masterarbeit - boku

Masterarbeit

Titel der Masterarbeit

Nanobiotechnology advanced Lab-on-a-Chip for

continuous blood glucose monitoring

Verfasserin

Maria Magdalena Picher

angestrebter akademischer Grad

Diplomingenieurin

(Dipl.Ing)

durchgeführt am Institut für Nanobiotechnologie der Universität für Bodenkultur

in der Abteilung für Nanosystemtechnologie des Austrian Institute of Technology (AIT)

Wien, im Jänner 2010

Studienkennzahl lt. Studienblatt: 006 418

Studiengang lt. Studienblatt: Biotechnologie

Betreuer: Ao.Prof. Dr.Dietmar Pum

Dr. Peter Ertl

Acknowledgment

First of all I want to thank my supervisor Peter Ertl at the AIT for his constant sup-port during the experimental work and also for his valuable advices for the writingof the thesis. He proved to be an excellent supervisor and never stopped having faithin my abilities. Another great thank-you goes to the members of the working groupat the AIT Verena Charwat, Gerald Birnbaumer, Lukas Richter and Michi Putscherfor the pleasant working atmosphere stimulating the work enthusiasm and their sup-portive words in difficult moments. Thanks also to Jakub Dostalek and Chen Huangfor the introduction into SPR measurements and for taking the time for additionalexplanations. Further I want to thank Hubert Brückl, the head of NanoSystems atthe AIT for the appropriation of the technical equipment.Also thanks to my supervisor Dietmar Pum at the University of Life Science Vi-enna to take over the universal supervision and for introducing me into this scientificarea. Further I want to thank Seta Küpcü my second supervisor at the Universityof Life Science Vienna for her great patience with me and her interests concerningthe progress of the work and her permanent availability for any kind of questions.Another great thank-you to Uwe Sleytr’s supportive interest concerning this project.Another great thank-you addresses Marcus Huber who also proved a lot of patiencewhen the work occupied my whole attention and for his support concerning the writingof the thesis. Last but not least I want to thank my mother for her financial supportmaking it possible to study Biotechnology and also for her faith in my person!

i

Zusammenfassung

In den letzten Jahrzehnten hat die Verwendung von miniaturisierten Analysesyste-men, auch Chiplabor, mikro Total Analyse System oder μTAS genannt, an Bedeutunggewonnen. Mit Chiplabor Systemen sind schnelle Analysen von komplexen Probelö-sungen unter der Verwendung geringer Probenvolumina im mikro- bis nanoliter Bere-ich möglich. Im Zuge dieser Arbeit wurde ein Mikro Total Analyse System für diekontinuierliche Messung von Glucose in komplexen Messlösungen entwickelt. DasChiplabor wurde passend für eine mögliche Integration in ein Dialysegerät konzipiertum Blutzuckermessungen vor und nach dem Dialyseprozess durchführen zu können,mit dem Ziel korrekte Glucoseverabreichungen nach der Blutreinigung zu ermöglichen.Derzeit werden standartisierte Dialysate mit konstanten Glucosekonzentrationen ver-wendet, was Symptome von Hypo- oder Hyperglycaemie in Partienten hervorrufenkönnen, die sich in Appetitlosigkeit oder Heißhunger äußern. Der mikrofluidische Chipbesteht aus vier separaten neanderförmigen Reaktoren und electrochemischen Detek-toren aus Platin. Die vier separaten Reaktoren ermöglichen unabhängige Messungenwährend simultaner Analysen und bilden zusätzliche Messkammern für Kalibratio-nen und die Erfassung von auftretendem Rauschen. Um kontinuierliche Messungenvon komplexen Medien ohne Sensitivitätsverlust durchführen zu können wurde indie Messkammern eine S-Schicht integriert. Das S-Schicht Protein SbpA von Lysini-

bacillus sphaericus CCM 2177 wurde an den Wänden der Reaktoren sowie auf denPlatinelektrodenoberflächen kristallisiert und die Eignung als Anti-fouling Oberfächemithilfe von Oberflächenplasmonenresonanz (SPR) nachgewiesen. Die Anti-foulingEigenschaften von dem S-Schicht Protein SbpA zeigten, dass Protein Adsorption anmodifizierten Oberflächen verhindert und auch Adsorption an Elektrodenoberflächenwährend elektrochemischer Messungen merklich reduziert werden konnten (um 96.2%). Das gesamte Chiplabor weist hohe Signalstabilität und Genauigkeit auf und kannfür Glucose Monitoring aus Blutplasma herangezogen werden.

iii

Abstract

The area of miniaturized or microfluidic analysis systems, also called "Lab-on-a-Chip(LOC)", has gained increased popularity over the past decade, fueled by the possi-bility of measurements of small volumes of complex fluids with efficiency and speed,without the need for a skilled operator. Portable LOC devices capable of automatedcomplex diagnostic procedures are able to provide outpatients with important health-related information. For instance, the lack of continuous blood analysis during extracorporal blood purification has shown to lead to an incorrect supply of glucose intothe patient´s bloodstream causing either ravenousness or anorexic symptoms. Toovercome these problems and also to limit patient suffering we have developed a Lab-on-a-chip device with a protein (S-layer) coated microfluidic system that is capableof continuously monitoring glucose concentrations during blood purifications. In thepresented work, a protective protein monolayer (S-layer) is applied to microfluidicchannels as anti-fouling strategy that eliminates adverse sensor surface-sample inter-actions. The Lab-on-a-Chip includes four independently addressable microreactionchambers containing integrated electrochemical detectors intended for internal cali-bration and background subtraction routines. In the presented work we demonstratehow to overcome non specific adsorption of biomolecules at microchannel surfaces us-ing S-layer technology. The protective and non fouling properties of the protein layerwere confirmed using Surface Plasmon Resonance (SPR) in the presence of humanplasma. The S-layer was integrated into the Lab-on-a-Chip system to prevent pro-tein and electrode fouling and to enable continuous electrochemical measurements.A three electrode system was used to continuously monitor glucose concentrationsbased on redox-mediated enzyme reactions. Consequently, the presented work ad-dresses aspects of chip design, fluidic flow profiles, description of anti-fouling layers,sensor characterization and chip application for continuous blood analysis.

