare electrode, and (c) sDNA.

The effect of ionic strength onto MB accumulation behavior

was also studied by our group [60]. Chronocoulometric and

voltammetric parameters for MB on binding to DNA at CPE were

monitored. It was found that 10 mM ionic strength is the critical salt

concentration. MB interacts to guanine electrostatically up to 10 mM

NaCl, in the presence of higher concentrations of 10 mM of NaCl, MB

intercalates to hydrogen bounds of dsDNA.

Hybridization and one-base mismatch detection was performed

by using self-assembled monolayer (SAM) on gold electrodes

in the presence of MB indicator first time [61]. 14-mer short

oligonucleotides were immobilized onto Au electrode surface by

using alkanethiol monolayer coupling at surface. Mercaptopropionic

acid (MPA) was used for monolayer production. Voltammetric

reduction signal of MB was monitored for hybridization and

mismatch detection.



Figure 13.3. Voltammetric hybridization detection with MB hybridization

indicator.

Peptide nucleic acid (PNA) is a structural DNA analogue con-

taining an uncharged N -(2-aminoethyl) glycine-based pseudopep-

tide backbone, which has been reported to form Watson–Crick

complementary duplexes with DNA. PNA, originally synthesized

as a gene-targeting antisense drug, has demonstrated remarkable

hybridization properties toward complementary oligonucleotides.

Compared to DNA duplexes, PNA hybrids have higher thermal

stability and can be formed at low ionic strengths. The neutral

peptide-like backbone of PNA provides the basis for the probe

to hybridize to target DNA sequences with high affinity and

specificity [62–63]. Due to these opportunities of PNA molecules,

our group used PNA for hybridization and mismatch detection

in the presence of MB indicator. Short oligonucleotides of PNA

sequences were immobilized onto mercury and carbon electrodes

[64] electrostatically and onto Au electrode by SAM method [65].



Consequently, mismatch detection was accomplished by voltammet-

ric transduction of MB accumulation.

The application of clinical analyses from real samples was

determined with HBV detection. For this purpose; we developed

a genosensor for clinical analysis of HBV based on MB indicator

[66]. In this study, real samples were used first time for medicinal

analyses. 24-mer capture probe related to HBV genome were

immobilized onto CPE electrostatically, the hybridiziation with

PCR amplicon and accumulation of MB was applied. Hybridization

detection was accomplished by monitoring the DPV reduction

signals of MB.

13.3.1.2 Electrochemical genosensing with labeled signalingprobe or labeled target DNA

Another approach to electrochemical biosensing of microbiological

and inherited diseases is to use labels attached onto capture probe

or target sequences. If a redox active molecule such as ferrocene has

been attached to probe sequence, the electron transfer of double-

stranded DNA has been insensitive due to the distance from the

electrode surface [67]. Ferrocene (Fc) and its derivates are attractive

redox active chemicals because of their stability [68]. Yu et al. [69]

prepared ferrocene-labeled oligonucleotides that were conjugated

with uridine. With the same technique, Yu [70] performed an

SNP detection based on DNA/RNA hybrids. Xu et al. [71] used

ferrocenecarboxaldehyde (FCA)-modified ssDNA probes bounded

at chitosan-modified electrode surfaces. Chitosan-modified graphite

surfaces provided a strong binding of probe sequence at the

surface, and hybridization detection was accomplished with DPV

transduction of FCA. 5’-FC–modified hairpin DNA probe was used

for sequence specific detection. Genosensing was performed by

transduction of AC voltammetry and differences in melting points

between FC-modified hybrid and unmodified target sequences [72].

The biotin–avidin system is used in a variety of biotech-

nological and diagnostic applications. It involves a chemical or

genetic (the bio-tag biotinylation) biotinylation step. Mascini used

biotinylated oligonucleotides for electrochemical genosensing [73].

Thiol-tethered capture probe was immobilized onto Au-SPE and



hybridization occurred with target sequence. Biotinylated signal-

ing probe was added onto PCR target and accumulation with

streptavidine–alkaline phosphatese conjugate was monitored by

impedance spectrometry. Based on this method, we used alpha

napthol as indicator [74]. Capture probe was modified onto

graphite surface and hybridization occurred with biotinylated target

sequence. Avidine–alkalinephosphatase complex was coupled with

hybrid and napthyl phosphate was added onto the surface. The

DPV reduction signal of napthol was used as an indicator. For

detection of different bacterial food contaminations [75], Legionella

pneumophila with hairpin DNA probe [76] was performed by using

this system.

13.3.2 Label-Free Electrochemical Genosensing

The main disadvantage of electrochemical nucleic acid biosensors

discussed above is requirement for an indicator to transduce

hybridization. Many scientists have focused on developing label-free

methods for directly monitoring the hybridization event. Wang and

coworkers [77] studied the oxidation signal of guanine base at about

1.00 V. Wang and coworkers [78] have determined that guanine is

the most electroactive base when compared with cytosine, timine,

and adenine. In this study, Wang used a pencil-based renewable

electrode for sensor surface. Former solution based electrochemical

reports have shown that the electron transfer from the uncatalyzed

guanine bases was slow at most electrode surfaces, however,

guanine oxidation could well be observed by using voltammetric

techniques when the guanine was adsorbed onto the CPE [79].

Tomschik et al. [80] observed the oxidation signals of guanine

and adenine at low concentrations of DNA and PNA by applying

chronopotentiometry and voltammetry with a suitable baseline

correction system at pyrolytic graphite electrode (PGE). By using

carbon nanotubes, Wang enhanced the surface area for label-free

detection of hybridization [81]. Prado et al. [82] used boron-doped

diamond electrodes for sensing surface. They monitored guanine

oxidation in ssDNA and dsDNA by cyclic voltammetry.

Label-free genosensing techniques have a great importance for

sequence specific detections in pharmaceutical and environmental



forensic science and clinical analysis. Our group has paid a

significant attention on label-free electrochemical genosensing

techniques for hybridization detection. Our studies are focused on

direct detections of microbiological and inherited diseases, DNA–

drug interactions and SNP analysis. Discrimination of single- and

double-stranded DNA was accomplished by using calf thymus ssDNA

and dsDNA. Electrostatically bounded ss and dsDNA at CPE surfaces

were monitored due to the oxidation signals of guanine and adenine

[83]. The electrochemical determination of hybridization between

DNA probe and target oligonucleotides and polynucleotides were

also accomplished by the dependence of peak heights of guanine and

adenine DPV oxidation signals. Figure 13.4 represents the label-free

electrochemical voltammetric genosensing of hybridization which

we follow in our laboratory.

A strong DNA immobilization method was developed by using

chitosan which is a cationic polymer that forms polyelectrolyte

Figure 13.4. Label-free voltammetric hybridization detection based on

guanine adenine oxidation.



complexes with DNA. Chitosan-modified CPE (ChiCPE) surfaces

were used as a sensing area for direct detection of hybridization in

this study [84]. Calf-thymus ss and dsDNA were immobilized onto

ChiCPE surafces and hybridization between PNA oligonucleotides

was determined by transduction of guanine oxidation. Thereby, a

cost-effective, rapid and direct genosensing method was developed

that provided highly strong DNA immoblization. A label-free SNP

detection was also performed in our laboratory by using PNA

oligonucleotides [85].

The detection of PNA–DNA and DNA–DNA hybridizations were

accomplished based on the oxidation signal of guanine by using

DPV at CPE. It was observed that PNA–DNA hybrids have significant

peak height differences when compared with DNA–DNA hybrids. In

addition, PNA probes have a weaker affinity to mismatch targets,

so detection of point mutation was performed based on guanine

oxidation signals.

Sequence-specific bioelectronic detection of PCR amplicons were

performed with unpurified PCR samples by Lai et al. [86]. GyrB

genes of Salmonella typhimurium were produced in PCR reaction

and detection was performed by applying AC voltammetry. Manalis’

group investigated a label-free microelectronic PCR quantification

[87]. A field-effect microelectronic sensor was developed which was

capable of quanification of DNA during PCR reaction at polylysine

covered surfaces.

Wang et al. [88] described an indicator-free electrochemical DNA

biosensor protocol, which involves the immobilization of inosine-

substituted (guanine-free) probe onto CPE and the detection of

hybrid formation was performed by using the appearance of the

guanine oxidation signal of the target in connection with chronopo-

tentiometric stripping analysis (PSA). Napier et al. [89] also used

inosine substituted probes, in the presence of ruthenium complexes

as hybridization indicator. Macsini [90] has developed an inosine-

based label-free genosensor for identification of mammalian species

by using bovine and sheep PCR amplicons. Guanine-free capture

probes were immobilized onto screen-printed carbon electrodes

(SPE), hybridization between positive real samples of porc, bovine

and sheep sequences were monitored by DPV oxidation signals

of guanine. Kerman et al. [91] monitored guanine oxidation at



about 0.73 V with square wave voltammetry (SWV) at l-cysteine

monolayer modified Au surfaces. 6-mer thymine-tag of the capture

probe was hybridized with the adenine probe, thus left the rest of

the oligonucleotide available for hybridization with the target.

The use of inosine-substituted probes and the appearance of

a guanine signal upon hybridization with the target opened a

new field in electrochemical research. We performed alle-specific

polymorphism detection in real samples by using inosine-modified

probe sequences, called yes/no system. Two capture probes related

to wild-type and mutant-type genoms were immobilized onto elec-

trode surface and hybridizations occured with denatured heterozy-

gous or homozygous amplicons. Favtor V Leiden and Achondropla-

sia G 380R point mutations were performed by this technology [92–

93]. It was observed that homozygous amplicons had only one signal

of guanine with their complementary strands, but on the other

hand, heterozygous amplicons had guanine signals with both probe

sequences [94]. Consequently, by using two different probes related

to both wild-type and mutant genomes, we could achieve rapid and

allele-specific detection. Figure 13.5 is the schematic representation

of voltametric allele-specific genosensing method based on yes/no

system. This method was able to detect down to 51.14 fmol mL1

target DNA. Similiar methods have been developed for the detection

of interleukin-2 DNA [95], Val108/158Met SNP in COMT gene

[96].

Optimizations of hybridization kinetics and washing conditions

including ionic strengths are the key points for effective detection

of microbiological and inherited diseases. Detection of optimum

probe sequence relative position in a long amplicon based on yes/no

system was studied [97]. 18-mer inosine-modified three capture

probes were chosen from several parts of HBV genome amplicon.

Two sequences were 5 base distance from primers, the 3rd sequence

was in the middle of the amplicon. The probes were guanine-free

besides including five cytosines in each sequence thus called as co-

equal captures. Capture probes were immobilized onto electrode

surfaces via carbodiimide chemistry. After hybridization occured;

optimum probe sequence position was identified by using the

differences between the responses of guanine oxidation signals. It

was observed that probe sequences chosen from the beginning and



Figure 13.5. Schematic illustration of electrochemical label-free allele-

specific genosensing method.

end part of the amplicon (close to primers) caused duplex formation

at the posterior of the long sequence, however, the probe chosen in

the middle section of the amplicon prevent the duplex formation

and stabilize the amplicon sequence for hybridization and provide

an optimum diagnosis.

Direct bioelectronic detection of multiple point mutations in

Mycobacterium tuberculosis amplicons related to rifampin drug

resistance was perfomed [98]. In recent studies, it was found that

95% of RIF-resistant bacteria strains possess mutations within



Figure 13.6. Schematic presentation of electrochemical genosensing of

multiple point mutations in PCR amplicon.

the 81-bp hotspot region between the 507th and 533rd codons

of the rpoB gene. Five different inosine-modified capture probes

represented several parts of rpoB gene area including several

SNPs were immobilized onto electrode surfaces. Hybridization

and mismatch detection was performed by monitoring guanine

oxidation. In conclusion, rapid, cost-effective, highly sensitive, and

sequence-specific array system which is capable of multiple SNP

detection at the same time was developed. This method was able to

detect down to 18.65 fmol/mL. Figure 13.6 represents detection of

multiple point mutations in mycobacterium tuberculosis amplicons

based on label-free electrochemical genosensing. Five capture

probes (P1, P2, P3, P4, P5) representing several parts of amplicon

were immobilized onto different sensing areas. After hybridization

with an amplicon, different responses of guanine oxidations were

obtained due to the region of the SNP.

13.4 Conclusion

Throughout this chapter, we demonstrated label-based and label-

free electrochemical genosensing techniques for the detection of

microbiological and inherited diseases devoted to clinical analysis.

The sensor technology is relatively cheap to produce, easily


References 421

stabilized, and voltammetric technique is stable. When compared

with conventional methods, it can be observed that electrochemical

techniques are also capable of sequence specificity and allele

specificity, time-consuming, and highly sensitive. Further research-

based mutation detection methods are under progress in our

laboratory.

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Chapter 14

Nanomaterial-Based ElectrochemicalDNA Detection

Ronen Polsky, Jason C. Harper, and Susan M. BrozikBiosensors & Nanomaterials, Sandia National Laboratories,PO Box 5800, MS-0892, Albuquerque, NM 87185, [email protected]

The combination of nanomaterials and biomolecules has led to a

new generation of DNA sensing devices. Taking advantage of the

size-dependent properties of nanomaterials and the unique inter-

facial phenomenon that result in their coupling with electrochem-

ical transducers, many different biosensing strategies have been

realized. The use of nanoparticles, various nanowires, nanotubes,

nanorods, etc. have all been incorporated into novel DNA sensing

schemes. Thus, the field of biotechnology has recently witnessed

extensive progress in the use of nanomaterial-based electrochemical

DNA sensors.

14.1 Introduction

The field of biotechnology has witnessed extensive progress over

the past decade in the use of nanomaterials to develop novel

biosensors and electrochemical bioassays [1]. Perhaps the most



428 Nanomaterial-Based Electrochemical DNA Detection

extensive growth in nanobiotechnology has been in the area of DNA

analysis. Electrochemical DNA biosensors are powerful tools for

nucleic acid analysis because they are often simple, rapid, reliable,

and cost effective. The transduction of DNA hybridization events

into electrical signals to construct sensing devices has potential

applications ranging from molecular diagnostics, drug screening,

medical diagnosis, food analysis, and environmental monitoring.

As a material system approaches molecular dimensions, it can

exhibit novel optical, electrical, mechanical, and chemical properties

that can be further manipulated and tailored by varying the size,

shape, and composition of the nanoscale material. The unique

electronic and structural properties of nanomaterials have enabled

new ultrasensitive electrochemical sensors [2] that would not have

been possible without the nanomaterials’ unique properties. The

progress made toward chemical functionalization of these materials

has led to successful interfacing of biomolecules, such as DNA,

with electrochemical signal transduction platforms providing an

enhanced electrochemical response.

For example, a number of different electrochemical techniques

such as cyclic voltammetry, differential pulse voltammetry, and

potentiometric stripping analysis can be used in combination with

nanomaterials to quantitatively detect extremely low concentrations

of oligonucleotides. This is due, in part, to the highly sequence-

specific hybridization of DNA coupled with the extraordinary

electron-transport properties, catalytic properties, and high surface

area of various nanomaterials. DNA hybridization is detected

on nanomaterial-modified electrodes using either a direct label-

free detection scheme or indirect methods. Direct methods are

usually based on the redox signal of DNA bases, most notably

the oxidation of guanine which can be further amplified using

electrocatalytic mediators such as [Ru(bpy)3]2+, or by measuring

changes in the interfacial properties of the nanomaterial-modified

electrode including impedance and conductivity. Indirect methods

make use of electroactive indicators that either intercalate into

hybridized double-stranded DNA (ethidium bromide, daunomycin)

or employ labels such as metal nanoparticles which enable a variety

of electrochemical enhancements. In comparison to nonmodified

surfaces, these electrochemical assays exhibit orders of magnitude


Nanoparticle-Based Electrochemical DNA Detection 429

increased sensitivity by combining conventional electrodes with

nanoparticles, nanowires, dendrimers, liposomes, or carbon nan-

otubes. The use of these nanoscale materials for electrochemical

DNA sensors is discussed in the following sections.

14.2 Nanoparticle-Based Electrochemical DNA Detection

Nanoparticles can be synthesized in size ranges similar to many

common biomolecular markers. This trait makes nanoparticles

particularly well suited to interface with biomolecules and to make

hybrid systems. Typically, nanoparticles are prepared by chemical

methods such as decomposition of metal complexes or reduction

of metal ions. Capping agents are often used to stabilize the

nanoparticle, control the size distribution during growth, and also

provide functional groups to allow modification with a variety of

linking chemistries for tailor-made functionalities. The types of

metal nanoparticles typically used in sensing applications include

coinage and noble metal (gold, silver, iron, platinum, etc.) magnetic,

solid oxide, and semiconductor nanoparticles containing group II

or III elements (e.g., CdS, ZnS, InP). Metal–nanoparticle integration

into sensing schemes consists of their use as supports to immobilize

DNA probes onto surfaces and as electrochemical labels by detecting

their intrinsic atomic makeup (i.e., stripping voltammetry after

dissolution of the metal), or nonstripping methods that take

advantage of catalytic properties of the material.