Contents

1 Introduction 1

1.1 Aim and structure of the thesis . . . . . . . . . . . . . . . . . . . . . 1

1.2 Monitoring blood glucose concentrations . . . . . . . . . . . . . . . . 2

1.2.1 Glucose sensors . . . . . . . . . . . . . . . . . . . . . . . . . . 4

1.3 Lab-on-a-Chip devices . . . . . . . . . . . . . . . . . . . . . . . . . . 17

1.3.1 Characteristics of LOC systems . . . . . . . . . . . . . . . . . 18

1.3.2 Materials commonly used for microfabrication . . . . . . . . . 25

1.3.3 Microfluidic considerations . . . . . . . . . . . . . . . . . . . . 27

1.4 Surface modification strategies to reduce protein fouling . . . . . . . . 32

1.4.1 Bovine serum albumin (BSA) . . . . . . . . . . . . . . . . . . 33

1.4.2 Poly(ethylene glycol) (PEG) . . . . . . . . . . . . . . . . . . . 35

1.4.3 Surface layer proteins . . . . . . . . . . . . . . . . . . . . . . . 37

2 Materials and Methods 43

2.1 Micro total analysis system (μTAS) . . . . . . . . . . . . . . . . . . . 43

2.1.1 Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 43

2.1.2 Method . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 46

2.2 Surface plasmon resonance (SPR) . . . . . . . . . . . . . . . . . . . . 60

2.2.1 Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 60

2.2.2 Method . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 61

2.3 Standard electrochemistry . . . . . . . . . . . . . . . . . . . . . . . . 66

2.3.1 Materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 66

2.3.2 Method . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 66

3 Results 69

3.1 Development of a glucose monitoring Lab-on-a-Chip . . . . . . . . . . 69

3.1.1 Simulations with fluidic designs . . . . . . . . . . . . . . . . . 75

vii

Contents

3.1.2 Characterization of the microfluidic biochip . . . . . . . . . . 793.2 Characterization of the enzymatic reaction . . . . . . . . . . . . . . . 853.3 Anti-fouling strategies to maintain surfaces . . . . . . . . . . . . . . . 96

3.3.1 Surface modification with S-layer protein monolayers . . . . . 963.3.2 Investigation of non-fouling properties of S-layer proteins . . . 100

3.4 Application of the glucose sensing LOC device . . . . . . . . . . . . . 107

4 Discussion 111

5 Conlusion 115

viii

1 Introduction

1.1 Aim and structure of the thesis

During extracorporal blood purification processes using dialysis devices, inaccurateblood sugar measurements can cause too high or too low glucose injections and re-sults in symptoms of hyper- or hypoglycemia. To overcome patients suffering fromthe symptoms of hyper- or hypoglycemia a sensitive glucose sensor needs to be inte-grated into the dialysis device to advance blood purification treatments.Lab-on-a-Chip (LOC) systems are miniaturized analysis systems, that compete withlaboratory-scale technologies. Analysis of complex biological samples need to bedemonstrated before real world applications of integrated microdevices are estab-lished. The analysis of biological samples translates into several processing stepssuch as sample preparation, analyte enrichment, labeling, signal amplification anddetection and can be performed on a single chip platform. So far, only a few microTotal Analysis Systems (μTAS) capable of delivering results from complex biologicalsamples in a single system have been developed. The aim of this thesis is the devel-opment of a microdevice capable of continuous measurements of blood sugar levelsusing a microfluidic device. Therefore an online glucose monitoring Lab-on-a-Chip de-vice with continuous feed line for continuous blood sugar measurements of untreatedwhole blood samples was developed. Further requirements on the LOC device are theintegration of a simultaneous calibration system and also the possibility of concurrentnoise reduction measurements that enable fast and sensitive glucose monitoring.The structure of the thesis includes:

An Introduction describing existing glucose monitoring systems and glucose detec-tion, Lab-on-a-Chip devices and their application field and nonfouling strategies forcontinuous electrochemical sensors.

1

1 Introduction

A Materials and Methods section outlining the configuration of the measurementstation and a short description of measurement principles of used methods such asAtomic Force Microscopy and Surface Plasmon Resonance.

The Result section presents results of all measurements from the electrochemicalcharacterization of the chip over the examination of surface modifications and theirinfluences on protein adsorption using Surface Plasmon Resonance to the continuousglucose measurements on chip.

The thesis ends with a discussion and possible future applications.

1.2 Monitoring blood glucose concentrations

Possible application area of a glucose monitoring LOC

"Dialysis is called a process of cleaning the blood by passing it through a specialmachine and is necessary when the kidneys are not able to filter the blood. Dialysisallows patients with kidney failure a chance to live productive lives." [1] The purifica-tion mechanism of dialysis devices is based on the diffusion of solutes including ureaand potassium or calcium excesses across a semi-permeable membrane (dialyzer). Theblood flows by one side of the semi-permeable membrane and the dialysate, a specialdialysis fluid flows by the opposite side. The dialysate contains equal concentrationsof blood components, that are not required to be removed and the final concentrationsof components like calcium and potassium normally highly concentrated in the bloodof patients suffering of kidney failure. The semi-permeable membrane is permeablefor small molecules like vitamins, glucose, ions and urea but blocks the passage oflarger molecules, such as cells and large proteins. The circulation process of the bloodoutside the body has to proceed until the waste and excess water in the blood areremoved or average a desired concentration. (Hämodialyse 100L). The amount ofglucose in the prepared dialysate is predefined and averages 1 g/L to prevent energylosses and hypoglycaemic symptoms of patients with diabetes during the procedure[2]. For the supply of representative concentrations of glucose similar to the concen-trations in the patient´s whole blood, a continuous glucose sensor can be applied tothe dialysis device to prevent sufferings from hyper- or hypoglycaemic symptoms, asshown in Figure 1.1.

2


Figure 1.1: Dialysis device and possible position of a Lab-on-a-Chip device

The LOC device has to be placed near the tubing system beneath the filtration mem-brane, so that blood coming from the body and feeding back into the body can beanalyzed simultaneously. A separate chamber for simultaneous calibration and back-ground measurements accelerates the measurement and increases the sensitivity.

3

1 Introduction

1.2.1 Glucose sensors

1.2.1.1 Glucose oxidase, the enzyme for many biosensor applications

Glucose oxidase, a flavin adenine dinucleotide (FAD) dependent enzyme, is in widespreaduse for biosensors, mainly for the measurement of blood glucose levels. Glucose oxi-dase is very often used caused by many advantageous properties such as a high speci-ficity, a high turnover and a high stability. The enzyme glucose 1-oxidase (GOD) wasintroduced by Müller in 1928 [3].

Structure

Glucose 1-oxidase is a dimer composed of two identical subunits, that are linkedvia disulphide bridges. 16 % of the secondary structure consists of α helices and 8to 12 % form a branched polysaccharide that partly surrounds a protein core. Thecarbohydrates do not influence the enzymatic reaction but increase the stability ofGOD by enhancing the resistance to proteases and increasing the solubility in water,inducing the enzyme a resistance to precipitation by e.g. trichloroacetic acid. The 3dimensional structure of glucose oxidase is shown in Figure 1.2.

Figure 1.2: 3 dimensional structure of glucose oxidase. The yellow cross marks the glucose

binding side [4].

4


Chemical and biological properties

Glucose oxidase is known for its remarkable stability. In lyophilized state at 0 ◦CGOD is stable for 2 years and in solutions the stability strongly depends on the pHvalue, where ranges between 2 and 8 are recommended. If the pH of the solutionoutstays the recommended range, the activity gets significantly lost, in case of a pHof 8.1 only 10 % of the activity remains after 10 min [5]. As already mentioned theenzyme is very resistant to proteolysis and also non-ionic detergents have little ef-fect on it. Problematic are ionic detergents like sodium-dodecylsulfate (SDS). Stronginhibitors of GOD are heavy metals like mercury, lead and silver, where even mi-cromolar amounts induce inhibitory actions. Competitive inhibitors are for instancealdohexoses like D-arabinose and 2-deoxy-D-glucose, both resulting in a decline ofenzyme activity [6].As can be obtained from Figure 1.2 glucose oxidase binds two molecules of the co-factor flavin adenine dinucleotide (FAD) that are tightly but not covalently bound tothe enzyme and do not cause denaturation when being removed.

Reaction mechanism

In general GOD catalyzes the reaction involving the oxidation of D-glucose with O2,producing D-gluconolactone and hydrogen peroxide (H2O2). D-gluconolactone canrapidly hydrolyze to gluconic acid, with a reaction velocity highly dependent on thepH value, and increasing reaction velocity directly proportional to the pH. The exactreaction mechanism can be separated into seven consecutive steps. Initially the en-zyme binds the substrate glucose and forms an enzyme-substrate complex, which isan extremely rapid reaction process (I). What follows is the reduction (II) of FAD,where the enzyme attacks the C1 hydrogen atom of glucose and starts the reductionreaction by transferring the hydride ions to the FAD. Another possible mechanismof this reaction is the induction a glycosidic bonding between glucose and FAD anda following abstraction of a proton from the C1 hydrogen atom of glucose, that isafterwards transfered to the FAD [7]. The reduced form of glucose oxidase (III) is inthis state unable to react with O2 and needs glucose to be converted into a reducedreactive state (IV), that can be rapidly oxidized by O2 (V). The oxidation reactionends with the release of the protonated H2O2 (VI). The enzyme is now inactive andunable to react with glucose and has to undergo a conformation change before start-ing a new oxidation reaction (VII) [8]. The chronological flowchart of the enzyme

5

1 Introduction

reaction is demonstrated in Figure 1.3

Figure 1.3: Flowchart of the enzymatic reaction between the enzyme glucose oxidase and

the substrate glucose

Mediating elements

In earlier glucose sensing devices, the removal of O2 was detected electrochemicallywith oxygen electrodes such as the famous Clark electrode. Another measurementprinciple is the measurement of H2O2 production. In this work another an electronacceptor acting as oxidizing substrate was used to detect glucose concentrations. De-pendent on the pH optimum various electron acceptors can be used, resulting in 4main mediator groups:

6


Table 1.1: 4 categories of mediating elements

pH optimum mediator

5.6 Quinones, Dioxygen7.5 Ferrocene, Phenoxazines, Tetrathiafulvalene,Benzylviologen< 3 Ferricyanide, Tris(2,2´-bipyridine)cobalt(III)perchlorate< 4 Indophenols

The catalyzed oxidation reactions of glucose oxidase including a specific mediatorare described by the following reaction-equations:

Eox +Gluc.ox−→red→ gluconolactone+ Ered (1.1)

Ered +Medox −→ Eox +Medred (1.2)

The possible usage of various mediating elements offering multiple chemical charac-teristics adds glucose oxidase further advantages for glucose detection in a varietyof measurement solutions. For further information consult the following reviews orarticles [8, 9, 5].

1.2.1.2 Glucose detection methods

Glucose concentrations can be detected in many ways, directly via optical detectionmethods or indirectly by using an enzyme to oxidize glucose and measure the reactionproduct. In Table 1.2 common detection methods are listed.

Table 1.2: Glucose detection methods

Method Examples Refs.

Optical detection Infrared spectroscopy (IR)Fluorescence measurementPolarimetryRaman spectroscopy

[10, 11,12]

Electrochemical detection Impedance SpectroscopyChronoamperometryCyclic voltammetry

[13, 14,15]

7

1 Introduction

In this work glucose concentrations were measured using the electrochemical mea-surement methods Cyclic voltammetry (CV) and Chronoamperometry (CA) and willbe explained in more detail.

Cyclic voltammetry

Cyclic voltammetry (CV) is an electrochemical method and can be employed tostudy the electron transfer kinetics and the transport properties of electroactive com-pounds. CV measurements are performed under static conditions and are character-ized as potential step methods. While the voltage is adjusted between working andreference electrode, the current is recorded between working and counter electrode.The scan rate affects the amount of the reduced or oxidized species, caused by thediffusion controlled measurement conditions. The scan rate v can be defined as thefollowing:

v = ΔE/Δt (1.3)

ΔE denotes the potential difference between upper and lower limit and Δt the re-quired time. Beginning at a start potential Ei the voltage is increased at a determinedscan rate until the potential maximum Emax is reached and proceeds with the back-wards scan using the same scan rate in direction to the lower potential limit Emin.The chronological sequence can be obtained from Figure 1.4.

ipC

ipa

EpaEpc

Figure 1.4: Cyclic voltammograms: a) Delta voltage plot plotting the voltage as a function of

time, with the potential max. V1 and the minimum V2. b) Cyclic voltammogram.

8


The scan range has to be adjusted to the standard electrode potential (EΘ) andhas to be set on a higher level, to enable a complete oxidation and reduction of theredox system. The relationship between concentration and voltage can be describedby the Nernst equation:

E = EΘ +RT

nFln

[cox]

[cred](1.4)

where E denotes the applied potential and cox/red the concentration of the oxidizedand reduced form of the analyte, R denotes the ideal gas constant 8.314 J/molK, Tthe temperature, n the amount of transferred electrons and F the Faraday constantof 96 485 C/mol. When the measurement solution contains only one electrochemicalreactant, the graph of the CV measurement looks like shown in Figure 1.4. The cyclicvoltammogram shows a potential maximum at 500 mV and a potential minimum at-300 mV . Ea

p denotes the anodic peak potential and Ecp the cathodic peak potential.

The curve progression in direction of higher current describes the oxidation reactionof the electroactive component and in the opposite direction the reduction reaction.The cathodic and anodic peak currents (ic and ia) can be taken from this graph bycalculating the current of the distance between the peak maximum and the linearelongation of the reaction start (marked with the red lines). If the electrode potentialexceeds the standard potential of a reduced component, a positive oxidative anodiccurrent (ia) starts to flow until the maximum is reached. The event on the electrodesurface can be describe by the following redox reaction:

cred − e− → cox (1.5)

The voltage at iamax is defined as anodic peak potential. The voltage separationbetween the two peak potentials is defined as:

ΔE = Eap − Ec

p =59

n[mV ] (1.6)

where Eap denotes the anodic peak potential, Ec

p the cathodic peak potential and nthe number of transfered electrons of the redox couple. By changing the scan rate,the peak voltages are not altered, only the peak high. Another characteristic of CVis that the ratio of the peak currents should equal 1 and that the peak currents ia

and ic are proportional to the square root of the scan rate:

iap and icp ∝√v (1.7)

Cyclic voltammetry is a very popular technique to study the kinetic behavior of acertain redox species or enzyme reactions of flavoenzymes such as oxidoreductases in

9

1 Introduction

solution. Another possibility is the examination of the surface behavior of certainelectrodes since surface charges and surface areas can be indirectly analyzed [16, 17].