14.2.1 Nanoparticle Modification of Electrodes and TheirUse as Supports for DNA Immobilization

Nanoparticles have been used extensively for the immobilization

of biomolecules [3]. In addition to their biocompatibility they can

produce a unique microenvironment that provides improvement

in the freedom of orientation for affinity binding with advantages

over planar substrates, an increase in surface area for higher probe

loading capacities, and enhanced diffusion of amplification agents.

Modification of electrode surfaces with nanoparticles can be carried

out by simple electrostatic adsorption or covalent attachments such



as chemical cross-linking, electron beam, or UV light irradiation, and

electro-deposition [4]. Electrostatic adsorption is straight forward

and the particle size can be strictly controlled from the previous

chemical synthesis of the nanoparticle. These surfaces, however,

are unstable and prone to particle desorption. Covalently cross-

linking nanoparticles to a surface can be quite versatile due to

the large range of functional groups available for cross-linking, but

first requires the modification of the surface which can hinder

electrochemical signals to the electrode. Nanoparticle synthesis

from electron beam and UV light irradiation does not suffer from

the insulating effects of covalent cross-linking; however, these

methods can be expensive and time consuming. Electrochemical

deposition of nanoparticles on the other hand, is a simple and

facile method to create nanoparticle-modified surfaces while the

final nanoparticle size and surface density can be controlled by

varying the deposition time, potential, and metal ion concentration

in solution. The following sections will focus on nanoparticle-DNA

immobilization methods in which electrochemistry was used to

modify surfaces with nanoparticles, or in which electrochemical

detection was combined with a DNA-nanoparticle–modified surface.

14.2.2 Gold Nanoparticle Supports

The chemisorption of thiol moieties onto gold makes the use of

gold nanoparticles a convenient support to immobilize sulfhydryl-

modified oligonucleotides for the construction of electrochemical

biosensors [5]. For instance, DNA hybridization was combined

with enzymatic electrochemical detection onto gold nanostructured

screen-printed carbon electrodes from the in situ generation of gold

nanoparticles using an applied constant current after which a 30-

mer oligonucleotide included in the SARS (severe acute respiratory

syndrome)-associated coronavirus genome was immobilized [6].

An alkaline phosphatase-modified detection probe was used to

monitor DNA hybridization events using a 3-indoxyl phosphate

substrate that produces a compound which was able to reduce

silver ions in solution into a metallic deposit. The deposited silver

was then electrochemically stripped into solution and measured by

anodic stripping voltammetry. Electrochemical deposition can also



be used to produce gold nanoparticles on planar gold electrodes

and was combined with a redox-active intercalating label to

create a DNA electrochemical biosensor [7]. The electrochemical

response of an immobilized long sequence single-stranded DNA

probe was monitored after target hybridization and measured by

cyclic voltammetry using methylene blue (MB) as an electroactive

indicator. It was shown that the immobilization of probe DNA onto

the nanogold aggregates (compared to the planar substrate) led

to a higher sensitivity and lower detection limit due to increasing

the number of probe molecules and improving molecular orien-

tation which increased the accessibility of target strands for DNA

hybridization.

Polyaniline is an attractive electropolymerizable polymer for

surface modifications due to its unique redox properties, high elec-

trical conductance, and ease of preparation. In addition, polyaniline-

modified surfaces retain a large specific surface area and can

remain conductive facilitating subsequent electron transfer. Feng

and coworkers [8] constructed a DNA impedance biosensor based

on gold nanoparticle/polyaniline nanotube membranes formed in

the presence of chitosan as shown in Fig. 14.1. Chitosan was used

Figure 14.1. Schematic diagram of the immobilization and hybridization

of DNA on Au/nanoPAN/GCE.



as a dispersant for aniline which was then electropolymerized onto

a bare glassy carbon surface to form polyaniline nanotubes. The

polyaniline nanotubes then served to nucleate and electrochemically

grow gold nanoparticles upon which single-stranded DNA oligonu-

cleotide probes could be immobilized. This technique combined the

large surface areas of two different nanomaterials, the polyaniline

nanotubes and the gold nanoparticles, to increase conductivity

and create a unique sensing composite membrane which was

characterized by cyclic voltammetry and electrochemical impedance

spectroscopy. DNA hybridization was monitored by impedance and

used to detect the sequence specific DNA of the phosphinothricin

acetyltransferase gene that exists in some transgenic crops. The

dynamic detection range was from 1 × 10−12 to 1 × 10−6 mol L−1,

the detection limit was 3.1 × 10−13 mol L−1, and the sensor showed

good selectivity, stability, and reproducibility.

14.2.3 Magnetic Particles

Magnetic (para- or super-) particles provide a means of both

immobilizing DNA and for separation and isolation from media

constituents in solution due to their ability to respond to an external

magnetic field [9]. Widely used as separation tools to purify many

biologically active compounds such as proteins, peptides, as well

as nucleic acids, they have also found use in electrochemical-

based DNA hybridization assays. The use of magnetic nanoparticle

probes has led to a “two-surface” strategy for improved biosensor

performance [10]. In traditional electrochemical DNA biosensors,

the probe recognition layer is directly immobilized onto the

electrode transducer with the hybridization and detection steps

being conducted on the same surface. The surface modification of

the transduction electrode, with immobilized single-stranded DNA

probes, can also act as an insulating layer and adversely affect the

electron-transfer kinetics for the detection method used. In contrast,

the two-surface approach allows for a separation of the hybridiza-

tion step, and after magnetic separation from nonhybridized DNA,

a fresh electrode can be used for detection. Additionally, the DNA-

bound magnetic particles which are suspended in the liquid phase



Figure 14.2. Some detection principles used in the double-surface DNA

hybridization techniques. (A) Label-free detection of target DNA (tDNA).

(B) Labeling of tDNA. Redox labels are covalently attached to the tDNA

strand outside the segment or on a secondary DNA strand recognized by the

capture probe. After hybridization and separation, the electroactive tags are

determined electrochemically (e.g., by ex situ adsorptive stripping voltam-

metry (a). Alternatively, electrochemical enzyme-linked immunoassay can

be used for detection of labeled tDNA at the MB surface (b).

allow for a higher degree of hybridization efficiency than DNA

probes immobilized on a flat substrate.

Figure 14.2 shows some general schemes where magnetic

particle-based DNA assays have been reported using a variety of

detection schemes utilizing two surface detection techniques. For

instance, a label-free approach has been developed where after DNA

hybridization and magnetic separation the target molecule can be

detected by cathodic stripping of nucleic acid bases (Fig. 14.2A) [11].

This approach can be applied directly; for instance, measuring gua-

nine oxidation with inosine-substituted DNA probes to lower back-

ground signals from guanines contained in the probe strand [12], or

by releasing purine bases by acid treatment for sub-nanomolar DNA

detection at silver, copper, platinum, or gold amalgam electrodes

[13–15]. The accumulation of guanine and adenine anodic signals

at carbon electrodes through a Cu(I)-purine complex can also be

used for an amplification effect. Alternatively, the labeling of tDNA,

or the corresponding secondary reporter probe in a “sandwich”

hybridization assay can be performed on magnetic particles as

shown in Fig. 14.2B. Redox labels, such as covalently bound osmium



tetroxide complexes, can be incorporated into DNA strands outside

the recognition sites or onto a secondary capture probe to be

measured electrochemically for determination of the amount of DNA

hybridization (Fig. 14.2B(a)). Enzyme reporter tags have been found

useful in detection strategies due to the catalytic signal amplification

from substrate turnover, and have also been used as reporter labels

in magnetic-based DNA assays (Fig. 14.2B(b)). These have been

reported either as an enzyme-linked immunoassay or by directly

linking the enzyme label to a secondary DNA probe.

14.2.4 Layer-by-Layer Immobilization Techniques

The sequential charge inversion of alternating polycation/polyanion

solutions to form multilayers, known as layer-by-layer assembly, is

a simple and efficient technique to form biologically active surfaces.

Several studies have used the layer-by-layer technique to immobilize

DNA functionalized multiwalled carbon nanotubes (MWCNT) with

nanoparticles that result in effective electrochemical DNA sensors.

In one report covalent attachment of Au nanoparticles and MWCNTs

was accomplished by first successively carboxylating the nanotubes

followed by cross-linking aminothiol groups to introduce thiol

functionalities [16]. Thiolated nanotubes were then adsorbed onto

a gold electrode followed by adsorption of gold nanoparticles. This

process was repeated 6 times until a final layer of gold nanoparticles

was used to adsorb probe DNA. In another configuration, cysteamine

was first attached to the gold electrode and acted as a molecular glue

to covalently attach carbodiimide ester-activated COOH-MWCNT

followed by treatment with a cysteamine/AuNP solution [17].

The additional cysteamine would subsequently conjugate to the

activated MWCNT while its free sulfur group would attach to

the gold nanoparticles. This process was then repeated to create

an alternating MWCNT/gold nanoparticle film with a controlled

number of bilayers. A final layer of cysteamine/silver nanoparticles

and activated MWCNT was used to covalently attach NH2-DNA

probes to create a reproducible and stable biosensor. In both

these works, detection of the DNA was carried out by monitoring

the voltammetric detection of the DNA intercalator doxorubicin

following the hybridization event.



Figure 14.3. Schematic representation of the immobilization, hybridiza-

tion, and detection of probe DNA.

In another configuration silver nanoparticles were electrode-

posited onto the surface of a previously electro-polymerized

poly(trans-3(-pyridyl) acrylic acid)-multiwalled carbon nanotube,

glassy carbon electrode while DNA hybridization events were mon-

itored by differential pulse voltammetry (DPV) after intercalation

of adriamycin and chemisorption of thiolated single-stranded DNA

onto the silver nanoparticles, as shown in Fig. 14.3 [18]. Multiple

DNA assays were performed by de-hybridizing DNA duplexes with a

1:1 H2O:HNO3 solution for 15 min to regenerate the single-stranded

DNA surface. Both detection schemes showed high sensitivity,

selectivity, and reusability and took advantage of the synergistic

effects of combining carbon nanotubes to increase conductivity

and metal nanoparticles to provide a suitable platform for DNA

immobilization.

14.2.5 Metal Nanoparticle Labels for DNA HybridizationDetection

14.2.5.1 Direct detection of the nanoparticle label

Costa-Garcia and coworkers [19] first reported on using gold

nanoparticles to electrochemically monitor an affinity binding



event by using an adsorbed biotinylated albumin layer to capture

streptavidin–gold conjugates on a pre-treated carbon paste elec-

trode. The colloidal gold label was detected following its oxidation at

a high potential in an acidic medium, and then reducing the released

AuCl–4 complex using differential pulse voltammetry. A modified

version of this detection principle was applied by the groups of

Limoges and Wang to detect DNA hybridization events based on

the oxidative dissolution of the particle in acidic bromine–bromide

solution and using highly sensitive stripping voltammetry [20,

21]. In the former case the 406-base pair human cytomegalovirus

DNA sequence was detected using oligonucleotide-modified gold

nanoparticle probes at probe-modified screen-printed microband

electrodes and had a detection limit of 5 pM. In the latter case

a two-surface technique was used to detect a DNA sequence

related to the BRCA1 breast cancer gene where magnetic bead

probe DNA complexes were used to hybridize to biotinylated DNA

that was conjugated to commercially available 10 nm streptavidin

gold nanoparticles. Following magnetic separation and nanoparticle

dissolution, the oxidized gold ions were used to determine the

amount of hybridized target at a thick-film screen-printed carbon

electrode using potentiometric stripping analysis. In both cases

a significant amplification signal can be attributed to metal

accumulation in the pre-concentration step of the stripping analysis

which makes the technique sensitive to the detection of trace

metals and particularly well suited for metal nanoparticle detection.

Further amplification can be performed after the hybridization

event by catalytically precipitating metals, such as gold and silver,

onto the nanoparticle label “seed”. Thus, more metal can be grown

in solution to increase the sensitivity of DNA hybridization binding

events [22]. Attomolar detection limits were achieved using a triple-

amplification strategy [23]. Instead of single nanoparticles being

used for each hybridization event, streptavidin-coated polystyrene

microspheres, each containing multiple biotinylated gold nanopar-

ticles and biotinylated DNA secondary capture probes were used.

Gold precipitation, acidic dissolution, and detection after DNA

hybridization resulted in a significant lowering of detection limits.

Wang et al. [24] also reported a solid-state detection method where,

after a silver-enhanced precipitation step, the enlarged gold–silver



magnetic bead–DNA conjugate was collected by positioning a

magnet behind a screen-printed carbon electrode. The attraction of

the conjugate directly onto the working area of the electrode allowed

the stripping detection step to take place without dissolution of the

metal obviating the need for the caustic acidic medium.

Inorganic semiconductor nanocrystals have also found their

use as electrochemical labels for DNA detection. Cadmium sulfide,

for instance, was reported to be a viable alternative for gold

nanoparticles. After dissolution in nitric acid Cd+2 ions can be

detected at a mercury or bismuth film electrode [25]. Taking

advantage of the wide potential window and the fact that multiple

group II and III metals can be detected simultaneously at mercury

and bismuth film electrodes, a multitarget DNA hybridization

assay was developed using three different inorganic nanocrystals

(ZnS, CdS, and PbS) to simultaneously detect three different DNA

targets in the same solution [26]. A general scheme of Wang’s

nanoparticle magnetic bead-based protocol for electrochemical DNA

detection consisting of gold nanoparticles (A), silver enhancement

(B), magnetic collection and solid state detection (C), the use of CdS

(D), and multiple inorganic semiconducting encoding nanoparticles

(E) is presented in Fig. 14.4.

Merkoci and coworkers [27] have reported several works

describing DNA electrochemical biosensors based on the direct

determination of gold nanoparticles which have been adsorbed

onto the rough surface of graphite–epoxy composite electrodes,

their electrochemical oxidation at +1.25 V, and the detection of the

resulting tetrachloroaurate ions by differential pulse voltammetry.

The use of 1.4 nm Au67 particles allowed the 1:1 conjugation of

nanoparticle to magnetic bead-DNA probe and prevented cross-

linking effects resulting in lower detection limits over previous

assays [28]. A magnet placed into the graphite–epoxy electrode

transducer collected the hybridized DNA after magnetic separation

and allowed for the direct detection of the gold nanoparticle label.

Two other gold nanoparticle assays were described based on this

method using larger gold nanoparticles conjugated to DNA using

biotin/streptavidin interactions with the first being a two-strand

detection technique to detect the BRCA1 breast cancer gene, and

the second a sandwich assay to detect a DNA sequence related to



Figure 14.4. Particle-based protocols for electrochemical detection of

DNA. These assays involve the introduction of the probe-attached onto the

magnetic particles, addition of the target/hybridization event, magnetic

removal of unwanted materials, binding of the metal, and amplified

electrochemical detection of the dissolved gold (Au) (A), silver (Ag) (B),

and cadmium sulfide (CdS) (D) nanoparticles. Me: metal tag. Also shown are

solid-state stripping (C) and multitarget (E) detection protocols.

the cystic fibrosis gene that could detect single- and three-base

mismatches [29]. The modification of gold nanoparticles with single

DNA bases was used to detect single nucleotide polymorphisms

(SNP), as described by Kerman et al. [30]. Phosphoramidite

chemistry was used to attach the monobases onto chitosan-modified

gold nanoparticles, which could then accumulate into a mismatched

DNA base pairing through Watson-Crick hydrogen base pairing in

the presence of DNA polymerase I. The electrochemical oxidation

signal of the gold nanoparticles could then be used to determine

the presence of mismatch sites in a synthetic 21-base DNA probe

related to tumor necrosis factor along with all its possible mutant

combinations. Liu et al. [31] subsequently reported a bioelectronic

method for coding SNPs using different encoding nanocrystals.



Adenosine, cytosine, guanosine, and thymidine mononucleotides

were linked to ZnS, CdS, PbS, and CuSnanoparticles, respectively,

and sequentially introduced to a DNA hybrid-coated magnetic

bead solution. Characteristic multipotential voltammetric peaks

were produced depending on the base pairing of the different

nanocrystal-mononucleotide conjugates with each mutation capa-

ble of identifying each of eight possible one-base mutations in a

single run.