Chronoamperometry

Chronoamperometry (CA) is an electrochemical technique recording current as afunction of time, while the voltage remains constant. CA is a often used analyticalmethod to determine the concentration of a certain electroactive species in test solu-tions. By initiating a voltage faradic and non faradic processes occur on the electrodesurface, resulting in rapidly changing currents at the beginning decreasing exponen-tially, converging against a minimal current, as can be obtained from Figure 1.5.If the applied voltage exceeds the standard potential of an electroactive component

Figure 1.5: Current-time traces of CA measurements. In the first period of the measurement

both, faradic and non-faradic reaction are detected and the peak exponentially

decreases. When the curve roughly remains at a constant level, faradic reaction

can be observed.

in a test solution, an oxidation reaction occurs on the electrode surface, resulting inthe faradic reaction defined by the electron transfer between the component in the so-lution and the working electrode. Positive currents describe oxidation reactions, whilereduction reactions induce negative currents. Oxidation and reduction processes are

10


described by the Cottrell equation:

i =nFAcOj

√Dj√

πt(1.8)

where i describes the current in A, n the number of transfered electrons, F the Faradayconstant (96 485 C/mol), A the surface area of the planar electrode, cOj the initialconcentration of the electroactive species in mol/cm3, Dj the diffusion constant ofthe species j in cm2/s and t the time in s.Chronoamperometry is a diffusion controlled electrochemical measurement, as can beobtained from the Cottrell equation. The current depends on the diffusion rate ofthe analyte to the electrode surface. The higher the applied voltage, the wider thediffusion layer and the higher concentrations of the analyte can be oxidized/reduced[16, 17].

1.2.1.3 Types of glucose sensors

Glucose sensing devices became an important issue for the treatment of type 1 di-abetes. With the increased number of accurate glucose sensors, it became easier tocontrol the blood sugar levels and maintain euglycaemia. The alternation between hy-poglycaemia and intermittent hyperglycaemia, the resulting conditions after too lowor high blood sugar levels of patients suffering type 1 diabetes became controllable.The first patent on a glucose sensing device was filed in 1971, the Ames ReflectanceMeter, a device that analyzed the color change of enzyme-based reagent strips. Thefirst glucose sensor suitable for blood-glucose monitoring at home was the Ames Eye-tone Meter investigated by Richard Bernstein, a US physician with type 1 diabetes,who wrote a guide to achieve normal blood sugars. After the development of the AmesEyetone Meter the research of glucose sensing devices made substantial progress [18].The wishlist of a today´s ideal glucose sensor contains the following properties: Thesensor has to be selective for glucose only, fast with predictable responses to glucosechanges and reversible with reproducible signal outputs. The fabrication should befast, reproducible and cheap on large scales. The operational lifetime under physio-logical conditions should be long and the acceptance from patients high.The trend is towards continuous glucose monitoring, non invasive techniques offeringcontinuous real-time information regarding glucose levels, to increase the understand-ing of fluctuation sources and eventually reasons for hyperglycaemie [18].

11

1 Introduction

In the following section the measurement principles of a few glucose sensors listed inFigure 1.6 will be described.

Figure 1.6: Schematic list of glucose sensing technologies already available or in development

phase [18]. MIR: mid infrared; NIR near infrared;

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Most of non invasive glucose sensors measure the glucose content of the interstitialfluid, abbreviated ISF. ISF is the term of the fluid surrounding the cells of a multicel-lular organism, or the extracellular fluid, except blood. The glucose concentration inthe ISF is dependent on blood sugar levels and metabolic rates. The rise of glucoseconcentrations in the ISF lags behind the concentration in the blood between 2 to45 min, with a mean value of 6.7 min. If the blood sugar level decreases, the con-centration starts to decrease in the ISF [19]. The lag time between blood and ISFglucose depends on the species, the sensor size, the applied stimulus and the depthof the tissue.

Non invasive glucose monitoring

Optical transducersInfrared absorption spectroscopy:Mid-infrared (MIR) light with a wavelength of 2.5 - 50 μm is launched to the bloodsample and the absorption of light is measured (intensity of the light-beam measuredbefore and after interaction with the matter). For glucose monitoring applicationsnear infrared (NIR) spectroscopy (0.7 - 1.4 μm) provides better performance, because90 - 95 % of the light passes the epidermis into the subcutaneous space, independentof pigmentation. The oxygenated and deoxygenated haemoglobin is monitored. Thespecific absorption depends on molecular structures and glucose absorption. To mea-sure one analyte only, multivariate analysis with calibrations are necessary to deriveglucose values. Near infrared spectroscopy is a fast reagentless method, that canbe used for in vivo measurements. Problematic is the high background signal in-duced by other molecules such as water and the absorption of light and scatteringappearances caused by fat contents of patients. The fat tissue variation between pa-tients is a reasons for heterogenous measurements and frequently needed calibrations.Candidates of sensing regions are the mucosa or measurements across the tongue [18].

Thermal infraredThermal infrared is based on the phenomenon, that cutaneous microcirculation isdependent on the local glucose concentration. By inducing periodic temperaturevariations in the skin, assessing MIR light scattering at different tissue depths, theglucose concentration can be derived from the degree of scattering after a foregone

13

1 Introduction

calibration process [18].

Photoacoustic spectroscopyPhotoacoustic measurements are based on the measurement of ultrasonic waves emit-ted by tissue after the absorption of light. After launching a light beam with a definedwavelength into a certain tissue, a part of the waves gets absorbed resulting in a lo-calized heating effect. The temperature of the irradiated tissue is dependent on thespecific heat capacity, which is furthermore dependent on the glucose concentration.After heating up the tissue, it starts to expand and therefore generates an ultrasonicpulse that can be detected. If the heat capacity is lower, the pulse frequency is higher.Advantages of photoacoustic spectroscopy are the absence of interferences by NaCl,albumin and cholesterol and a better detection sensitivity than with other opticaltransmission techniques [18].

Transdermal sensorsReverse iontophoresisIn reverse iontophoresis an electric current is applied between two electrodes acrossthe skin. The generated electric field induces active flow of charged and also un-charged molecules across the dermis, including ion moving to maintain the skin´sneutrality. As the skin is negatively charged under physiological pH, cations, mainlysodium molecules, are able to move. Iontophoresis induces a flow of glucose from thecathode to the anode, where it is collected. The concentration of glucose in the fluidcollected by iontophoresis is 1000 fold lower than in blood and can be measured by anenzyme based electrochemical measurements. An advantage of reverse iontophoresisis, that the collected fluid is free from larger molecules and prevents electrode fouling.Glucose is measured indirectly by a subsequent sensing device [18].

Skin suction blister techniqueSkin suction blister is a frequently used method to determine blood sugar levels. Vac-uum is applied to a small spot on the skin, inducing the formation of blisters witha diameter of a few millimeters. The fluid collected of the blister can be analyzed,because the composition is similar to serum, with a lower protein concentration. TheSkin suction blister technique is like reverse iontophoresis dependent on a subsequentglucose sensing device. In general skin suction blister is well tolerated among patients,

14


painless, with a low possibility of infection with a glucose concentration in the ISFcorrelating well with the glucose concentration in the blood [18].

Minimally invasive glucose monitoring

Micropore and MicroneedleMicropores can be generated by pulsed laser beams or the local application of heat.The pores penetrate the stratum corneum and not the full skin. ISF can be collectedvia vacuum pump and afterwards analyzed to derive the blood sugar level.Microneedles can be used to sample capillary blood. The needle itself has a diameterof 17.5 μm and samples the the fluid by capillary forces. The needles are almostsensation less and the sample can be collected and analyzed by enzyme based sensorsystems. The correlation of glucose measurements in the collected ISF and referencesensors are well tolerated. Disadvantageous are the blister formations over the vac-uum site [18].