Ying has described two solid-state approaches based on the

incorporation of silver nanoparticles into DNA duplexes followed

by the direct detection of the nanoparticles based on an Ag/AgCl

cycling process shown in Fig. 14.5. In the first approach neutral

PNA, which can significantly increase DNA hybridization efficiency

due to a lack of electrostatic repulsion of the DNA target, was

used as the probe capture molecule [32]. Following target DNA

hybridization the surface would become negatively charged and

could then be labeled with positively charged dodecylamine-capped

Ag nanoparticles (Fig. 14.5A). In the second approach a normal

thiolated DNA mixed mono recognition layer was used in connection

Figure 14.5. Schematic for biosensing strategy using (A) dodecylamine-

capped Ag nanoparticles and (B) doxorubicin-modified Ag nanoparticles.



with doxorubicin-modified silver nanoparticles [33]. As doxorubicin

is a well-known DNA intercalator, the particles could then intercalate

into the DNA duplex after hybridization (Fig. 14.5B).

14.2.5.2 Non-stripping-based nanoparticle electrochemicalDNA detection methods

Ruthenium hexamine (RuHex) is a positively charged electroactive

complex that can bind to the anionic phosphate backbone of

DNA strands. Zhang et al. [34] constructed an electrochemical

DNA biosensor by creating a mixed monolayer of DNA probes

onto a gold surface, shown in Fig. 14.6. A sandwich assay was

used to bind DNA-coated gold nanoparticles and bring them in

proximity to the electrode surface. The RuHex marker could then

be bound to DNA strands through electrostatic interactions and

its signal measured as a direct function of DNA hybridization

(Fig. 14.6A). The resulting sensor produced fM detection limits

A: DNA-AuNPs technology

B: Modified bio bar codes technology

C: This method

bridge DNA

Figure 14.6. Schematic diagram for the DNA biosensor fabrication based

on a one-to-one recognition tri-gold nanoparticle DNA probe. And the

comparison of DNA biosensor fabrication based on Au NPs modified with

only one kind of DNA (A: DNA–Au NPs), Au NPs modified with two kinds of

DNA (B: modified bio-bar code technology), and a one-to-one recognition

tri-gold nanoparticle DNA probe technology (C).



with excellent differentiation for single base mismatches. The use

of gold nanoparticles provided a significant signal amplification

effect in that hundreds of DNA reporter strands were immobilized

on each particle thus increasing the amount of reporter RuHex

molecules that could bind. A modified “bio-bar code” technique that

mixes both complementary and noncomplementary DNA probes

on the modified gold nanoparticles limits the number of strands

available for hybridization of target molecules on the surface

(Fig. 14.6B) [35]. This subsequently decreases the number of DNA

interconnects on the transducer surface and has a profound impact

on the reproducibility and sensitivity of the technique. A DNA probe

bridge could be constructed that could combine two different gold

nanoparticle bio-bar codes. The DNA bridge gold nanoparticle bio-

bar code conjugate contained three gold nanoparticle labels and

only one linking DNA molecule for target binding (Fig. 14.6C). The

resulting tri-gold nanoparticle DNA probe combined the maximum

synergy of signal amplification, from the electrostatic binding of

ruthenium hexamine onto 486 DNA reporter probes on the three

gold nanoparticles, and increased selectivity from the one-to-one

recognition of the single target binding site to achieve a detection

limit of 53 aM. Li et al. [36] reported another version of this tech-

nique where an avidin/polyamidoamine (PAMAM) dendrimer/3-

mercaptopropionic acid layer was used to immobilize DNA probes.

The use of the PAMAM served as an additional amplification effect,

along with the use of gold nanoparticles to bind RuHex, due to

the increased amount of DNA probes that could be attached when

compared to a flat substrate and led to a low detection limit of

1.4 × 10−14 mol L−1.

Enzymes have found wide use as labels in biological assays due

to their ability to produce catalytic signals from the generation of

electroactive products. However, there are some inherent drawbacks

with using biological labels associated with their thermal and

environmental instabilities. The large surface area-to-volume ratio

of nanoparticles makes them superior catalysts when compared to

their bulk metal counterparts. Taking advantage of these catalytic

properties, Willner and coworkers [37] introduced the use of metal

nanoparticles as inorganic analogues to traditional enzyme tags by

using single-stranded DNA probe-modified platinum nanoparticles



as electroactive labels. Amperometric currents were generated

from the Pt-catalyzed reduction of H2O2 following DNA target

capture onto a DNA-probe mixed monolayer gold electrode and

secondary DNA-Pt nanoparticle hybridization with a detection limit

of 10 pM for target DNA. The substitution of an enzyme with an

inorganic nanoparticle combines the advantages of high sensitivity

from substrate turnover and increased stability for the amplified

detection of biomolecules. Yang and coworkers [38] described the

detection of DNA hybridization onto an ITO electrode using DNA-

conjugated gold nanoparticles to catalytically oxidize hydrazine.

Because of the high overpotential and slow electron transfer kinetics

of hydrazine oxidation, a NaBH4 treatment was used to enhance

the catalytic signals to produce a detection limit of 1 fM. The

pre-treatment hydrolyzed NaBH4 and induced sorption of atomic

hydrogen onto the gold nanoparticles. This process, however,

occurred at very slow rates at higher pH. The substitution of

gold nanoparticles with Pd nanoparticles increased the catalytic

hydrolysis time, even at high pH, and allowed the construction of

a DNA hybridization detector using the Pd catalyzed oxidation of

NaBH4, shown in Fig. 14.7. ITO electrodes were modified using

silanization with a copolymer containing carboxylic acid groups (to

Figure 14.7. Schematic view of DNA detection using the catalytic and

electrocatalytic oxidation of NaBH4 on Pd NPs and the rapid enhancement

of electrocatalytic activity of DNA-conjugated Pd NPs.


Nanowires, Nanorods, and Nanofibers 443

conjugate probe amine-terminated DNA) and poly (ethylene glycol)

units (to limit nonspecific adsorption). After subsequent target

binding followed by DNA–nanoparticle capture, the Pd nanoparticles

catalyzed the hydrolysis of NaBH4 and the sorption of many atomic

hydrogens which were used to generate catalytic currents with a

detection limit of 10 aM.

14.3 Nanowires, Nanorods, and Nanofibers

The use of nanowires, nonorods, nanofibers, etc. has also attracted

considerable attention for use in detection of DNA and other bio-

molecules [39]. Similar to carbon nanotubes, these one-dimensional

nanostructures posses unique electrical properties due to their high

surface-to-volume ratio and extreme sensitivity of carrier charge

mobility that can be exploited for sensing [1a]. Additionally, the

dimensional scale of these materials is comparable to that of the

biological species being interrogated, providing interesting oppor-

tunities for use as labels or signal transducers for electrochemical

sensing. The extremely small footprint of these nanomaterials may

allow assembly of numerous sensors onto a small area, facilitating

development of devices capable of detecting a host of analytes.

Synthesis and characterization of nanowires remains a signif-

icant focus area of nanotechnology [40]. Nanowires composed

of metals, semiconductors, conducting polymers, diamond, and

other materials have been reported. Although several methods

exist for producing nanowires, the use of porous templates for

the synthesis of nanowire tubes and -rods is the most commonly

used and has been extensively investigated. In this approach an

inert porous membrane, anodized alumina, for example, is used

as the template for forming well-defined free-standing nanowires

that can be oriented or non-oriented. The nanowires are formed

by electrochemical or electrophoretic deposition of the desired

material(s) into the porous template which can be subsequently

removed or left as a scaffold for the nanowire array. Other methods

for producing nanowires include evaporation/condensation, dis-

solution/condensation, vapor/liquid/solid (vapor deposition), and

substrate ledge or step induced growth.



14.3.1 Nanorods as Labels

Use of nanorods as labels for electrochemical detection of DNA was

first reported by Wang and coworkers [41]. In this work conical

indium/gold nanorods, approximately 3 to 5 μm in length, were

synthesized via sequential electrodeposition of Au and indium into

alumina membranes. Following synthesis, the alumina template

was dissolved in 3 M NaOH yielding free nanorods (Fig. 14.8B).

These rods were then modified with thiolated oligonucleotide

detection probes, complementary to a portion of the target DNA, via

Figure 14.8. (A) Schematic representation showing sandwich hybridiza-

tion linking magnetic beads and indium/gold nanorods through the DNA

target, magnetic collection of the DNA-linked particle assembly onto the

thick-film electrode transducer, and solid-state derivative chronopotentio-

metric measurements of the captured indium rods. P1, DNA probe 1; T, DNA

target; P2, DNA probe 2; MR, indium/gold nanorods; MB, magnetic beads;

and M, external magnet. SEM images of (B) indium/gold rods and (C) DNA-

linked particle assembly (after sandwich hybridization assay).



thiol–gold interactions. In the presence of target DNA, a sandwich

was formed between magnetic beads modified with capture

probe DNA, the target DNA, and the detection probe modified

indium/gold nanorods. This is shown schematically in Fig. 14.8A.

Detection occurred by either solid-state chronopotentiometry of the

indium/gold nanorods collected at a mercury-coated screen-printed

carbon electrode by an external magnet, or chronopotentiometric

stripping of the indium label, following dissolution under acidic

conditions, at a mercury-coated carbon-fiber electrode. The use of

the nanorods as labels allowed detection of 30 ng/L (250 zmol)

target DNA.

14.3.2 Nanowires Interfaced with Electrodes as anImmobilization Matrix

Nanowires have also been used as an immobilization matrix for

probe DNA with inherent enhanced electron-transfer kinetics and

higher surface area. Electroactive reporter molecules are then used

to measure immobilized DNA. Kelley’s group has reported a platform

for electrochemical DNA detection using arrayed gold nanowires

generated by electroless deposition of gold onto polycarbonate

membranes (see Fig. 14.9A), which were exposed by subsequent

oxygen plasma etching (Fig. 14.9B) [42]. Thiolated probe DNA

was deposited onto the gold nanowire array and [Ru(NH3)6]3+

and [Fe(CN)6]3− were used as electrocatalytic reporters for the

amount of hybridized target DNA, as shown in Fig. 14.9C, yielding

an attomole-level detection limit. The authors showed that catalytic

currents and diffusional mobility of Ru3+ ions at the nanowire array

are markedly different than that obtained at bulk macroelectrodes

allowing for improved signal-to-noise ratio and sensitivity [43].

These results demonstrate the utility three-dimensional nanoscale

systems can posses over bulk macroscale systems. Very recently,

Kelley’s group reported an extension of their work in which peptide

nucleic acid (PNA) probes were immobilized onto gold nanowire-

modified polycarbonate membranes and again used [Ru(NH3)6]3+

and [Fe(CN)6]3−as electrocatalytic reporters [44]. Unlike DNA, the

peptide backbone of the PNA probe resulted in a neutral charged

surface that provided significantly decreased background signals. In



Figure 14.9. Scanning electron micrographs and schematic illustrations

of 2D (A) and 3D (B) gold nanowire electrodes. (C) Modification of the gold

nanowire electrodes with thiolated probe DNA, subsequent hybridization of

target DNA, and detection via electrocatalysis of Ru(III)/Fe(III).

this work femtomolar levels of DNA, as well as an RNA sequence

relevant to prostate cancer, were detected in unamplified patient

samples.

Gold nanowire arrays were also used by Andreu and coworkers

for DNA detection [45]. Anodic aluminum oxide membranes were

used as templates for galvanostatic Au electrodeposition followed

by treatment in base to dissolve away the template leaving free-

standing gold nanowires 330 nm in diameter and ∼2 μm in

length. [Ru(NH3)6]3+ was used to measure charge before and after



hybridization of the target DNA via chronocoulometry using the

method of Tarlov [46]. The authors reported that the large surface

area of the electrode resulted in large measured currents (hundreds

of μA to nearly a mA) and large IR drops requiring the use of

resistance compensation to perform effective DNA quantification

measurements.

The use of conducting polyaniline nanowire–modified electrodes

for electrochemical DNA detection has also been reported. Zhu et al.[47] directly deposited polyaniline nanowires onto a glassy carbon

electrode from an aniline containing electrolyte solution yielding

nanowires with diameters ranging from 80 to 100 nm. Probe DNA

with a free carboxyl group was covalently linked to free primary

amines on the polyaniline nanowires via carbodiimide chemistry.

Hybridization of target DNA was monitored using differential pulse

voltammetry and methylene blue (MB) as the electroactive reporter.

MB binds to guanine bases of ssDNA with higher affinity than dsDNA

in which the guanine residues are less accessible. This resulted in

a decrease of current, or a “signal off” detection mechanism, with

a detection limit of 1 pM. In a similar work, Chang and coworkers

[48] electrochemically deposited ordered polyaniline nanowires

onto a graphite electrode using a porous aluminum layer template.

The porous aluminum template was prepared by deposition of

aluminum onto the electrode via magnetron sputtering followed by

anodization. Carbodiimide was also used to link carboxyl-modified

DNA probes to the nanowires (40 nm diameter). In this work,

daunorubicin, which binds with higher affinity to dsDNA, served

as the electroactive reporter. This “signal on” approach yielded a

significantly improved detection limit of 1 fM which the authors

attribute to enhanced conductivity and faster hybridization kinetics

at the oriented nanowires.

The first use of vertically aligned conducting diamond nanowires

for electrochemical DNA detection was also recently reported

[49]. Boron-doped diamond posseses many advantages over other

materials used for producing nanowires including high chemical

stability, low background current, wide potential window, and high

biocompatibility. In this work metal-like diamond nanowires were

fabricated from boron-doped single crystalline diamond produced

by chemical vapor deposition and subsequently exposed to reactive



Figure 14.10. (A) SEM image of vertically aligned conducting diamond

nanowires. Examples of detection of DNA hybridization by (B) cyclic voltam-

metry, and (C) differential pulse voltammetry. Target DNA concentration

was 10 nM.

ion etching using diamond nanoparticles as hard etch masks to form

nanowires, as shown in Fig. 14.10A. The tips of the wires were

functionalized with aminophenyl groups by electrodeposition of

nitrophenyl diazonium followed by electroreduction of nitro groups

to amines. A heterobifunctional crosslinker was used to covalently

link the free amine groups on the diamond nanowires to thiol-

modified DNA probes. [Fe(CN)6]3− was used as a redox probe in

which peak currents would decrease upon hybridization of target

DNA yielding an ∼2 pM detection limit (see Fig. 14.10B, C). This

conducting diamond nanowire sensor proved 100 to 1000 times

more sensitive than sensors composed of smooth gold or diamond

surfaces.

14.3.3 Nanowire Conductance Based DNA Detection

Nanowires have been used to bridge two closely spaced electrodes

for DNA detection by monitoring the conductance of the nanowire

during hybridization. Binding of the negatively charged DNA strand

to the nanowire increases the net negative surface charge density

leading to an increase in conductance between the two electrodes.

This method is analogous to field-effect transistor (FET) switches

used in microelectronics in which the electrodes serve as the

electron source and drain while the nanowire serves as the

modulating gate [39]. In addition to being label-free and reagentless,



this approach also allows for real-time detection of target DNA [1a].

As these devices are very sensitive to changes in conductance, they

suffer from high sensitivity to the sample solution ionic properties

and impurities found in complex detection matrices.

Silicon nanowires were employed by both Hahm and Lieber

[50] and Li et al. [51] for the label-free real-time detection of

ssDNA sequences via conductometric monitoring of target DNA

hybridization. In the work of Li, ssDNA probes with acrylic

phosphoramidite functionality were immobilized to a silicon

nanowire which had been previously exposed to the vapor

of 3-mercaptopropyltrimethoxysilane. By monitoring changes in

conductance target DNA 12-mer strands could be detected at

concentrations as low as 25 pM, and the sensor showed excellent

discrimination against single-base mismatch sequences. Hahm and

Lieber employed biotinylated PNA probes conjugated to a silicon

nanowire, previously modified with biotin followed by avidin, to

detect 31-mer DNA strands. PNA probes were chosen over DNA

probes due to their higher affinity for DNA, greater stability,

and neutral charge. A detection limit of 10 fM was reported for

this system along with good discrimination against single-base

mismatch sequences, and similar changes in conductance from

device to device. A top-down approach was also recently reported

for producing an array of highly ordered silicon nanowires for

DNA detection [52]. This method resulted in high uniformity and

reproducibility and allows for simpler scaling and manufacturing of

the sensor. Similar to the work of Hahm and Leiber, this sensor was

modified with PNA probes and yielded a 10 fM detection limit.

Multisegment CdTe-Au-CdTe nanowires have also been used

for FET-based sensing of DNA [53]. Synthesized by consecu-

tive electrodeposition onto an anodized alumina template these

metal-semiconductor nanowires exhibit a p-type behavior. Thiol-

terminated ssDNA probes were bound via Au–thiol interaction to the

Au segment of the nanowires. Target DNA could be detected at 1 μM

and higher concentrations.

In a recent report, Kong and coworkers [54] developed a

conductometric DNA sensor in which captured target DNA serves

as the template for electroless silver deposition forming silver

nanowires. In this work, interdigitated electrodes were formed onto



Figure 14.11. Schematic representation of the DNA-pectin templated

silver nanowire formation between two interdigitated electrodes.

a silicon substrate with 500 nm gaps between the electrodes. The

silicon substrate between the interdigitated electrodes was modified

with 3-aminopropyl triethoxysilane, as shown in Fig. 14.11, allowing

cross-linking to amine terminated PNA probes. Upon binding of

target DNA, zirconium-phosphate-carboxylate chemistry was used

to bind the polysaccharide and pectin to the DNA. Oxidation of

the pectin under acidic conditions yielded aldehyde groups which

served as sites for silver deposition via Tollen’s reduction. The

formation of the DNA templated silver nanowires significantly

reduced the resistance measured between the interdigitated elec-

trodes allowing detection of DNA as low as 3 fM. Use of DNA as

a template for growth of silver nanoclusters was also reported by

Wang’s lab [55]. In this work, ssDNA probes were immobilized

via carbodiimide chemistry to cystamine modified Au electrodes.

Following binding of target DNA, silver ions were loaded onto the

DNA by Na+/Ag+ exchange/electrostatic interactions under basic

conditions. Hydroquinone was then used to catalyze silver reduction

forming DNA templated silver nanoclusters. These aggregates were



dissolved in nitric acid and the silver ion solution was transferred

to a screen-printed carbon electrode for potentiometric stripping

analysis yielding a highly linear response (current peak area) with

DNA concentration and a detection limit ∼100 ng/mL.

14.3.4 Electrochemical Impedance Spectroscopy atNanowires for DNA Detection

Electrochemical Impedance Spectroscopy (EIS) is a method used to

characterize electron-transfer reactions by perturbing the system in

a sinusoidal manner over a wide range of frequencies. This method,

which is very sensitive to the properties of the electrode interface,

provides information regarding electron-transfer kinetics, diffusion

of charged species, charging/discharging, and system conductance.

Very recently Chen and coworkers [56] demonstrated the use

of EIS for label-free electrochemical detection of DNA sequences

relevant to anthrax lethal factor on gallium nitride (GaN) nanowires.

The GaN nanowires were grown on a silicon substrate coated with

Au catalyst using Ga as the source material and NH3 as the reactant

gas in a tubular furnace via air pressure chemical vapor deposition.

EIS measurements of the “as grown” GaN nanowires, observed in

the Nyquist plot in Fig. 14.12A, exhibited a semicircle and a straight

vertical line, indicative of finite impedance at the GaN/electrolyte

Figure 14.12. Electrochemical impedance spectroscopy based in situDNA sensing: (A) Nyquist plots and (B) corresponding Bode plots of as

grown, DNA probe against anthrax lethal factor (pLF)-modified, and dsDNA-

modified GaNNWs at different concentrations of LF targets (arrows indicate

increasing concentration, in situ DNA hybridization detection).



interface and suppressed diffusion-limited electrochemical behav-

ior. Interestingly, upon binding of thiolated probe DNA to 3-

mercaptopropyl trimethoxysilane-modified GaN nanowires, two

semicircles were observed. The first semicircle was now indicative

of charge transfer at the GaN–DNA interface, with the second

semicircle indicative of charge transfer at the DNA–electrolyte

interface. These two phenomena are more clearly observed in the

two peaks in the Bode plot shown in Fig. 14.12B. The deconvolution

of the charge transfer properties of these two interfaces allowed for

monitoring the extent of DNA hybridization (decrease in resistance

to charge transfer in the semicircle corresponding to the GaN/DNA

interface, shown in Fig. 14.12A) while the second interface served

as a fingerprint for modification of the nanowires with DNA.

Picomolar concentrations of target DNA, even in the presence of

noncomplementary and mismatched sequences, were reported.

14.3.5 Dendrimers

Dendritic polymers, or dendrimers, are three-dimensional nano-

sized synthetic molecules possessing a regularly branched tree-

like structure. Dendrimers can be described as covalent micelles

having well-defined cavities, being nontoxic/biocompatible, and can

contain several functional groups allowing for functionalization

and/or immobilization of the dendrimers [57]. Several schemes

utilizing dendrimers for electrochemical detection of DNA have

been reported. In these reports dendrimers are either loaded

with electroactive reporter molecules and used as labels for DNA

detection, or immobilized on electrodes as scaffolds for DNA

immobilization providing higher probe densities and improved

electron transfer to the electrode.

Commercially available poly(amidoamine) (PAMAM)

dendrimers are the most commonly used dendrimers for electro-

chemical DNA detection. Recently, Zhu and coworkers [58] reported

the use of a new class of PAMAM dendrimers with a trimesyl core

and terminal carboxyl groups for DNA detection. Amine-modified

target DNA was immobilized to the dendrimer using carbodiimide

chemistry. Amine functionalized probe DNA was also immobilized

to a mercaptoacetic acid self-assembling monolayer (SAM)-modified



Au electrode via carbodiimide chemistry. EIS in the presence of

[Fe(CN)6]3− [also referred to as faradic impedance spectroscopy

(FIS)] was employed to detect dendrimer-labeled target DNA

hybridization. Upon hybridization, the negatively charged den-

drimer on the electrode surface induced electrostatic repulsion of

the negatively charged [Fe(CN)6]3− reporter. Monitoring resistance

to charge transfer by FIS resulted in a detection limit of 2.5 pM. This

sensitivity was two orders of magnitude lower than that obtained

for target DNA without the PAMAM dendrimer label. Similar results

were obtained by Humenik et al. [59] utilizing detection probe

DNA conjugated with PAMAM dendrimers that had been loaded

with esterase enzymes. These polyvalent esterase dendrimer DNA

clusters were hybridized to captured target DNA immobilized on

an Au electrode in a sandwich assay format. The amperometric

signal of p-aminophenol produced by the esterase enzymes was

used indirectly to detect DNA and resulted in a detection limit of

20 fM. This provided a 100-fold signal enhancement over use of

monovalent esterase-detection DNA probe conjugates.

Immobilization of PAMAM dendrimers on Au electrodes previ-

ously modified with SAMs has also been reported. Zhu et al. [60]

utilized a carboxyl terminated SAM to crosslink amine-terminated

PAMAM dendrimers via carbodiimide chemistry to an Au electrode.

This was followed by immobilization of phosphate-modified probe

DNA by phosphoramidate bond formation. Daunorubicin was used

as an electroactive indicator of target DNA hybridization. In a similar

work, Li and coworkers [61] used glutaraldehyde to immobilize

amine-terminated PAMAM to an Au electrode modified with an

amine-terminated SAM, followed by conjugation of the dendrimer

with amine-modified probe DNA, again with glutaraldehyde. FIS

with [Fe(CN)6]3− reporter was used to monitor changes in surface

charge and electron-transfer properties upon binding of target DNA.

The detection limit for both systems was similar: 8 pM using DPV

and daunorubicin, and 3.8 pM using FIS and [Fe(CN)6]3−. Both

reports demonstrated higher sensitivity for DNA when dendrimers

were used as compared to SAM-modified electrodes alone. This was

attributed to the higher surface area of dendrimer-modified sur-

faces significantly improving the immobilization capacity of probe

DNA.



Figure 14.13. (A) Formation of mixed SAM on Au electrode, (B) immobi-

lization of ferrocene functionalized dendrimers (Fc-D), (C) immobilization

of thiolated capture probe with bifunctional linker, (D) hybridization with

target, (E) hybridization with biotinylated detection probe, (F) association

with avidin-alkaline phosphatase, (G) description of the process of the

electrocatalytic reaction of p-aminophenol ( p-AP) via electronic mediation

of Fc-D.

Incorporation of electroactive ferrocene groups into PAMAM

dendrimers for enhanced electrochemical signal has also been

reported [62]. In this work ferrocene functionalized dendrimers

were immobilized onto a SAM-modified Au electrode, as shown in

Fig. 14.13, and served as an immobilization matrix for the capture

probe DNA, and as an electrocatalyst for p-aminophenol oxidation.

p-aminophenol was produced by alkaline phosphatase labeled

detection probe DNA used in a sandwich-type enzyme-linked DNA

assay. The authors show that use of the ferrocene functionalized

dendrimers lead to a significant enhancement in electrochemical

signal resulting in a 100 pM detection limit.

Gibbs et al. [63] reported the formation of dendrimers from

norbornene block copolymers with detection probe DNA and



ferrocenyl side chains for use in a sandwich type assay. Interestingly,

the DNA-diblock copolymer dendrimers used for detection showed

higher binding affinity and sharper melting profiles than the

ssDNA used to form the dendrimer. Two ferrecenyl derivatives,

ferrocenyl and dibromoferrocenyl, were used to form the dendrimer

allowing one to tailor the redox characteristics of each DNA-

diblock copolymer dendrimer probe. Using this strategy detection

of multiple targets simultaneously and detection of point mutations

was possible. Target DNA could be detected at 100 pM and higher

concentrations.

Nanoscale dendrimers of DNA have also been utilized for

electrochemical DNA detection. Wei and coworkers [64] employed

a polymer-DNA dendrimer surface for DNA and RNA detection.

Streptavidin functionalized DNA dendrimers were incorporated into

polypyrrole on an Au electrode via electropolymerization. This

was followed by immobilization of biotin-terminated capture probe

DNA. The capture probe was designed to form a hairpin loop in

the absence of target DNA, and contained a FITC label on the

end opposite of the biotin group. In the presence of target DNA

or RNA, the hairpin loop opened exposing the FITC group. An

anti-FITC antibody-horseradish peroxidase conjugate bound to the

exposed FITC group. In the presence of substrate and mediator,

the horseradish peroxidase produced an electroactive signal. The

authors report a detection limit of 10 aM which they attribute to the

conducting polymer-DNA dendrimer interface providing enhanced

electron-transfer kinetics and high probe density.

14.3.6 Apoferritin Nanovehicles

Apoferritin is a spherical protein shell composed of 24 protein sub-

units, forming an outer diameter of 12.5 nm and an aqueous interior

about 8 nm in diameter [65]. This protein cage is capable of holding

about 4500 iron atoms and can be reversibly dissociated into its 24

subunits at low pH (2.0), and reassembled at high pH (8.5). Modula-

tion of pH can thus serve as a method to load and release electroac-

tive markers allowing apoferritin to be employed as an electroactive

label. Such an approach avoids the use of harsher acid dissolution

of quantum dot NP labels and complicated semiconductor NP



synthesis [66]. Loading of apoferritin with Zn, Cd, and Pb phosphate

NPs followed by release and electrochemical stripping analysis has

been reported [67]. The use of the different metal phosphate NPs

allowed for simultaneous detection of the different NPs at different

potentials, or identification of compositionally encoded nanoparti-

cles which may prove efficacious for multianalyte detection.

Electrochemical detection of DNA using apoferritin as a nanove-

hicle label was reported by Liu and coworkers [68]. Dissoci-

ated apoferritin subunits were reassembled in the presence of

[Fe(CN)6]3− producing electroactive apoferritin, each loaded with

∼150 [Fe(CN)6]3− molecules. Free carboxyl groups on the exterior

of the apoferritin were coupled to amine-terminated DNA probes

via carbodiimide chemistry. This DNA-apoferritin conjugate served

as the detection probe in a magnetic bead based sandwich

hybridization assay. Following bioassay, [Fe(CN)6]3− was released

with 0.1 M HCl/KCl solution and subsequently detected by square

wave voltammetry at a screen-printed carbon electrode resulting in

a detection limit of 3 ng/L (460 fM). Cadmium phosphate loaded

apoferritin modified with a monobase residue (guanine in this

work) via phosphoramidite chemistry was used for detection of

single-nucleotide polymorphisms or SNPs [69]. In this magnetic

bead based sandwich assay shown in Fig. 14.14, the guanine-

modified apoferritin bound to the complementary base at the

mutation site of the sample DNA, cytosine, as this residue did not

bind with the mismatched base on the capture probe. Following

collection, the sample was exposed to acetate buffer (pH 4.6) to

release the cadmium, which was detected by stripping analysis at

a mercury film coated screen-printed carbon electrode. This system

could detect 21.5 attomol SNP DNA, which the authors state should

enable quantitative analysis of nucleic acid without polymerase

chain reaction (PCR) preamplification.

14.3.7 Silica Nanoparticles

Silica nanoparticles (Si NPs) have been successfully used for

electrochemical DNA detection. As silica is inherently inactive

electrochemically, these particles are either loaded with elec-

troactive molecules and used as labels, or employed as scaffolds



Figure 14.14. Schematic of an electrochemical SNP quantitative assay

based on nanoparticle probe and sequential dna hybridization.

for DNA immobilization resulting in higher probe density and

improved electron transfer to the underlying electrode. The use

of [Co(bpy)3]3+-doped Si NPs as labels for electrochemical DNA

detection was reported by Zhu et al. [70]. In this work [Co(bpy)3]3+

molecules were loaded into Si NPs during NP synthesis, and

then conjugated to amine-terminated ssDNA detection probes via

trimethoxysilylpropydiethylenetriamine and glutaraldehyde. This

now electroactive Si NP functionalized detection probe was used

in a sandwich assay format with capture probe DNA immobilized

onto a glassy carbon electrode. The high loading of [Co(bpy)3]3+

molecules in the Si NPs resulted in a 200 pM target DNA detection

limit. The response from a three-base pair mismatch sequence and

noncomplementary sequence was negligible.

Recently, Ma and coworkers [71] reported the use of Si NP

films for enhanced electrochemical DNA detection. Si NPs were



deposited onto a p-aminothiophenol SAM on an Au electrode by

either electrodeposition of the Si NPs from a silica sol, or adsorption

of the Si NPs upon dipping the SAM-modified electrode into a silica

sol/Si NP solution for several hours. Electrodeposition of the Si

NP provided the best Si NP loading, increasing immobilized ssDNA

probe density and improved electron-transfer kinetics compared

to the SAM-modified electrode alone. FIS allowed for detection

of target DNA hybridization with a detection limit of 1.5 pM and

discrimination between single or double base pair mismatched

DNA sequences. A similar Si NP-SAM-modified Au electrode system

employing [Co(bpy)3]3+ as the electroactive reporter and differen-

tial pulse voltammetry for DNA detection has also been reported

[72].

14.3.8 Liposomes

Liposomes are aggregates of amphiphilic block copolymers or sur-

factant molecules that self-assemble into spherical nanostructures

in aqueous solution. Typically, liposomes consist of a bilayer in

which hydrophilic blocks of the polymer form the outer and inner

shell of the bilayer while the hydrophobic blocks lie between the

inner and outer shell. This configuration shields the hydrophobic

blocks from the external aqueous solution and the aqueous internal

core of the liposome. Liposomes can be functionalized with various

biomolecules and loaded during the self-assembly process with

reporters facilitating use of liposomes as effective labels for DNA

detection.

Patolsky et al. [73] reported the use of 220 ± 20 nm diam-

eter negatively charged liposomes with maleimide functionality

for electrochemical detection of DNA. Thiol-terminated detection

probe DNA was immobilized onto the maleimide functionalized

liposomes yielding 50 to 60 bound DNA probes per liposome.

These DNA-modified liposomes were hybridized to captured target

DNA which was previously immobilized onto a probe DNA-

modified Au electrode in a sandwich assay format, as shown in

Fig. 14.15A. The strong negative surface charge of the DNA-modified

liposomes prevented nonspecific interactions with the negatively

charged electrode surface, providing very low background signals.



Figure 14.15. (A) Amplified electrochemical sensing of an analyte DNA

using oligonucleotide-functionalized liposomes and FIS as a means of

transduction. (B) Electrochemical sensing of an analyte DNA using a

biotinylated oligonucleotide, avidin, liposome labeled with biotin as an

amplification conjugate, and FIS as a means of transduction.



In addition, hybridization of the liposome-labeled detection probe

DNA to the electrode led to a significant negatively charged electrode

interface, which repelled the negatively charged [Fe(CN)6]3– redox

probe. This increase in resistance to charge transfer to the redox

probe was monitored using FIS resulting in a detection limit of

1.2 pM. The authors also used a biotinylated detection probe DNA

that after binding to target DNA on the electrode surface (see

Fig. 14.15B), was treated with avidin, allowing subsequent capture

of biotinlyated liposomes. This was again followed by treatment

with avidin and biotinylated liposomes forming large aggregates of

liposomes yielding a very high negatively charged surface density.

The detection limit for this system was 50 fM.

In an extension of this work biotin-labeled liposomes were also

modified with horseradish peroxidase (HRP) via periodate oxidation

chemistry [74]. The HRP loaded biotin-labeled liposomes catalyzed

oxidation of 4-chloro-1-naphthol in the presence of H2O2 yielding an

insoluble product which precipitated onto and fouled the electrode.