Invasive glucose sensors

Subcutaneous needle-type sensorsSubcutaneous needle-type sensors are used to measure blood sugar levels directly fromblood samples by using subsequent enzyme based sensing devices.Enzyme based electrochemical sensors use the ability of enzymes as glucose oxidase(GOD) to oxidize glucose and transfer of the electrons to acceptor molecules andfinally to the electrode. The electron transfer is detected via amperometric or volta-metric sensing methods, resulting in a concentration dependent current or voltage.The enzyme GOD transfers the electrons to H2O producing H2O2 or to an appliedmediator. The advantage of mediator based enzyme sensors is the independenceon oxygen concentrations, the mediator regeneration at low potentials, resulting innegligible or no interferences with other blood components beside the absence of in-terferences by urate, ascorbate and paracetamol [18]. The mechanism of enzymebased electrochemical glucose measurement was used as measurement principle of theLOC devices developed in this work and is therefore additionally shown in Figure 1.7.

15

1 Introduction

glucose

glucose oxidase

glucose

glucose oxidase

glucose oxidase

2xFAD+

FAD+

FADH2

gluconolactone

glucose oxidaseFADH2

2x

2x

1) 2)

3)4)

........ mediator [K3Fe(CN)6]

........ enzyme [GOD]

........ FAD+

........ FADH2

........ substrate / product[glucose / gluconolactone]

gluconolactone

Figure 1.7: Description of the expiration of a mediated glucosesensor: a) Components nec-

essary for the enzyme reaction, b) Binding-event of the substrate to the enzyme

and oxidation reaction of glucose. The electrons are transported through an

electron-channel to the FAD and induce the reduction reaction of the enzyme.

c) Transfer of the electrons to the mediator , causing a reduction of the mediator;

d) Diffusion of the mediator to the electrode, where the voltage has to be high

enough to induce the oxidation of the mediator

The adduced glucose sensors only provide a brief insight into the broad field ofglucose monitoring to get an impression of the diversity of measurement methods.For further information consult [18].

16

1.3 Lab-on-a-Chip devices


In recent years Lab-on-a-Chip devices gained popularity, because they mimic labora-tory conditions on a miniaturized scale and provide controlled conditions for scientificmeasurements, see concept in Figure 1.8.

Modern microfluidic devices can be

Figure 1.8: Concept of a LOC device

traced back to the year 1979, to the de-velopment of a microfluidic based on gaschromatography on a 5 cm silicon wafer[20]. In the same year IBM published afabrication process for planar silicon ink-jet printheads that became one of themost successful commercially availablefluidic applications [21]. Both techni-cal expertises highlight the day of birthfor Lab-on-a-Chip, abbreviated LOC de-vices.Closely related fields to LOC systems are

micro Total Analysis Systems, also termed μTAS and micro Electro- Mechanical Sys-tems, also called MEMS. The former field focuses primarily on the integration andminiaturization of devices for analytical chemistry and the latter term describes sys-tems with integrated electrical elements. LOC, μTAS and MEMS are often overlap-ping technologies, sharing the same elements. The newer generation of microfluidicsstarted in 1990 with the elaboration of miniaturized total chemical analysis systemsas new concept of chemical sensing [22]. In this theoretical work, A.Manz, N. Graberand H.M. Widmer illuminated the conversion of Total Analysis Systems (TAS) intoμTAS. By analyzing the theory behind hydrodynamics and diffusion, the results indi-cated a faster and more efficient transport and electrophoretic separation time. Theyalso mentioned a decreased consumption of reagents and a possibility to increasesensor performance. Since then the investigation on micro Total Analysis Systemsproceeded in many different areas and found a broad range of application possibilitiessuch as medical and molecular diagnostics [23, 24], biological and chemical analysis[25, 26, 27], clinical and forensic analysis [28, 29] and pathogen detection [30].

17

1 Introduction

1.3.1 Characteristics of LOC systems

Microfluidics involve technical, economical and eventually also ecological advantages.The most important characteristics are cited in the following list:

• small scaling factor

• highly laminar flow conditions

• requirement of small amounts of reactants and samples

• various kinds of fluid transport systems

• miniaturization of analytical equipment

Ad small scaling factor

By scaling down the analytical tool, the surface to volume ratio is drastically in-creased, resulting in high surface interactions with the measurement solution, alsocausing interactions between the surface and components of the analyte. Concerningheat conductions the small scaling factor enhances the achievement of equilibriumconditions.The high surface to volume ratio results in advantages concerning enzyme based re-actions on chip. By immobilizing the enzyme on the surface of the microchannel ahigh amount of active enzymes is able to react with a sample solution from all fourchannel-walls, resulting in a high contact surface for enzymatic reactions. The surfacecoverage of the enzyme on the high surface area increases the reaction rate while thethroughput of the enzyme is reduced, resulting in a cost reduction of the measurement.

Ad highly laminar flow conditions

Small geometries yield Reynolds numbers (Re) < 1. Low Re numbers indicatestrictly laminar conditions in micro channels resulting in mixing effects induced onlyby diffusion. Increased turbulences can be achieved by specific structures such assharp edges. The highly laminar flow conditions also cause precisely controlled flowconditions within the microfluidic device and can be adjusted to desired require-ments by specialized geometric designs and technical tricks like cell arrangement by

18


electrophoresis or the changing of the microchannel width to enable variable flow ve-locities for specific reactor chronologies [31].

Ad requirement of small amounts of reactants and samples

The requirement of small amounts of sample volumes accompanies a reduced con-sumption of analytical compounds, leading to significant cost reductions for analysisdealing with expensive reactants. For parallel applications, necessary in case of dif-ferent kinds of arrays small sample volumes are highly recommended [32].

Ad various kinds of fluid transport systems

Capillary forces can be used for fluid transport systems. Also continuous flow anal-ysis seem to be a minor challenge, since continuous flow injection devices have beenestablished [33].

Ad miniaturization of analytical equipment

The miniaturization of total analysis systems (TAS) into small LOC devices enablesthe integration of μTAS into technical equipment. An example for an integration pos-sibility is the utilization of a LOC-glucose sensing device into a blood dialysis devicefor sensitive blood sugar measurements during blood purification processes [34, 32] .

Application fields for microfluidic devices are systems for DNA analysis, devices forseparation based detections, devices for cell handling, sorting and general analysis,systems for protein based analysis, immunoassays and also for chemical analysis,detection and processing. In this work only a few microfluidic devices of differentapplication fields are mentioned.

19

1 Introduction

Table 1.3: Lab-on-a-Chip applications

Example Description Refs.

Polymerase chain reactions(PCR)

control channel (grey neanders)fluidic channel (black)heating plates (light grey aroundchannel)

The rotary LOC device consists of three lay-ers, the top layer with the Polydimethylsilox-ane (PDMS) control channels, a thin PDMSmiddle layer containing the fluid channel andthe heating element as bottom layer. Theloop structure of the fluid channel can besplit into two semi-cycles with different chan-nel width to vary the fluid velocity in bothsections. This is necessary to enable dis-tinct residence times of the sample at specificstages of the PCR cycle.