FIS was used to monitor resistance of electron transfer to the

[Fe(CN)6]3− redox probe resulting a similar detection limit of 650 fM

for a DNA sequence relevant to Tay-Sachs disorder. The authors also

extended these various approaches to probe and amplify the signal

from single-base mismatches in analyte DNA [75].

Loading the aqueous interior of liposomes with electroactive

molecules has also been reported. Liposomes prepared with

cholesterol-labeled detection probe DNA and loaded with

[Fe(CN)6]3− were used in a magnetic bead based sandwich assay

in a glass-chip PDMS microfluidic device [76]. Collected by the

magnet upstream of an interdigitated ultramicroelectrode array

(IDUA), the liposomes were lysed by addition of detergent. The

released [Fe(CN)6]3− was subsequently detected at the downstream

IDUA. The assay took less than 30 minutes to perform, including

hybridization time, and could detect 1 fmol DNA. This electroactive

liposome magnetic bead-based sandwich assay was also recently

used by the authors to detect the mRNA amplified from a single

oocyst (an immature ovum) within a PMMA biosensor [77].

Liposomes functionalized with reporter DNA and loaded with

[Ru(NH3)6]3+ were also recently reported for effective electrochem-

ical DNA detection [78]. Used in a competitive assay format on Au


DNA Detection Using Carbon Nanotubes 461

NP–modified screen-printed carbon electrodes, the reporter DNA-

modified liposomes hybridized directly to thiol-terminated capture

probes bound to the Au NPs via Au–thiol interaction. In the presence

of target DNA strands specific to E. coli O157, liposome-labeled

reporter DNA was displaced from the electrode surface. Remaining

[Ru(NH3)6]3+ loaded liposomes were quantified via square wave

voltammetry. This “signal-off” mechanism provided a detection limit

of 150 fM (0.75 amol in 5 μL).

14.4 DNA Detection Using Carbon Nanotubes

There is enormous interest in utilizing carbon nanotubes (CNTs)

in biosensors primarily due to the high surface area, extraordinary

mechanical properties, electron-transport properties, and high

thermal and electrical conductivity of these materials. These one-

dimensional materials (1D) are attractive for the detection of

minor surface perturbations due to binding events. In the case of

single-walled carbon nanotubes (SWCNTs) the structure is such

that every carbon atom is on the surface, thus any event such

as DNA hybridization strongly influences the electronic behavior

of the material. Based on their structure, CNTs can be either

single- or multiwalled (MWCNTs), and envisioned as cylindrical

roll-ups of one or more sheets of graphene. These nanomaterials

have a high aspect ratio with diameters as small as 0.4 nm

for SWCNTs and 2 to 100 nm for MWCNTs and lengths from

tens of nanometers to several micrometers. Although both have

been studied as biosensor materials, MWCNTs, because of their

higher complexity, have been studied more frequently as a bulk

material where ordered structuring may not be as critical. When

integrating CNTs onto a substrate, controlling geometric structure

and orientation can provide enhanced electrochemical responses

due to their fast electron-transfer characteristics.

Their unique structural, mechanical, and electrical properties

differ greatly from other carbon materials used in electrochemical

measurements such as diamond, graphite, and glassy carbon. As

compared to graphite, SWCNTs have a greater surface area and

a much lower density. The unique differences of these materials



are strongly dependent on physical properties such as chirality,

diameter, and length. For example, CNTs are either metallic

conductors or semiconductors based on chirality of the structure

[79], whereas diamond is insulating and graphite is semimetallic.

Historically, MWCNTs were the first to be observed in 1991 by

Dr. Sumio Iijima [80] and shortly thereafter SWCNTs were syn-

thesized by arc discharge [81]. Now CNTs are synthesized by

arc discharge of graphite, laser vaporization, and chemical vapor

deposition methods. To date, the use of CNTs for electrochemical

biosensing has been summarized in several excellent reviews [82],

with recent reviews specifically on DNA functionalization of CNTs

[83]. Presented here are recent innovations in electrochemical

DNA detection using CNTs followed by a description of their

implementation into sensing devices. Before expanding on these

areas, a brief overview of key methods used for functionalization

of CNTs is provided since it is a prerequisite to immobilize

biomolecules on CNTs in a reliable manner.

14.4.1 Functionalization of Carbon Nanotubes with DNA

The potential use of CNTs as electrochemical DNA sensors depends

greatly on their solubility in aqueous media as well as routine

assembly into integrated devices. DNA and other biomolecules

have been successfully immobilized on CNTs by various covalent

and noncovalent binding methods [84]. For covalent attachment,

CNTs are typically activated by chemical oxidation in strong acids,

resulting in the formation of various oxygenated functional groups,

the most prevalent being carboxylic acid groups at the reactive

open ends of the tube, or defect sites at the side walls. This not

only increases their solubility but also presents opportunity for

further modification of the nanotubes. Esterification or amidation

reactions can then be carried out on the oxidized CNTs using

either acid chlorides as intermediates, or carbodiimide coupling

agents [85]. The modified CNTs then react directly with DNA

either targeting an amine site or a thiol site introduced at the

5’ end of the DNA molecule. This approach is primarily used to

functionalize the ends of the nanotubes, and although it is fairly

simple, it is not specific. In comparison, sidewall functionalization



is more difficult but can be achieved using highly reactive species

such as fluorine, nitrenes, arylation using diazonium salts, 1,3-

dipolar cycoadditions, and addition of carbenes to name a few [86].

These methods allow the incorporation of various reactive groups

(–––COOH, –––NO2, OH, H, and ====O) with high specificity for attach-

ment of DNA or other biomolecules. Furthermore, photochemistry

has been used to functionalize the sidewalls of MWCNTs [87]. CNTs

photoetched with azidothymidine serve as photoadducts, with a

reactive group on each photoadduct for the subsequent in situsynthesis of DNA oligonucleotides. This method may potentially

enable photolithographic patterning of different DNA sequences on

CNTs arrayed on genomic chips. The covalent modification of CNTs

can completely change the electronic properties of the CNTs as a

consequence of the transformation of the sp2 hybridization of CNTs

to sp3 hybridization. This can lead to partial loss of conjugation

affecting electron-acceptor and/or electron-transport properties. A

vast amount of work has been conducted on CNT functionalization in

the last decade to overcome these challenges since covalent coupling

of biomaterials to CNTs is critical to the development of biosensors

as well as bioelectronic devices. In contrast to the traditional

approach of covalent modification, noncovalent modification of the

sidewalls for sensor applications has been shown to preserve the

desired electronic and optical properties of CNTs while improving

their solubilities. The earliest work on DNA linkage to CNTs was

through noncovalent interactions [88] and has continued to be used

as a nondestructive functionalization method in the construction

of field-effect transistor (FET)-based biosensors [89]. Sidewalls are

functionalized noncovalently through π stacking or hydrophobic

interactions. DNA bases interact with CNTs via π stacking on the

nanotube surface, with the hydrophilic sugar–phosphate backbone

exposed to the solvent, thereby achieving solubility in water. Zheng

et al. [90] demonstrated DNA-assisted dispersion of CNTs in water

during sonication. Noncovalently wrapped DNA-CNTs were then

separated via ion exchange chromatography. Further, the wrapping

of SWCNTs with ssDNA was found to be sequence dependent [91].

Through a systematic search of a ssDNA library, it was found that

the selected sequence, d(GT)n, n = 10 to 45, self-assembles into

a highly ordered structure around individual nanotubes in such



a way that the electrostatics of the DNA–CNT hybrid depends on

tube diameter and electronic properties. This assembly enabled

improved metal from semiconducting tube separation and also

diameter-dependent separation. DNA can also enter the internal

cavities of CNTs; and electrophoretic transport through a single

MWCNT cavity has been imaged through fluorescence microscopy

[92]. The main disadvantage of noncovalent interactions, however,

is their lack of specificity, and in some cases, denaturing of the

biomolecule upon adsorption.

14.4.2 CNTs for Electrochemical DNA Sensing

Carbon nanotube electrodes have been used for the electrochemical

characterization of DNA. Wang et al. [93] conducted voltammetric

studies on the electrochemical oxidation of guanine and adenine

residues in DNA at SWCNT-modified electrodes. Compared to

nonmodified electrode materials, the electrochemical response

corresponding to the oxidation peaks was greatly enhanced at

the modified electrode. Guo et al. [94] covalently attached both

single-stranded and double-stranded calf thymus DNA molecules

onto MWCNT-modified gold electrodes and characterized electro-

chemical differences by cyclic voltammetry and electrochemical

impedance analysis. This was done using both a redox indicator

[Fe(CN)3−6 /Fe(CN)4−

6 ] and an electrochemical intercalator (ethidium

bromide). Both of these studies suggested that further application

of CNT-modified electrodes might be exploited for detecting DNA

hybridization.

Baker et al. [95] were the first to report the formation of

DNA–SWCNT adducts in solution for DNA hybridization. The

DNA–SWCNT complexes were synthesized by reacting oxidized

SWCNTs with thionyl chloride and ethylenediamine to form amine-

terminated sites, shown in Fig. 14.16. The amines were further

reacted with succinimidyl 4-(N -maleimidomethyl)cyclohexane-1-

carboxylate (SMCC) forming maleimide groups which reacted with

thiol-modified DNA. To confirm covalent attachment and to test

the accessibility of the DNA-modified SWNTs, hybridization studies

were conducted using fluorescently labeled DNA oligonucleotide

targets. Shortly thereafter, the use of carbodiimide-assisted coupling



Figure 14.16. Scheme for fabrication of covalently linked DNA-nanotube

adducts.

of amine functionalized DNA to oxidized SWCNTs in solution was

demonstrated [96].

However, Cai et al. [86] were the first to demonstrate the

use of CNTs in an electrochemical DNA biosensor fabricated by

covalently immobilizing a DNA probe onto a MWCNT-modified

glassy carbon electrode and detecting the hybridization of target

DNA by differential pulse voltammetry (DPV) using an electroactive

intercalator, daunomycin, as an indicator, illustrated in Fig. 14.17.

The MWCNTs served as a method of covalent attachment of probe

DNA, but also improved the sensitivity of this electrochemical assay.

A detection limit of 1.0 × 10–10 M was achieved whereas previous

results reported by Marrazza et al. [97] using similar experiments

with the probe DNA directly attached to nonmodified carbon

electrodes gave a detection limit of 1 μg/ml of target sequence. The

use of MWCNTs led to an increased rate of heterogeneous electron



Figure 14.17. Schematic representation of the enhanced detection of DNA

hybridization based on a DNA-MWCNT sensor using daunomycin as the

electroactive indicator.

transfer between the electrode and intercalator, but also increased

the effective area of the electrode.

Hybridization events were also monitored by DPV measurement

of the reduction of intercalated daunomycin on a glassy carbon

electrode modified with MWCNTs and platinum nanoparticles

dispersed in Nafion [98]. Probe DNA was attached in a similar

manner through the formation of amide bonds between the –COOH

on the MWCNTs and –NH2 of the oligonucleotides. With the addition

of the Pt nanoparticles, the detection limit of this glassy carbon-

modified electrode to hybridized complementary DNA sequences

was lowered by an order of magnitude to 1.0 × 10−11 M as compared

to that reported by Cai (1.0 × 10−10 M) using only MWCNT-

modified GCEs. Whereas CNTs promote electron-transfer reactions,

the nanoparticles further amplified the signal due to their high

catalytic activity toward daunomycin reduction.

He and Dai [86a] prepared aligned SWCNT–DNA sensors by

chemically coupling ssDNA probes on both the tip and wall of

plasma-activated aligned carbon nanotubes on gold electrodes.



Gold-supported aligned nanotubes were generated from pyrolysis

of iron(II) phthalocyanine. They were next chemically activated

with an acetic acid-plasma treatment followed by covalent coupling

of the generated carboxylic acid groups with amine-terminated

ssDNA. The strong oxidation peak measured at 0.29 V, due to

ferrocene-labeled complementary oligonucleotides, was used to

verify hybridization events. Also observed was a much higher

amperometric response from the aligned SWCNT–DNA modified

electrode as compared to electrodes immobilized with ssDNA

probes without SWCNTs (ca. 20 times). Improved electrochemical

performance of almost all electrode materials has been observed

when modified with nanotubes. MWCNT-modified glassy carbon

electrodes have shown an enhanced signal when used for label-free

DNA analysis based on the oxidation of guanine bases [99]. Similar

amplification of the guanine response has been reported at MWCNT

carbon paste electrodes [100], SWCNT glassy carbon electrodes

[101], and on graphite pencil electrodes modified with MWCNTs

[102]. Electrochemical AC impedance measurements provided

another label-free approach to DNA hybridization detection on a

DNA probe-doped polypyrrole film on MWCNT-modified electrodes

[103]. A 5-fold enhancement in sensitivity was reported.

Ultrasensitive detection of DNA hybridization was shown by

combining a CNT-modified nanoelectrode array with [Ru(bpy)3]2+

mediated guanine oxidation, shown in Fig. 14.18 [104]. Vertically

aligned MWCNTs were grown by plasma-enhanced chemical vapor

deposition on UV-lithographic patterned electrodes on a Si [100]

wafer. DNA probes were covalently coupled to the nanotubes

through carbodiimide chemistry. The hybridization of subattomole

DNA targets was detected using cyclic voltammetry, improving the

sensitivity of DNA detection by orders of magnitude compared to

methods where DNA is immobilized directly on a conventional

electrode material. Interestingly, by lowering the nanotube density,

greater sensitivity was achieved. This was due in part to the

use of AC voltammetry (ACV). With higher density samples, ACV

results were inconsistent, however, this electrochemical technique

worked well with low-density arrays. The unstable exponential

background current, characteristic of the CNT arrays, was filtered

out by the phase-sensitive ACV technique, and only the Faradaic



Figure 14.18. SEM images of array of MWCNTs at UV-lithography

patterned Ni spots (left) and polished MWCNT array (right). The schematic

mechanism of [Ru(bpy)3]2+ mediated guanine oxidation.

current associated with [Ru(bpy)3]2+oxidation was measured. This

MWCNT-modified nanoelectrode array was also applied for label-

free detection of PCR amplicons [105].

The use of CNTs as carriers of metal tags has been used

to amplify DNA hybridization detection [106]. CdS nanoparticles

were loaded onto acetone-activated CNTs and further function-

alized with streptavidin. The SWCNT-CdS–streptavidin conjugates

reacted with biotinylated DNA probes. Hybridization of these

probes to complementary oligonucleotides anchored on a support

was detected by stripping voltammetric measurements of the

dissolved CdS particles. Approximately 500 particles were loaded

on a single nanotube, effectively lowering the detection limit

by 500-fold when compared to that achieved using a single

nanoparticle label typical of this type of sandwich assay. Another

effective amplification method developed by Wang’s group used

CNTs in a dual amplification role in both the recognition and

transduction events [107]. CNTs were used as carriers of alkaline

phosphatase (ALP) enzyme tags (9600 enzyme molecules/CNT) and

as transducers for accumulation of the product of the enzymatic

reaction, α-naphthol. The enzyme-functionalized CNTs were further

modified with DNA probes. Magnetic particles were functionalized



Figure 14.19. Electrochemical DNA detection using ALP-loaded CNT tags.

with a second DNA probe, which hybridized to a complementary

oligonucleotide. This complex then hybridized with DNA probes

attached to the CNT-enzyme conjugates, demonstrating the first

example of using DNA for linking particles to CNTs, as shown

in Fig. 14.19. The catalytic hydrolysis of α-naphthylphosphate to

the electrochemically detectable α-naphthol product by the bound

enzymes gave a 104-fold improvement in the sensitivity compared

to a single ALP tag. Further amplification was achieved by using

a CNT-modified glassy carbon electrode, increasing the electrode

area for the chronopotentiometric detection of the enzymatic

product. Coupling the two amplification steps (CNT-enzyme tags

and preconcentration of CNT transducers) yielded a dramatic

enhancement in sensitivity, allowing an extremely low detection

limit of 1.3 zmol in a 25 μl sample. This corresponds to 820 copies

in the sample size. Further amplification, with detection of DNA

down to 80 copies, was achieved with the enzyme-coated CNT tags

when they were prepared by using a layer-by-layer self-assembly

technique, maximizing the ratio of enzyme tags per binding event

[108].

A sensitive, indirect method of detecting hybridized DNA was

conducted by preparing ferrocene (Fc)-SWCNT adducts coupled

with a DNA probe [109]. Ferrocene noncovalently interacts with

SWCNTs through π–π interactions. The Fc-SWCNT adducts were

then further conjugated with DNA probes covalently through the

amide linkage between the primary amine at the 3′ end of the DNA

probe and the carboxylic acid groups on the CNTs. The Fc-SWCNT–

DNA probe hybridized to a target sequence already hybridized



to another DNA probe immobilized on a gold electrode. In this

sandwich assay the amplified electrochemical response was due to

the ferrocene-catalyzed reduction of H2O2.

14.4.3 Progress toward CNT-Based Sensors for DNADetection

The integration of CNTs into large-scale assemblies or circuits

requires precise control in the placement of the carbon nanotubes.

Several groups have used DNA to direct the assembly of CNTs

between metal electrodes to form a simple circuit. For example,

Hazani et al. [110] assembled a monolayer of probe thiol-terminated

ssDNA on two neighboring gold contacts. SWCNTs were modified

at the ends with complementary DNA sequences. Hybridization

between the complementary strands and the immobilized probes

resulted in bridging the gold contacts by the SWCNT. Current–

voltage (I –V ) curves were measured on electrode pairs using both

complementary and noncomplementary DNA-SWCNTs to bridge the

electrodes. The currents measured from the noncomplementary

interactions were an order of magnitude smaller compared to

currents measured from covalently bridged electrodes. Taft et al.[111] demonstrated that selective coupling of DNA to either the ends

or sidewalls of CNTs was highly specific and based on the DNA–CNT

linkage scheme. Either amine-terminated DNA was immobilized

on CNTs through free carboxyl groups, or pyrene-modified DNA

was immobilized noncovalently through hydrophobic interactions.

Complementary DNA was immobilized on gold particles and

hybridized with the CNT–DNA probes (covalently or noncovalently

attached). SEM images showed that the gold particles were bound

primarily at the ends of the CNTs when covalent functionalization

of probe DNA was used, or were located at the sidewalls when

probe DNA was attached noncovalently. Similarly, transmission

electron microscopy (TEM) was used to verify the binding of gold

nanoparticles modified with complementary strands of DNA to

probe DNA attached to SWCNTs [112]. Two methods for attaching

DNA to SWCNTs in either aqueous solution or in organic solvent

were developed and both methods resulted in selective attachment

of the DNA-modified gold particles at the tips of SWCNTs. These

methods for binding DNA provide versatility for modification of



CNTS with DNA, but also demonstrate the potential for controlled

placement of nanotubes into sensors, FETs, or other electronic

devices. For practical device applications, the density of CNTs needs

consideration; an individual nanotube has a high risk of channel

failure. Baek et al. [113] recently demonstrated that the performance

of SWCNT devices for DNA hybridization was dependent on SWCNT

film density. SWCNT networks of varying density were deposited

onto glass slides. Using photolithography and reactive ion etching,

defined lines of SWCNT networks were patterned. Finally, using e-

beam lithography, metal evaporation of Co/Au was performed in

order to fabricate the final two-terminal device. Covalent attachment

of probe DNA via amide coupling to the SWCNT film spanning

the electrodes was conducted followed by complementary DNA

hybridization. The electrical behavior of hybridization at varying

film densities was determined from I –V curves and it was shown

that as the nanotube network density decreased, conductance

increased, with an optimum range of film density. This behavior is

likely due to an optimization in the number of reaction sites as well

as the conductance of the film.

SWCNTs were first used to fabricate FET devices in 1998 [114].

Since then several groups have fabricated CNTFETs, where absorbed

molecules that modulate the nanotube conductance replace the

solid-state gate. Only in the past few years have a small number

of research groups applied CNTFET devices for DNA detection.

DNA molecules can be selectively attached to either the nanotube

or at the metal electrodes. Hybridization of complementary DNA

at the nanotube is thought to mostly influence the electronic

response of the FET by electron depletion in the channel, whereas

binding at the electrodes modifies the metal work functions, that

is, the Schottky barrier [115]. Tang et al. [116] examined this

experimentally and found that the electrical conductance change,

observed when DNA hybridized to the device, was due to binding

at the gold electrodes instead of the sidewalls of the nanotube. Thus

the Schottky barrier modulation appeared to play a more significant

role in DNA detection. CNTFET devices have been fabricated using

peptide nucleic acid (PNA) oligonucleotides immobilized on the

gold surfaces [117]. Binding of complementary DNA resulted in an

increase in conductance corresponding to the increase in negative

surface charge density associated with binding of the negatively



charged complementary DNA molecules. As a practical application,

an allele-specific assay was developed to detect the presence of

single nucleotide polymorphisms (SNPs) [91a]. A network of carbon

nanotubes within an FET device was functionalized with either

mutant or wild-type alleles while DNA hybridization to the corre-

sponding allele was measured through a decrease in conductance.

One last example is an amplification method reported in a CNTFET

device. As noted earlier, the incorporation of nanoparticles has been

shown to further enhance an electrochemical response of a CNT-

modified electrode [99]. Dong et al. [118] report on the use of a

SWCNTFET to detect target DNAs labeled with Au nanoparticles.

Hybridization events of CNT–DNA probes were detected down to

100 fM due to an increased conductivity through the NPs in close

proximity.

14.5 Conclusion

The integration of nanotechnology with biology and electrochem-

istry has produced many advances for novel DNA sensing strategies.

The ability to synthesize many nanomaterials in similar size

ranges to biomolecular markers makes their coupling with DNA

extremely efficacious toward the design of sensors to transduce

DNA hybridization events. Nanoparticles, nanorods, nanowires,

nanotubes, etc. have been used as labels in novel DNA detection

assays that are fast, sensitive, and reliable. Their use as supports

for DNA immobilization, tags that take advantage of their intrinsic

atomic make-up or ability to load and pre-concentrate secondary

labels, and ability to modulate interfacial phenomenon have been

described. Further advances are expected to lead to new generations

of electrochemical DNA sensors with implications in fields such as

medical diagnostics, drug discovery, detection of biothreats, and

environmental monitoring.

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Chapter 15

Electrochemical Genosensor Assay forthe Detection of Bacteria onScreen-Printed Chips

Chan Yean Yeana*, Lee Su Yinb, and Manickam Ravichandranb

aDepartment of Medical Microbiology and Parasitology, School of Medical Sciences,Universiti Sains Malaysia, Kota Bharu, MalaysiabFaculty of Applied Sciences, AIMST University, 08100 Semeling, Kedah, Malaysia*[email protected]

Electrochemical genosensors for the detection of bacteria were

introduced about a decade ago. Miniaturization and advanced

microfabrication technology have made it compatible with bacteria

DNA diagnostic. This technology is cost effective, fast, and accurate.

The bioaffinity and biocatalysis reactions generate amperometric,

voltametric, impedimetric, or conductimetric signals on screen-

printed transducer chips (SPC), which is proportional to the number

of immobilized DNA copies on the SPC surface. Electrochemical

genosensor assays give quantitative rather than qualitative results.

Furthermore, the use of a hand-held portable reader makes this

assay suitable for use in the field, especially for point-of-care (POC)

tests at the patient bedside, during surveillance and environmental

studies of microorganisms.



482 Electrochemical Genosensor Assay for the Detection of Bacteria on Screen-Printed Chips

15.1 Introduction

In clinically and epidemiologically severe infectious diseases, the

rapid identification and detection of the causative organism is

crucial for effective control, management, and prompt treatment of

the infection. The conventional laboratory methods involve culture,

microscopy, and biochemical tests [1]. This process is laborious and

takes 2 to 4 days or longer to obtain a result. Culture methods often

lack sensitivity, especially for poorly handled samples or clinical

samples from patients previously treated with antibiotics [2].

Polymerase chain reaction (PCR) has been used extensively

as a diagnostic tool in various fields, such as genetic screening,

infectious disease diagnosis, forensics, environmental monitoring,

and veterinary science. PCR is an enzymatic process in which

specific regions of DNA are amplified in vitro. This process amplifies

the target DNA exponentially to generate billions of copies from

a single copy in less than 1 h [3]. The conventional detection of

PCR amplicons by electrophoresis exposes the user to hazardous

chemicals, such as ethidium bromide and ultraviolet light. Other

safer detection techniques, such as capillary blotting and enzyme-

linked immunoassays, require multiple hybridization and washing

steps, which are labor-intensive and time consuming.

15.2 Methods for the Detection and Identification ofMicroorganism Utilizing Enzyme-Based Genosensorson Screen-Printed Chips

15.2.1 Electrochemical Genosensors for the Detection ofBacteria

Electrochemical genosensors for the detection of bacteria were

introduced about a decade ago. Miniaturization and advanced

microfabrication technology have made it compatible with bacteria

DNA diagnostic. This technology is cost effective, fast, and accurate.

The bioaffinity and biocatalysis reactions generate either amper-

ometric, voltametric, impedimetric, or conductimetric signals on

screen-printed transducer chips (SPC), which is proportional to the

number of immobilized DNA copies on the SPC surface.


Methods for the Detection and Identification of Microorganism 483

Electrochemical methods are well suited for molecular diagnos-

tics of microorganisms on the genomic and proteomic level. Electro-

chemical reactions can be designed to produce a direct electronic

signal using a portable handheld and inexpensive electrochemical

analyzer (AndCare, PalmSens, DropSens etc.) that is commercially

available in the market, without any expensive signal transduction

equipment [4].

Numerous electrochemical platforms have been developed for

DNA detection, including direct electrochemistry of the DNA

bases [5], electrochemistry of different polymer-modified screen-

printed chips [6], electrochemistry of DNA-specific redox indicator

molecules or enzymes [7, 8], electrochemistry of signal amplification

with nanoparticles (NPs) such as gold, silver or magnetic particles

[9, 10], and dsDNA π -stacked mediated charge transport chemistry

[4, 7, 11, 12].

A genosensor for bacterial detection should possess the fol-

lowing criteria: sensitive (able to detect the bacterium in a small

sample), specific (able to distinguish the target from non-target

strains), precise, rapid and able to perform direct measurement

without pre-enrichment. In addition, it would be desirable if

the genosensor is portable or handheld, affordable and can be

performed even by untrained personnel.

Biosensors for bacterial detection involve biological recognition

components such as presence of the biomarkers, nucleic acid,

antibodies or aptamer attached on a transducer. However, in this

chapter, bacterial nucleic acid will be described in detail [13].

Electrochemical genosensors for detection of various bacterial

species have been described in Refs. 8 and 13–19. Reliable detection

assays have been developed for pathogenic bacteria such as

Salmonella sp, E. coli 0157:H7, Staphylococcus aureus and Vibriocholerae that cause major worldwide foodborne outbreaks [8, 13, 16,

19, 20].

The general scheme of a genosensor assay development starts

with the immobilization of the specific nucleic acid sequence

(“probe”) on the transducer surface. The presence of the com-

plementary sequence (“target”) in the sample is recognized and

captured by the probe through hybridization.



A probe is an oligonucleotide (single stranded DNA or RNA)

with the size of 18 to 25 bases. The difference between probe and

primer is the function of the oligonucleotide. Primer is used during

polymerase chain reaction (PCR); however probe is used during

capturing of the target DNA during hybridization.

Short single-stranded synthetic targets (∼20–60 bases oligonu-

cleotides) are used to evaluate the technology platform and

specificity of the DNA sequences in identification of bacteria before

the assay is evaluated with PCR amplicons.

The DNA duplex formation can be detected based on the

incorporation or association of a hybridization indicator or changes

accrued from the hybridization event. Different indicators can be

used in detection of DNA based on the appropriate electrochemical

activity selected, it can be either label-free (e.g. guanine, adenine),

or label-based (enzyme-based, ferrous and ferricyanide, Ruthenium

bipyridine [Ru (bpy)], methylene blue, Ethidium bromide etc).

The hybridization event is detected via the increase or decrease

in signal of the redox indicator or changes in conductivity or

impedance/capacitance.

Most genosensors are designed to detect bacterial DNA that

is first amplified by PCR. The species-specific detection of the

bacteria mainly depends on hybridization of the specific probes to

complementary sequences in the PCR amplicons. PCR primers and

hybridization probes are designed using bioinformatic softwares

to ensure high specificity and sensitivity. The bacterial genetic

sequences can be obtained from GenBank of National Center

for Biotechnology Information (NCBI) (www.ncbi.nlm.nih.gov). The

target genes selected will depend on the purpose of the test,

either for identification of infectious bacteria, detection of antibiotic

resistant genotypes (MRSA, VRE, ESBL), bioterrorism (anthrax,

Burkholderia pseudomallei or melioidosis, Yersinia pestis or plague),

food pathogens (V. cholerae, Clostridium botulinum, Escherichia coli0157:H7), fecal contamination (Enterococcus species) or environ-

mental surveillance. Then target gene sequences will be down-

loaded from Genbank and aligned by multiple sequence alignment

softwares available such as VectorNTI (Invitrogen Corporation,

California, USA). The highly specific gene regions (conserved

regions) will be selected for primer designing.



Not all bacteria are pathogenic or harmful to humans. Some

microorganisms are harmless or even some are very useful for

human beings. An example is the lactobacilli in human stomach

that helps in converting lactose and other sugars to lactic acid.

However, these bacteria will cause disease if they are detected in

environments that are not their normal habitat. Thus, the presence

of certain bacteria out of their normal habitat is an indicator of a

certain disease or contamination. For example, Enterococcus species

is used as an indicator of fecal pollution in environmental waters,

while the detection of species-specific Enterococcus faecium is used

as an indicator of human fecal pollution [13]. On the other hand,

the presence of some bacteria almost certainly indicate an infection;

for example, Mycobacterium tuberculosis causes tuberculosis, and

Streptococcus and Pseudomonas cause pneumonia.

An enzyme-based genosensor for amperometric detection of

PCR amplicons on screen-printed carbon (SPC) chips was recently

described in Ref. 8. The SPCs were pretreated with streptavidin

before each experiment. Covalent agent (200 mM 1-ethyl-3-

[3-dimethylaminopropyl]carbodiimide and 50mM N -hydroxy-

succinimide prepared in 0.05 M phosphate buffer) was added to the

working electrode of the SPC and incubated at room temperature.

The electrodes were washed by dipping them once in deionized

water. Streptavidin (0.05 mg/mL) was then pipetted onto the

working electrode again to form a meniscus and incubated at room

temperature. The electrodes were washed by dipping them once in

deionized water (a schematic diagram is shown in Fig. 15.1A). The

unbound area on the streptavidin-treated SPC reservoir area was

blocked with 1 M ethanolamine chloride.

The PCR amplicons were captured on the electrodes and detected

using a portable pulse amperometric reader (AndCare, Durham,

NC). A schematic diagram of the detection process is shown in Fig.

15.1B. Briefly, the biotin- and fluorescein-labeled PCR amplicons

were diluted with an equal volume of 0.05 M phosphate buffer,

and the diluted PCR amplicons were applied to the surface of the

working electrode for 5 min. During the incubation step, the biotin-

labeled strands of the PCR amplicons were specifically captured

on the streptavidin precoated working electrode. The excess PCR

amplicons were removed by dipping the electrode 10 times into



Figure 15.1A. The electrode was washed by dipping once in deionized

water.

0.1× saline sodium citrate (SSC) containing 0.5% sodium dodecyl

sulfate (SDS). After the washing step, the SPC was incubated with

horseradish peroxidase (HRP)-conjugated anti-fluorescein antibody

diluted in the ratio of 1:200. During this step, the antibody is bound

to the fluorescein-labeled strand of the PCR amplicons. The SPC was

then washed in 0.1× SSC containing 0.5% SDS [8].

An HRP substrate was prepared containing a mixture of 3,3′,5,5′-tetramethylbenzidine and H2O2 in a 1:10 ratio, and this substrate

mixture was applied to the SPC reservoir area to cover the

working, counter and reference electrodes. The enzymatic reaction

occurring on the working electrode was detected using a portable

pulse amperometric reader. The reader used intermittent pulse

amperometry in which a 15 s incubation period was followed by an

applied potential of −0.1 V (vs. a silver pseudoreference electrode)

with a measurement time of 10 s and a pulse time of 10 s at a

frequency of 5 Hz and a current range of 10 μA [8].

15.2.2 Principles of Enzyme-Based PCR Amplicons TargetDNA Detection Methods

15.2.2.1 Direct method

In general, this approach for species-specific identification of

bacterial pathogens involves immobilization of single-stranded



Figure 15.1B. The PCR amplicons were captured on the electrodes

and detected as described in the AndCare company protocol, with some

modifications.



oligonucleotide capture probes onto the transducer surface and

hybridization of single-stranded oligonucleotide target which is

labeled with haptens (biotin, fluorescent, digoxigenin, etc.) on one

end during PCR. Thus, this method can only be applied on PCR

amplicons that are labeled with haptens on one end. The two-

component complex on the transducer surface will enable binding

of a specific anti-hapten-conjugated HRP or alkaline phosphatase

(AP) reporter enzyme to the probes-target complex. Addition of the

enzyme-specific redox substrate and application of a fixed potential

between working and reference electrodes on the transducer

surface generates an enzyme-mediated redox cycle and detected

in the form of a current (Fig. 15.2A). The electroredox current

amplitude reveals the concentration of the probe-target complexes.