[35]

Electrophoresis A Lab-on-a-Chip device with fully integratedelectrochemical detection methods for cap-illary electrophoresis was established. TheμTAS consists of capillary electrophoresischannels and high voltage electrodes made ofplatinum specially designed for electrochem-ical measurements. The detection limit forcatechol and dopamin, two electroactive sub-stances is in the range of 4 - 5 μM and the an-alytical performance had been obtained overmonths of usage.

[36]

20


Table 1.4: Lab-on-a-Chip applications part 2


Micro coulter particle counter(μCPC)

The μCPC can be used for diagnostic appli-cations, mainly cell counting and separation,by measuring the special impedance of sin-gle cells or even particles. In a subsequentseparation step, cell sorting is a possiblemechanism, whereby the first on chip flow-cytometer had been published. The μCPCdevice can be used for counting, sizing andpopulation studies.

[37]

Protein synthesis This LOC system consists of a microreactorwith integrated temperature control chip andreaction chamber. The temperature control-ling subunit is placed beneath the reactionchamber to control the reaction temperatureduring the cell free protein synthesis. Thesmall reactor is ideal for fast controls overreaction-temperatures. The performance ofthe synthesis chip was evaluated by synthe-sis of green fluorescent protein (GFP) andblue fluorescent protein (BFP) and subse-quent detection of fluorescence emission ofthe products.

[38]

21

1 Introduction

Table 1.5: Lab-on-a-Chip applications part3


Sandwich immunoassay The aim of the cited work was the detectionof D-Dimer (DDi), the final product duringblood coagulation. The enzyme horseradishperoxidase was linked to the secondary anti-body to detect DDi. The detection limit wasin the range of 0.1 to 100 nM .

[39]

Glucose and lactate sensing de-vice

The sensing device consists of a reactorwith dual enzyme-modified microelectrodesfor the monitoring of glucose and lactate.To enable the lactate and glucose measure-ment one carbon electrode was modified witha layer of glucose oxidase and another elec-trode with a layer of lactate oxidase. Toprevent electrode cross talk during measure-ments, a flow separator was integrated intothe LOC system. The interfering effect ofL-Ascorbic acid was eliminated by incorpo-rating L-ascorbate oxidase upstream of thesensing electrodes, which lead to the remark-able result that both lactate and glucose weredetectable in the range between 5 μM and 5mM , which covers the concentration rangesufficient for brain glucose and lactate mea-surements.

[40]

22


1.3.1.1 LOC devices for glucose monitoring

Lab-on-a-Chip devices for glucose monitoring applications provide a heterogeneousresearch area, with various on chip sample preparation techniques and creative ideasfor performing multiple analysis simultaneously. Multiple enzymatic analysis can beobtained of analysis of enzymatic reactions with different reaction products and adetection system, capable of distinguishing between products. Examples for glucosemonitoring LOC devices and techniques for multiple enzymatic analysis are listed inTable 1.6.

Table 1.6: Glucose monitoring μTAS

Example Principle Refs

Enzyme-Release Capillary Capillary assembled microchips arereagent-release capillaries with non-covalent immobilized enzymes insidethe square glass capillary and are capa-ble of measuring different analytes byintegrating various biosensors.

[41]

Biosensor array microsystem The microsystem consists of two lay-ers, the upper part with the integratedbiosensor array and the lower part con-taining the gold counter electrodes andelectrical interconnection lines. Bymodifying electrochemical transducersand utilizing photopatternable enzymemembranes crosstalk free, long termmeasurements were performed.

[28]

23

1 Introduction

Table 1.7: Glucose monitoring μTAS part 2

Example Principle Refs

Multi-enzyme microchipfor simultaneous detectionof sugars

The microfluidic chip is capable ofquantifying sucrose, D-glucose and D-fructose simultaneously using an inte-grated optical detection system. Theoptical detector measures the changeof NADH concentrations, the reaction-product of the enzymatic reactions.The chip consists of a long reactionchannel, where the three sugars areconsecutively detected.

[26]

Microseparation chip for simultaneousmeasurements of ethanol and glucose

In the analytical platform a capillaryelectrophoresis microseparation tech-nique is used to enable multiple enzymereactions. Glucose and ethanol con-centrations are measured using the en-zymes glucose oxidase and alcohol de-hydrogenase. The reaction productshydrogen peroxide and NADH are sep-arated and detected via amperometricthick-film detectors.

[29]

24


1.3.2 Materials commonly used for microfabrication

Table 1.8: Materials for microfabrication

Application Substrate Refs.

Micromachining SiliconGlassPolymer

[32]

Microfluidics PDMSSU-8PMMA

[32]

PDMS

Microfluidics are devices with integrated channel networks of various dimensions,mainly between 5 - 500 μm, and are suitable for transports of small fluid volumes inthe range of microliters to femtoliters. Microfluidics are mostly made of polymers,although glass is also sometimes used. Poly(dimethylsiloxane) abbreviated PDMS isan attractive material for chip developments, because it is an excellent for easy andrapid productions and replications. PDMS is elastic, therefore hard to break, and isalso relatively cheap. Micro Total analysis Systems often contain microfluidic parts,often placed between other materials like e.g. glass and need to be covalently bondedto seal the channel network. For sealing purposes adhesives are used for PMMA orother polymer fluidics. PDMS can be covalently bonded after activating free reactivegroups using air or oxygen plasma (as described in Section 2.1.2.4). After plasmatreatment the oxidized and therefore reactive Si − (OH) groups can be sealed tothemselves and other materials, by forming covalent bondings. This feature enablesthe fabrication of complex 3D structures and microchannel networks by multilayerPDMS prototyping approaches[42].A disadvantage of PDMS is its penchant for absorbing small, hydrophobic, biologicaland drug molecules from solutions, a characteristic influencing measurements of drugscreening microfluidic devices [43].The reasons why PDMS is used in this work are mentioned above, the possibilityof fast chip fabrication that accelerates the microfluidic developing process, the easyhandling with PDMS, the possibility to fabricate complex microfluidic geometries like

25

1 Introduction

rectangular structures without additional problems and also the binding capacities ofPDMS microfluidics resulting in a seal microfluidic device. For further details consult[44, 25].

SU8

The epoxy-based UV-sensitive polymer SU-8 provides an alternative to PDMS.The hard polymer SU8 presents a thermally and mechanically stable material. SU-8 is impermeable for small molecules, provides high biocompatibility and chemicalinertness against most biological substances. To seal the SU-8 microfluidic betweentwo glass plates the soft baked fluidic can be bonded to the glass substrate usingone thin layer of liquid SU-8. The epoxy groups on the surface can be changed intohydroxyl groups by treatment with strong acids and can be further used for covalentbinding of suitable biomolecules like enzymes [45, 46]. Additionally the bondingcan be accomplished by applying adhesives and an additional soft baking step. Asa favorable alternative it is also possible to use soft SU-8 spinned onto soft bakedmicrofluidic. The additional layer provides a binding substance to other materialssuch as glass and hardens in the correct assembly by a shared hard baking step. Inthe course of this project SU-8 was used for first microfluidic fabrications. For furtherinformations about microfluidic chips made of SU-8 consult [47, 48].