15.2.2.2 Indirect method

This approach for species-specific identification of bacterial

pathogens involves immobilization of a single-stranded oligonu-

cleotide capture probe onto the transducer surface, followed by

hybridization of single-stranded oligonucleotide targets and a

detection probe which is labeled with haptens (biotin, fluorescent,

digoxigenin, etc.) on one end. This method can be performed even

without labeling the PCR amplicons. The detection of the three-

component “sandwich” complex (capture probes-target-detection

probes) on the transducer surface is the same as using the direct

method (Fig 15.2B).

15.2.2.3 Rapid method

In this approach double-stranded oligonucleotide PCR amplicons

are labeled with a hapten on one end and another different hapten

on the other end during PCR. The transducer surface (gold or carbon

screen-printed chips) is treated with protein based anti-haptens

that capture one of the hapten label on the PCR amplicons. The

second hapten label on the PCR amplicons will bind to a specific

anti-hapten-conjugated HRP or AP reporter enzyme to labeled PCR

amplicons. Addition of the enzyme-specific redox substrate and

application of a fixed potential between working and reference



Figure 15.2. Principles of enzyme-based PCR amplicons target DNA

detection by (A) direct method, (B) indirect method, and (C) rapid method.

See also Color Insert.



electrodes on the transducer surface generates an enzyme-mediated

redox cycle and detected in the form of current (Fig. 15.2C). The

electroredox current amplitude reveals the concentration of the

labeled PCR amplicons in the test sample.

15.2.3 Screen-Printed Transducer Surface

The surface structure and chemistry of electrochemical transducers

with regard to the detection of DNA hybridization and electron

transfer (current measurement) have been thoroughly investi-

gated in several studies [13, 21–23]. The immobilization of the

oligonucleotide probe onto the transducer surface will influence the

genosensor performance. The oligonucleotide probe orientation on

the surface will determine the accessibility of the probe to target

DNA. The type of probe, either labeled or unlabeled will depend

on the transducer surface used. For example, on the screen-printed

carbon electrode, the immobilization method of the oligonucleotide

probe on the surface can be either covalent bonding, adsorption

or electrostatic by applying a fixed potential between working and

reference electrodes on the transducer surface.

15.2.3.1 Screen-printed gold chip genosensors

In recent electrochemical genosensor studies, researchers have

started using screen-printed gold chips and modified thiolated

oligonucleotide probe to form a self assembly monolayer on the

chip’s surface. There are many ways to detect the hybridization

of the probe and target DNA, such as enzyme-based (HRP or

AP) or label-free oxidation of guanine bases, anthraquinone-2,6-

disulfonic acid (AQDS) anthraquinone-2-monoisulfonic acid (AQMS)

methylene blue, Ruthenium bipyridine [Ru(bpy)], hexaamineruthe-

nium(III) chloride or ferrocene.

However, on the screen-printed gold electrode, the self-assembly

immobilization of the oligonucleotide probe on the gold surface,

the probe can be either 3’-thiol or 5’-thiol labeled with C3 or C6

linker depending on the signal substances [such as AQMS, enzyme,

MB, Ru (bpy), ferrocene] and the blocking agent used [2-Mercapto-

1-ethanol (MCE), 6-mercapto-1-hexanol (MCH) or 11-mercapto-1-



undecanol). The usage of the different carbon linkers (C3 or C6)

and blocking agents (MCE MCH or MCU) are based on the different

biosensor requirements [12].

A few examples of pathogen detection using enzyme-based

electrochemical genosensors have been developed [8, 13–19]. In

addition, Farabullini et al. (2007) and Liao et al. (2006) published

their microfabrication of simultaneous detection of different food

pathogenic bacteria (Salmonella species, Lysteria monocytogenes,

Staphylococcus aureus and Escherichia coli) and uropathogens

(Escherichia coli, Proteus mirabilis, Pseudomonas aeruginosa,Enterococcus species, Klebsiella species, Enterobacter species andthe Enterobacteriaceae group) in clinical urine specimens by means

of a disposable electrochemical gold genosensor array.

These analytical methods relied on the immobilization of

specific-thiolated probes with the optimized concentration on the

screen-printed arrays of gold electrodes. The unlabeled or unmodi-

fied PCR amplicons from the bacteria genomic DNA were captured

onto the capture probes on the transducer surface via sandwich

hybridization (indirect method). The biotinylated hybrids were

bound to a streptavidin-alkaline phosphatase (AP) or horseradish

peroxidase (HRP) conjugate and then exposed to their subtrates,

α-naphthyl phosphate or 3,3’5,5’-tetramethylbenzidine (TMB)-

hydrogen peroxide (H2O2). Finally, differential pulse voltammetry

measurement was used to detect the signal [20]. Electrochemical

detection can be achieved by monitoring the oxidation or reduction

signal of a substrate after its hybridization with an enzyme-tagged

probe [24]. The analytical signals were observed only at the specific

positions with the corresponding capture probe. The non-specific

signal observed at other position of the array was comparably

negligible.

15.2.3.2 Screen-printed carbon-chip genosensors

In our recent articles, we described the detection of a food-borne

pathogen, Vibrio cholerae, which causes cholera disease. The assay

relied on detection of Vibrio cholerae-specific PCR amplicons using

an electrochemical genosensor on screen-printed carbon chips. The

signal was measured by intermittent pulse amperometry (IPA) using

a portable handheld reader AndCare (Alderon, Durham, NC).



In the assay described above, the screen-printed carbon chips

were first pretreated with covalent agent [N -hydroxy succin-

imide (NHS) and 1-ethyl-3-(3-dimethyl-aminopropyl) carbodiimide

hydrochloride (EDC)] to immobilize streptavidin on the transducer

surface. The surface was inactivated and blocked with ethanolamine

and bovine serum albumin (BSA) to avoid non-specific adsorption

of antibody-HRP-conjugate onto the electrode surface. The labeled

PCR amplicons produced from amplification of V. cholerae DNA were

captured onto the transducer surface without hybridization (rapid

method). The biotinylated hybrids were bound to streptavidin on the

transducer surface and hybridized with antibody-HRP-conjugate.

Addition of the substrate 3,3’5,5’-tetramethylbenzidine (TMB)-

hydrogen peroxide (H2O2) resulted in a signal that was detected

using amperometry (IPA) measurement [8].

Signals were produced only with the specific-labeled PCR

amplicons labeled with the corresponding haptens. The background

signal was low and negligible compared to the signal produced by

positive samples. The enzyme-based electrochemical genosensor

assay concept has shown promising results in the detection of

various analytes [16, 17, 24]. The combination of horseradish per-

oxidase (HRP)-coupled hybridization schemes with electrochemical

biosensors allow highly sensitive detection of targets because the

signal is amplified [18, 24].

15.3 Advantages of the Enzyme-Based ElectrochemicalGenosensors in Detecting Bacteria onScreen-Printed Carbon Chips

Conventionally, PCR amplicons are detected by agarose gel elec-

trophoresis which takes 45 minutes to one hour and the use of

expensive chemicals, such as SYBR Green dye or harmful agents

such as UV light and ethidium bromide. As an alternative method

for PCR amplicon detection, many enzyme-based electrochemical

genosensor assays have been developed and have shown promising

results [16, 18, 25, 26]. Electrochemical DNA hybridization sen-

sors have been reported for pathogens such as Cryptosporidium,

Escherichia coli, Giardia, Mycobacterium tuberculosis Salmonella


Discussions 493

enteritidis, Streptococcus sobrinus and hepatitis B virus [14, 15,

19, 27]. However, most of these electrochemical genosensors have

the drawback that they require an extra hybridization step with a

probe before the PCR amplicon signal is detected The rapid method

for detection of DNA on screen-printed carbon chips described in

this chapter eliminates this hybridization step by labeling the PCR

amplicon with both biotin and fluorescein via modified primers. The

PCR amplicon is directly applied to the modified SPC and the HRP

enzymatic reaction is read within 15 s [8].

Thus, in this assay, rapid method has eliminated the two steps

that are normally included in the conventional electrochemical

genosensor assay: the denaturation of the PCR amplicon and

its hybridization. Here, we merely immobilized the biotin- and

fluorescein-labeled PCR amplicon on a streptavidin-modified SPC,

followed by incubation with HRP-conjugated anti-fluorescein anti-

body, and the direct detection of the amperometric signal [8].

However, there is a need to incorporate PCR and electrochemical

analysis into a single device for this method to be fully usable for

field applications [13].

15.4 Discussions

Electrochemical enzyme-based biosensor techniques can be used

for DNA and immunoassays (antigen–antibody) based on amper-

ometry [16, 18, 26, 28]. Although the SPC was designed for

DNA detection, it can also be used for the detection of bacterial

cells using antigen–antibody interactions. Rao et al. [2] reported

an antibody-based V. cholerae electrochemical biosensor assay

using alkaline phosphatase (AP) and the Autolab PGSTAT 12

potentiostat/galvanostat equipment However, the lowest detection

limit was around 105 CFU/mL, compared to 10 CFU/mL with a

genosensor, hence it is less sensitive than a genosensor. Moreover,

the assay used an AP enzymatic system for detection which requires

more time (10 min) to read the oxidation signals [8].

Conventional DNA microarrays are based on sequence-specific

DNA detection, but their application in diagnostic tests for field

settings is limited by the large biological samples required and



costs and complicated procedures involved. Current electrochemical

genosensors can overcome these drawbacks as these assays are

more affordable, rapid and easier to perform, while maintaining high

sensitivity and specificity [29].

The newly developed concept of the “lab-on-chip” integrates

the chip, and the components for DNA extraction, amplification

and detection, with the advantages of a detection system that

requires only a small sample and few reagents. It is cost effective,

has enhanced rapidity, high-level performance and can be highly

automated [29].

The optimized genosensor procedure used in this study is unique

and universal in that it can detect both biotin- and fluorescein-

labeled PCR amplicons from any organism, allowing the early and

precise diagnosis of infectious agents [8].

Furthermore, the integration of self-assembled monolayer (SAM)

nanoscale chemical structures with an electrochemical sensing

system allows rapid and ultra low concentration sensing assays that

will preclude the need for PCR amplification in the future [19].

15.5 Conclusion

Genosensor assays are more useful and informative than agarose gel

and DNA chromatography-based tests for DNA detection as they give

quantitative rather than qualitative results. Furthermore, the use of

a hand-held portable reader makes it suitable for use in the field.

Therefore, in the future, genosensors will be applicable to a wide

variety of applications, which include identification of antimicrobial-

resistance determinants, other microorganisms or mutant genes in

hospitals and environmental settings.

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M. Ozsoz, Enzyme-linked amperometric electrochemical genosensor

assay for the detection of PCR amplicons on a streptavidin-treated

screen-printed carbon electrode, Anal. Chem. 80, 2774–2779 (2008).


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13. I. Palchetti and M. Mascini, Principles of Bacterial Detection: Biosensors,Recognition Receptors and Microsystems (Amperometric Biosensors forPathogenic Bacteria Detection), Springer, New York (2008).

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Chapter 16

Introduction to Molecular BiologyRelated to Electrochemical DNA-BasedBiosensors

Yalcin Erzurumlu and Petek BallarEge University, School of Pharmacy, Biochemistry Department,35100, Izmir, [email protected]

16.1 Introduction

Molecular recognition is central to biosensor technology. Receptors,

enzymes, antibodies, aptamers, molecular beacons, and nucleic

acids are mainly used as molecular recognition elements in

biosensor development (Chambers et al., 2008). Since 1990, nucleic

acids, especially deoxyribonucleic acid (DNA) have been used as

biorecognition elements in biosensor technology. These biosensors

are named as DNA-based biosensors.

DNA was first isolated in 1869 by Friedrich Miescher as a

phosphorous-containing substance called nuclein (Dahm, 2008). In

1943, Oswald Avery and his colleagues discovered that DNA is the

bearer of genetic information by permanently transforming a non-

virulent form of the organism into a virulent form via transforming



500 Introduction to Molecular Biology Related to Electrochemical DNA-Based Biosensors

DNA taken a heat-killed virulent strain of the bacterium Strepto-coccus pneumonae (Avery et al., 1944). James Watson and Francis

Crick proposed a double helical structure for DNA in 1953 (Watson

and Crick, 1953). In 1958, Kornberg discovered and isolated DNA

polymerase in order to make DNA in a test tube. Kary Mullis and

colleagues invented a technique for multiplying DNA sequences invitro by the polymerase chain reaction (PCR) in 1980 (Mullis KB,

1990).

Diagnosis of genetic disorders is clearly the focused aim of

many research groups, since genetic disorders are an important

health problem among the world. More than 4000 genetic diseases

are known, many of which are debilitating or fatal (McKusick,

1991). Cystic fibrosis (CF), an autosomal recessive disorder, occurs

approximately once in every 3500 live births (Lommatzsch and

Aris, 2009). Exocrine glands and small airways are affected in CF

resulting in death in early twenties. Over 800 mutations leading to

CF have been found. Another example is alpha-1 antitrypsin (A1AT)

deficiency that affects approximately one in 2000 individuals. A1AT

deficiency is a condition in which the liver does not make enough of a

protein that protects the lungs and liver from damage. It is the most

common genetic liver disease in children. This condition can lead

to emphysema and cholestasis, late hemorrhagic disease, or chronic

liver disease (Fairbanks and Tavill, 2008; Gooptu et al., 2009).

Mutations of DNA in cells are the reason of most of the genetic

disorders. Some genetic diseases can be identified by detecting

the defective protein, product of mutated gene. However, there are

many genetic disorders that do not have a characterized change of

a protein. Moreover, many of these genetic disorders are formed

by even a single mismatch (single nucleotide polymorphism [SNP]).

Since detection of these defined sequences of DNA is very important

for the diagnosis of these diseases, development of DNA-based

biosensors is crucial for correct and cost-efficient diagnosis.

Besides the genetic diseases, there are different types of DNA

damages potentially leading diseases like cancer induced endoge-

nously by attack of reactive oxygen species or exogenously by many

different sources such as radiation, ultraviolet light, toxins, and

mutagenic chemicals. Analyzing DNA damage is vital to understand

those diseases and screening new treatments.


Nucleic Acids 501

16.2 Nucleic Acids

Nucleotides consist of a nitrogenous base, a pentose group, and a

phosphate molecule. A molecule that does not contain a phosphate

molecule is named as nucleoside. The nitrogenous bases are

derivatives of pyrimidine and purine (Fig. 16.1).

Nitrogenous base molecules are weak basic compounds, thus

called bases. The base is covalently bound to the 1′C of pentose in

an N-β-glycosyl bond via removal of water and the phosphate is

esterified to 5′C. The purine derivative bases are adenosine (A) and

guanine (G), the pyrimidine derivatives are cytosine (C), thymine

(T), and urasil (U) (Fig. 16.2).

Nucleic acids have two kinds of pentoses: 2′-deoxy-D-ribose

and D-ribose. Both of them are present in their furanose form

in nucleic acids (Fig. 16.3). These sugar residues can easily

bend and twist into different conformations thus making nucleic

acids dynamic in structure. Depending on the kind of pentose,

nucleotides are subclassified into deoxyribonucleotides and ribonu-

cleotides (Fig. 16.4). Ribonucleotides might contain A, G, C, and U

while deoxyribonucleotides contain A, G, C, and T. The bases of DNA

and RNA are important for the structure and e− distribution of

nucleic acids.

The successive nucleotides in deoxyribonucleic acid (DNA) and

ribonucleic acid (RNA) are covalently linked via phosphodiester

bond. In this binding, the 5’OH group of one nucleotide is bridged

to the 3’OH of the next nucleotide (Fig. 16.5).

Figure 16.1. Structures of pyrimidine and purine.



Figure 16.2. Structures of adenosine, guanine, cytosine, thymine, and

urasil.

The backbone of DNA and RNA is hydrophilic due to pentose

group. At pH 7, the phosphate groups of DNA and RNA are ionized

and negatively charged. The purine and pyrimidine bases are

hydrophobic.

There are two important modes of interactions between bases in

nucleic acids:

1. Hydrophobic stacking interactions.

In this type of interaction, the planar and rigid bases are

positioned with the planes of their rings parallel. It involves

combination of van der Waals and dipole–dipole interactions to

minimize water contact with bases thus stabilizing the three-

dimensional structure of nucleic acids.

2. Hydrogen bonds.

This bond involves the amino and carboxyl groups and is

important for complementary association of two nucleic acid


Nucleic Acids 503

Figure 16.3. Structures of pentoses in nucleic acids.

strands (Fig. 16.6). Hydrogen bonds are individually weak;

however, the large number of hydrogen bonds along a nucleic

acid chain provides sufficient stability to hold the two strands

together. While hydrogen atoms of amino group serves as the

hydrogen bond donor, carbonyl oxygen and ring nitrogen serve

as acceptors.

Like proteins, nucleic acids have different modes of structure.

Nucleotide sequence and covalent structure form the primary

structure of nucleic acids. When nucleotides form regular and

stable structures, it is referred as secondary structure. The ternary

structure is considered as the folding of large chromosomes within

the chromatin.



Figure 16.4. Deoxyribonucleotides (dAMP, dGMP, dTMP, dCMP) and

ribonucleotides (AMP, GMP, UMP, CMP).


Nucleic Acids 505

Figure 16.5. Phosphodiester bonds in DNA and RNA.



Figure 16.6. Hydrogen bond.

16.3 Deoxyribonucleic Acid

A DNA (deoxyribonucleic acid) molecule formed by two polynu-

cleotide chains consists of nucleotide subunits. There are four types

of nucleotide subunits of DNA, which are composed of deoxyribose

attached to phosphate groups and four bases: adenosine (A),

cytosine (C), guanine (G), and thymine (T). These nucleotides are

covalently bound to each other via the deoxyribose and phosphates.

One end of the polynucleotide chain has 3’ hydroxyl and the other

end has 5’ phosphate.

The two chains form double helix as two sugar-phosphate

backbones wind around each other as three-dimensional structure.

The bases are stacked inside the helix and their hydrophobic

structures are very close to each other, whereas the hydrophilic


Deoxyribonucleic Acid 507

Figure 16.7. DNA structure. See also Color Insert.

backbones of alternating sugar-phosphate portions are on the

outside (Fig. 16.7) facing the surrounding water. To hold two

polynucleotide chains together hydrogen bonds are formed between

base parts of the polynucleotide chains (Fig. 16.7). During hydrogen

bond formation, A always pairs with T and G with C (Fig. 16.7).

The distance between the vertically stacked bases inside the

double helix is 3.4 A◦. During the formation of dsDNA, ssDNA

strand wind around each other in a way forming two grooves

(minor and major grooves) spiraling around the outside of duplex.

Major and minor grooves create perfect adaptation for the binding

various molecules. Major groove is rich in chemical information. The

secondary repeat distance is about 36 A◦ which is accounted for one

complete turn of double helix happening in every 10.5 base pairs

(Fig. 16.8).



Figure 16.8. Minor and major grooves.

There are three different forms of DNA. B-form DNA also called

Watson-Crick structure is formed by two individual DNA strands

aligned in an antiparallel manner. B-form is the most stable form of

DNA molecule under physiological conditions. It is long, thin, and

right-handed. The number of base pairs per helical turn is 10.5.

B-DNA has wide major groove and narrow minor groove. A-form

DNA is also right handed, but the helix is shorter and wider than

B-form. There are 11 base pairs per each helical turn of A-form DNA.

The major groove of A-DNA is deeper and thus the minor groove is

shallower. The present of A-DNA in cells is uncertain. Alternating

runs of (CG)n·(CG)n or (TG)n·(CA)n dinucleotides in DNA under

superhelical tension or high salt can adopt a left-handed helix called

Z-DNA. In this form, the two DNA strands become wrapped in a

left-handed helix, which is the opposite sense to that of canonical

B-DNA. The number of base pairs per helical turn is 12 in Z-form.

The structure is thinner and longer. While the minor groove of Z-

form is deep, its major groove is hardly apparent. There are some

prokaryotic and eukaryotic examples for Z-form DNA. It has been

suggested that Z-form DNA functions in genetic recombination or

regulation of some genes’ expression.

In addition to the specificity of the hydrogen bonding between

complementary bases, unwounding and rewounding of the double


DNA in Electrochemical DNA-Based Biosensors 509

Figure 16.9. DNA denaturation and renaturation.

helix in exactly the same configuration is one of the most important

features of DNA. Unwinding of the double helix to form two single

strands occurs by disruption of base stacking and hydrogen bonds

between paired bases. During this denaturation process, no covalent

bonds in DNA gets broken (Fig. 16.9). Heat and extreme pH features

can cause denaturation of double-stranded DNA. When temperature

or pH is returned to the physiological range unwound strands

rewind or anneal to yield intact double helix, therefore this seper-

ation of DNA strands is reversible. Each DNA molecule has a charac-

teristic denaturation temperature or melting point (Tm). Since there

are three hydrogen bonds between G and C and two hydrogen bonds

between A and T, separation of paired DNA strands is more difficult

when GC ratio is higher than AT ratio. The transition from double

helix to the single-stranded denaturated form can be detected by

monitoring the absorption of UV light at A260. Denaturation of

double-stranded nucleic acid causes an increase in absorption.

16.4 DNA in Electrochemical DNA-Based Biosensors

DNA-based biosensors are mainly based on hybridization, which

consists of DNA base pairing between two complementary nucleic



acid strands. When mixtures of denaturated DNA from same or

different sources is slowly cooled on same medium, artificial

hybrid DNA molecules to be formed that is called hybridization.

The specificity of such biosensor systems is dependent on probe

selection and hybridization conditions. Such systems are based on

immobilization of the probe (a single-stranded DNA) to recognize its

complementary target strand in a mixture by hybridization. Target

strand is the complementary sequence of the probe. The major

aspect is the electrochemical transduction of DNA hybridization

(Grieshaber et al., 2008; Sassolas et al., 2008; Wang et al., 2008)

(Natsume et al., 2007). There are several approaches to obtain

transduction. Using redox active molecules having the ability to

bind DNA is the most commonly used approach, often referred as

labeled approach. Binding of redox active molecules might occur

via different ways such as intercalating a planar aromatic ring

between base pairs, binding in minor groove, or interaction with one

of the bases. Intercalation happens when intercalating molecules

with appropriate size and chemical properties fit in between

base pairs of DNA. These molecules are mostly planar aromatic

structures. Ethidium bromide and doxorubicin are examples of these

intercalator molecules. These intercalator molecules do not interact

significantly with single-stranded DNA (Sassolas et al., 2008).

On the other hand, changes to the electrical properties of an

interface, change in flexibility from ssDNA to the rigid dsDNA, or

the electrochemical oxidation of guanine bases are used approaches

for label-free methods for detection of DNA hybridization. There

Figure 16.10. Comparison of guanine and inosine structures.


Nucleic Acid Variants Used in Electrochemical DNA-Based Biosensors 511

is a commonly used technique based on the easy oxidisability of

guanine base. Inosine also selectively binds to cytosine bases but its

oxidation signal is different from the guanine peak (see Fig. 16.10 for

comparison of guanine and inosine structures). Substituting inosine

for guanine in the probe strand causes guanine signal loss prior

to hybridization during the process. After the target strand and

probe strand hybridization, the guanine peak from the target DNA

is observed (Sassolas et al., 2008; Wang et al., 2008).

16.5 Nucleic Acid Variants Used in ElectrochemicalDNA-Based Biosensors

In addition to conventional DNA, other variants can be used as probe

DNA. In order to distinguish single-base mutations such as disease

related mutations, having stronger hybridization is essential, and

this can be achieved by using novel oligomers such as PNA and LNA.

16.5.1 Peptide Nucleic Acid (PNA)

Originally synthesized as a DNA-targeting antigene drug, PNA is a

DNA analogue, in which the negatively charged sugar-phosphate

backbone of DNA is replaced with a structurally neutral pseudopep-

tide backbone (Fig. 16.11). This peptide-like backbone is neutral

and consists of repeated N-(2-aminoethyl) glycine units linked by

amide bonds (Nielsen and Egholm, 1999). The purine (A, G) and

pyrimidine (C, T) bases are attached to the backbone through

methylene carbonyl linkages (Nielsen and Egholm, 1999). It has

been used as a novel oligomer due to its ability to hybridize with

single-stranded DNA with high affinity and specificity owing to its

neutral backbone and proper interbase spacing (Fig. 16.11).

The interaction of PNA–DNA is suggested to be more stable

than of DNA–DNA because of higher melting temperatures that is

strongly affected by the presence of imperfect matches (Brandt

and Hoheisel, 2004). Furthermore, there is a larger positive charge

on the hydrogen atoms in the hydrogen bonds of PNA–DNA,

which might explain the greater binding energies for PNA–DNA

double strands than those for the DNA–DNA. Such presence of



Figure 16.11. Protein, PNA structures, and PNA–DNA interaction.


Nucleic Acid Variants Used in Electrochemical DNA-Based Biosensors 513

mismatches in a PNA/DNA duplex is much more destabilizing

than a mismatch in a DNA/DNA duplex, thus a PNA-modified

transducer surface can distinguish perfect complementary DNA

strand from one with a single mismatch. It is suggested that the

uncharged nature of PNA is accounted for greater thermal stability

(Natsume et al., 2007). The lack of electrostatic repulsion between

the two strands in a PNA/nucleic acid duplex leaves the melting

temperature largely independent of salt concentration. In addition

to greater mismatch discrimination, PNA biosensors have higher

biological stability and operation over a wide range of hybridization

conditions (compared to their DNA counterparts). Therefore, PNA

probes become attractive oligonucleotide recognition elements in

biosensor technology (Wang, 1998).

16.5.2 Locked Nucleic Acid (LNA)

LNA, also referred as inaccessible RNA, is an RNA analogue

exhibiting C3’-endo conformation similar to the RNA. In LNA, the

furanose ring of the ribose sugar is chemically locked by the

presence of a methylene linkage between 2′ oxygen and 4′ carbon

of the ribose ring (Fig. 16.12). This linkage locks the pentose in

the 3′-endo conformation found in the A-form DNA and RNA. LNA

has high affinity toward both DNA and RNA single strands. Due to

its restricted conformation, the base stacking and thermal stability

Figure 16.12. LNA.



of LNA are increased. It was shown that the melting temperatures

of LNA–RNA and LNA–DNA double strands are higher than those

of DNA–RNA and DNA–DNA. Therefore, it is expected that LNA

single strand may accomplish stronger hybridization with DNA as

well as RNA. It is used to increase the specificity and sensitivity of

oligonucleotide-based DNA-biosensors (Mukhopadhyay et al., 2005;

Natsume et al., 2007).

References

O. T. Avery, C. M. McLeod, and M. McCarty, Studies on the chemical nature of

the substance-inducing transformation of pneumococcal types, J. Exp.Med. 79, 137–158 (1944).

M. J. Bessman, I. R. Lehman, E. S. Simms, and Arthur Kornberg, Enzymatic

synthesis of deoxyribonucleic acid: II. General properties of the

reaction, J. Biol. Chem. 233, 171–177 (1958).

O. Brandt and J. D. Hoheisel, Peptide nucleic acids on microarrays and other

biosensors, Trends Biotechnol. 22, 617–622 (2004).

J. P. Chambers, B. P. Arulanandam, L. L. Matta, A. Weis, and J. J. Valdes,

Biosensor recognition elements, Curr. Issues Mol. Biol. 10, 1–12 (2008).

R. Dahm, Discovering DNA: Friedrich Miescher and the early years of nucleic

acid research, Human Genetics 122, 565–581 (2008).

K. D. Fairbanks and A. S. Tavill, Liver disease in alpha 1-antitrypsin

deficiency: a review, Am. J. Gastroenterol. 103, 2136–2141, quiz 2142

(2008).

B. Gooptu, U. I. Ekeowa, and D. A. Lomas, Mechanisms of emphysema

in alpha1-antitrypsin deficiency: molecular and cellular insights, Eur.Respir. J. 34, 475–488 (2009).

D. Grieshaber, R. MacKenzie, J. Voros and E. Reimhult, Sensors, 8, 1400–1458

(2008).

S. T. Lommatzsch and R. Aris, Genetics of cystic fibrosis, Semin. Respir. Crit.Care Med. 30, 531–538 (2009).

V. A. E. McKusick, in Mendelian Inheritance in Man Johns Hopkins Univ.

Press, Baltimore, (1991).

R. Mukhopadhyay, M. Lorentzen, J. Kjems, and F. Besenbacher, Nanome-

chanical sensing of DNA sequences using piezoresistive cantilevers,

Langmuir 21, 8400–8408 (2005).


References 515

K. B. Mullis, The unusual origin of the polymerase chain reaction, Sci. Am.262, 56–65 (1990).

T. Natsume, Y. Ishikawa, K. Dedachi, T. Tsukamoto, and N. Kurita, Hybridiza-

tion energies of double strands composed of DNA, RNA, PNA and LNA,

Chem. Phys. Lett. 434, 133–138 (2007).

P. E. Nielsen and M. Egholm, An introduction to peptide nucleic acid, Curr.Issues Mol. Biol. 1, 89–104 (1999).

A. Sassolas, B. D. Leca-Bouvier, and L. J. Blum, DNA biosensors and

microarrays, Chem. Rev. 108, 109–139 (2008).

J. Wang, DNA biosensors based on peptide nucleic acid (PNA) recognition

layers: a review, Biosens. Bioelectron. 13, 757–762 (1998).

Y. Wang, H. Xu, J. Zhang, and G. Li, Electrochemical sensors for clinic analysis,

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structure for deoxyribose nucleic acid, Nature 171, 737–738 (1953).

March 20, 2012 18:32 PSP Book - 9in x 6in Ozsoz-color-Insert

Color Insert

Figure 3.3


C2 Color Insert

Figure 3.4

Figure 3.5


Color Insert C3

Figure 3.6

Figure 3.7


C4 Color Insert

Figure 3.8


Color Insert C5

red

Figure 4.7

Figure 4.13


C6 Color Insert

Figure 4.14

Figure 5.1

Figure 5.2


Color Insert C7

Figure 5.3

Figure 5.4

Figure 5.5


C8 Color Insert

Figure 5.6

Figure 5.7

Figure 5.8


Color Insert C9

Figure 5.9

Figure 6.4

Figure 8.4


C10 Color Insert

Figure 8.9

Figure 8.10

Figure 8.15


Color Insert C11

Figure 8.17

Figure 9.2

Figure 9.3


C12 Color Insert

Figure 9.17

double EVLS

I (μA) EVLS

df(I)

E (mV)

or f(I)or df(I)

Figure 11.2

Figure 11.4


Color Insert C13

Figure 11.6

scan

Figure 11.7


C14 Color Insert

2 nm

minor groove

major groove

B DNA form

3.4 nm(1 pitch)

Figure 12.2

Figure 13.1


Color Insert C15

Figure 15.2


C16 Color Insert

Figure 16.7

ELECTROCH

EMICA

L DN

A BIOSEN

SORS

ELECTROCHEMICAL DNA

B I O S E N S O R S

Edited by

Mehmet Ozsoz

Ozsoz

Mehmet Sengun Ozsoz is a professor of analytical chemistry in the Faculty of Pharmacy at Ege University and also teaches biosensor technology courses in the Biotechnology Department at Izmir Institute of Technology. Prof. Ozsoz holds a BS in chemical engineering from Middle East Technical University, Ankara, Turkey, and a PhD in analytical chemistry from the Faculty of Pharmacy, Ege University, Izmir, Turkey. He was a postdoctoral fellow with Dr Joseph Wang at New Mexico State University, Las Cruces, between 1989–1991 and 1996–1997. He is a recipient of

the 2008 Scientific and Technological Research Council of Turkey (TUBITAK) science award. Prof. Ozsoz conducts well-recognized international work on electrochemical DNA biosensors.

“The marriage of natural and synthetic nanotechnology in electrochemical DNA sensors is

a fascinating object of research. The reader gets an easy access to the complex matter by

the well-written introductory chapter. This volume builds a bridge from molecular biology

to the applications in medical diagnostics and microbiology.”

Prof. Frieder SchellerUniversität Potsdam, Germany

“This book is a very welcome contribution to the literature of electrochemical DNA

biosensors. It offers extremely useful insights into this exciting and important field.”

Dr. Joseph WangUniversity of California, San Diego, USA

This book focuses on the electrochemical applications of DNA in various areas, from basic

principles to the most recent discoveries. It comprises theoretical and experimental analyses

of various properties of nucleic acids, research methods, and some promising applications.

The topics discussed in the book include electrochemical detection of DNA

hybridization based on latex/gold nanoparticles and nanotubes; nanomaterial-

based electrochemical DNA detection; electrochemical detection of microorganism-

based DNA biosensor; gold nanoparticle-based electrochemical DNA biosensors;

electrochemical detection of the aptamer–target interaction; nanoparticle-induced

catalysis for DNA biosensing; basic terms regarding electrochemical DNA (nucleic

acids) biosensors; screen-printed electrodes for electrochemical DNA detection;

application of field-effect transistors to label-free electrical DNA biosensor arrays;

and electrochemical detection of nucleic acids using branched DNA amplifiers.

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