26


1.3.3 Microfluidic considerations

In microfluidic devices a variety of driving force for fluid transport were investigatedand are in use, as can be obtained from Table 1.9.

Table 1.9: Fluid dynamics in LOC systems

Force Description Refs.

Pressure driven flow The flow rate can be determined byinducing pressure from outside, usingpumps such as syringe pumps or mi-cropumps.

[49, 50]

Capillary/ Evaporation effects The driving force is a pressure differ-ential arising from a meniscus near theoutlet reservoir.

[51]

Electrophoresis With the induction of a magnetic field,charged molecules start to move. Theflow velocity and the flow direction aredependent on the charge, form and sizeof the charged components beside theinduced electric field. Electrophoresiscan also be used to orientate cells ormolecules.

[31]

Electromagnetic forces By inducing a magnetic field to a per-meable core a magnetic flux is inducedand the magnetic resistance reduced.The magnetic force minimizes reluc-tance of a magnetic system to enablefluid transport.

[52]

Electroosmosis Electroosmosis defines a electrokineti-cal driven flow with induced charge os-mosis. The charge osmosis can also beused to enhance micromixing.

[53, 54]

27

1 Introduction

In this work fluid flow was afforded by pressure forces using a syringe pump. Forthat reason the microfluidic fundamentals in this work focus on pressure driven flow.In fluid dynamics, the law of conservation of mass is formulated by the continuityequation:

ρ1A1ν1 = ρ2A2ν2 (1.9)

where ρ describes the density, A the area of a cross section and ν the fluid velocity.The continuity equation testifies that mass can not get lost and therefore the mass ofthe incoming fluid has to equal the mass of the out coming fluid, as shown in Figure1.9.

Figure 1.9 demonstrates the flow behavior in a tube with altering diameter. At

Figure 1.9: Behavior of a fluid in a channel with decreasing diameter. The red color indicates

high flow velocity and blue rather low flow velocities. By decreasing the diameter

of the tube, the velocity (ν2) increases to transport the same fluid-volume in a

specific fraction of time compared to the wider channel.

position A1 the amount of fluid passing in a defined time interval remains the samecompared to position A2, so the fluid velocity increases. If A1 has an area of twice A2

the mean velocity of v2 is twice as fast as v1. Microfluidics consist per definition offlow-channels in a micrometer scale with channel widths of about 100 μm and channelheights of about 50 μm, which affects the Reynolds number of the fluid transportedthrough the microchannel. The Reynolds number is a dimensionless value and is ameasure for the ratio of inertial forces (ρν) to viscous forces (ν /2r) and describes the

28


flow conditions. The Reynolds number is defined as:

Re =2rρν

η=

dQ

νA(1.10)

In Equation 1.10 r denotes the radius, d the diameter of the channel, ν the averagefluid velocity, ρ the fluid density, η the dynamic and ν the kinematic viscosity. TheReynolds number classifies flow behavior into three groups, laminar flow conditions atRe < 2300 followed by the transition area at 2300 < Re < 4000. At Re > 4000 tur-bulent flows are possible, depending on factors as pipe roughness and flow uniformity.Equation 1.10 denotes the Reynolds number of flow conditions within a tube. Tocalculate the Reynolds number of rectangular shapes the hydraulic diameter dH hasto be introduced. The hydraulic diameter is defined as presented in Equation 1.11.

dh =4A

pwet

(1.11)

pwet denotes the wetted perimeter of the channel. The Re in rectangular channels isdefined in Equation 1.12.

Re =ρvdhμ

=QdhνA

(1.12)

Assuming a microchannel with the height of 50 μm and the width of 100 μm thewetted perimeter averages 300 μm. If water is transported in this microchannel withv = 1 m/s, ρ averages 1000 kg/m3 at 25 ◦C and μ=0.891 kg/ms the Re is calculatedusing Equation 1.14.

dh =4 ∗ 5 ∗ 10−9

3 ∗ 10−4(1.13)

Re =100016.67 ∗ 10−5

0.891= 0.07 (1.14)

A Reynolds number of 0.07 describes extremely laminar flow conditions. Forces af-fecting fluids are shear forces within the fluid and shear forces between fluid and thewall. The shear forces arise from the displacement of fluid-elements in flow-direction,the influence of shear forces on the fluid behavior is shown in Figure 1.10 on thenext page. Figure 1.10 shows the velocity distribution of particles in a fluid betweentwo plates. One plate is moving the other remains stagnant. The particles of thefluid near the plate attach to the surface surface and show the same flow behavior,therefore move with the same flow velocity. This behavior is called the conditionof detention. The velocity of a single particle in a solution between the two plateschanges dependent on the distance of the particle to both plates. The nearer a par-ticle is positioned to the moving plate the higher is its flow velocity, resulting in a

29

1 Introduction

Figure 1.10: The left picture show a fluid between two plates, where the upper one moves

with the velocity v and the lower plate remains stagnant (v=0). The right

picture describes the velocity distribution of the tube´s cross section. The

arrow-length is direct proportional to the velocity of the particles in this section.

velocity-gradient perpendicular to the vector of flow velocity. The velocity-gradientin a channel is responsible for the shear stress within fluid-elements and is definedaccording to Newton as:

τ = ηdv

dy(1.15)

F = ηAdv

dy(1.16)

η denotes the dynamic viscosity, that decreases with increasing temperature.Many microchannels have rectangular cross section with four rigid walls, resulting ina parabolic flow profile as illustrated in Figure 1.11.

30


Figure 1.11: a)Pressure drop along a pipe, from high pressure (red) to low pressure (blue)

b)Flow profile in a rectangular channel. Shear forces between channel surface

and fluid induce a parabolic flow profile, with a flow rate converging against

zero at the fluid channel interface and a maximum flow rate in the middle of

the channel. The flow profile of the fluid in the tube is described with a color

scale, ranging from high velocity (red) to low velocity (blue) [55].

The shear forces of both interactions between fluid-elements and between fluid andwall induce a pressure drop within the microchannel over a distance l, that is directproportional to the flow rate. The proportional constant is the fluidic resistance Rand can be described with the Hagen-Poiseulle equation:

Δp = QR (1.17)

Q denotes the volumetric flow rate and R the fluidic resistance. By inserting theRe from equation 1.10 combined with equation 1.11 as R into the Hagen-Poiseulle

equation, the pressure drop in a microchannel with rectangular cross-section can becalculated using the following relationship:

Δp =128ηlQ

πd4h(1.18)

The mean velocity νm can be evaluated at a given Δp by inserting the parametersinto Equation 1.19.

ν̄ =Δp

8ηlr2 (1.19)

To calculate the flow profile in a microchannel under the consideration of flow velocity,pressure, temperature and density, the Navier-Stokes equation has to be applied. The

31

1 Introduction

Navier-Stokes equation is a partial differential equation, needed to calculate the flow-velocity on various positions in a flow-channel. The Navier-Stokes equation is definedas:

ρ(∂ν

∂t+ ν∇ν) = −∇p + η∇2ν (1.20)

∇ denotes the nabla operator and stands for the gradient of a scalar field, or thedivergence of a vector field. This partial differential equation is used to calculaterelations involved in unknown functions of several independent variables and theirderrivates.

1.4 Surface modification strategies to reduce protein

fouling

Unspecific adsorption of proteins or lipids on surfaces is of major concern in numerousimportant biomedical and analytical applications such as cell culturing [56], tissue en-gineering, implantable devices [57], analytical measurements [58] and more. Duringanalysis of samples containing proteins, proteins tend to adsorb on surfaces, startingwith single ones that interact with more proteins, causing a cascading reaction ofprotein adsorption. Proteins with a high adiabatic compressibility can unfold on thesurface magnifying the fouling. Concerning sensing surfaces fouling plays a key chal-lenge, including electrochemical or affinity biosensors and is the subject of extensiveresearch [59, 60, 61]. During analysis of complex samples, including blood, saline,urine and environmental samples like sea water samples protein adsorptions reducethe selectivity (affinity sensors) and the sensitivity (electrochemical sensors) of thesensor. For electrochemical measurements the preservation of a stable and unalteredelectrode surface is a premise for reproducibility and repeatability of analytical mea-surements. Protein adsorptions on electrode surfaces (electrode fouling) decrease theelectrode area and therefore also the measurement signal.Lab-on-a-Chip devices provide high surface to volume ratios offering a large attractionarea for protein adsorption and therefore result in an alternation of the measurementsolution. The problems concerning protein adsorption boost the study of protectivesurface modifications and made significant progress in the last years [62, 59]. Ta-ble 1.10 lists common nonfouling surface modifications.

32

1.4 Surface modification strategies to reduce protein fouling

Table 1.10: Nonfouling surface modifications

Category Examples Refs.

Carbohydrates AgaroseMannitolSorbitol

[63, 64,65]

Synthetic polymers Poly(ehylene glycol) (PEG)PolycrylamidePlurinicsPoly(ethylene oxide) (PEO)Poly(carboxybetaine)HydrogelsAlkanethiolate derivates

[59, 66,67, 60,65]

Proteins Bovine serum albumin (BSA)Surface layer proteinLaminin

[68, 69,70, 65]

For this work the nonfouling characteristics of bovine serum albumine, poly(ethyleneglycol)and surface-layer proteins, very heterogeneous surface modifications were examined.

1.4.1 Bovine serum albumin (BSA)

Bovine serum albumin is a negatively charged very soluble and stable protein of arather small size having a molecular mass of about 66.4 kDa. It is considered as amodel protein in many applications and gained a lot of popularity in the past decades[71, 72, 73, 74]. BSA is a well studied protein and the high availability as the proteinmost common in blood boosts further studies [75] and is used in a broad area ofapplication as for Enzyme-linked immunosorbant assays (ELISAs) [76], as blockingagent for antibody labeling experiments [77], for analytical purposes like proteinquantification and as surface modifying proteins forming nonadhesive regions for cellpatterning approaches [65, 70, 69]. BSA is also used to stabilize enzymes duringcatalytic reactions, for instance during digestion of DNA. Figure 1.12 shows the 3dimensional structure of albumin.

33

1 Introduction

Figure 1.12: Quaternary structure of albumin

BSA covered surfaces are used for cell patterning experiments and was thereforeused as coating material of structured silicon wafers to perform cell patterning essays.Bovine serum albumin treated surfaces avoid cell adhesion for two days under serum-free conditions and enable the formation of cell patterns. The dependency of BSAon serum-free conditions for cell repulsion can be explained by the attachment ofserum proteins on BSA surfaces facilitating cell adhesion [65, 69]. Figure 1.13 showsstructures of cell patterning experiments.

34


Figure 1.13: Cell patterning experiments on differently modified surfaces. a-c) present first

experiments to create a liver on chip [31]. a)presents a radiate pearl-chain pat-

tern of hepatic cells on specialized electrodes. b) a heterogeneous integration

pattern of hepatic cells (green) and endothelial cells (red) and c) the control

group with two randomly distributed cells without dielectrophoresis manipu-

lation. d)in this case a membrane based patterning experiments covered with

BSA had been used to create stable, and sharp BCE cell patterns [69].

1.4.2 Poly(ethylene glycol) (PEG)

Poly(ethylene glycol), abbreviated PEG is termed after polyols and has a molecularweight below 20 kDa. If the molecular mass is higher the molecules are termed poly(ethylene oxide) or poly(oxyethylene). In general PEG seems to have the structure of asimple molecule with one chain, that exists in a variety of different forms. Both linearand branched molecules with variable molecular weights are available. The chemicalprecursor of PEG is the trace hydroxide acting as initiator, a difunctional polymerthat grows in both directions. The structure of PEG and a few derivatives can beobtained from Figure 1.14. PEG forms complexes with metal cations, is highly mobile,is used to precipitate proteins beside nucleic acids and forms two-phase systems withother polymers in aqueous solutions. It is also a non-toxic molecule and beyond thatwith a low immunogenicity and antigenicity resulting in an increased utilization forbiomedical and biotechnological applications. By activating the hydroxyl-groups orif modified the carboxylic-groups of the chain ends, proteins such as antibodies andenzymes can be covalently bonded having a small impact on their chemistry and

35

1 Introduction

activity. What can be controlled is the solubility of certain compounds and theirsize. The protein or larger molecule rejecting property of PEG and also remains afterPEG immobilization reactions on surfaces. If surfaces are covered with PEG, themolecule remains mobile, hydrated and forms elastic spheres, able to deform wheninteracting with proteins, resulting in a a temporary deformation, but the proteins arenot able to attach or significantly compress the sphere [62]. Figure 1.14 shows furtherapplication possibilities published in recent years. First the modification of PEG withthe copolymer poly(L-lysine)-graft-poly(ethylene glycol), abbreviated PLL-g-PEG, athin polymer layer where little amounts cover a larger surface (1 g for 1000 m2)and act as a waveguide. Therefore the PLL-g-PEG copolymer forms a biospecificinterface [78]. Second PEG used as a shielding polymer on virus surfaces to preventinteractions between the coated virus and native receptors or neutralizing antibodies.The incorporation of ligands on the PEG layer enables re-targeting [79].

Figure 1.14: Structure of PEG and derivates and application possibilities. a) linear and neu-

tral polyether with the two active hydroxy groups; b) PEG chains with different

length and functional groups to modify surfaces; c) PEG Sigma-Thiol; d) PEG

chains attached to a surface to immobilize recognition elements. The PEG chain

had been modified with the copolymer poly(L-lysine)-graft-poly(ethylene gly-

col) (PLL-g-PEG) [78]; e) Detailed structure of PLL-g-PEG, the Poly(L-lysine)

backbone with the PEG side chains; d)PEG as shielding polymer [79];

36


The discovery of PEG lead to a whole new era of material modification possibilitiesand to an extensive study of PEG behavior. Five main disciplines of PEG-usage canbe distinguished.

1. PEG can be used to collect proteins and nucleic acids from solutions, providingideal purification possibilities [80].

2. When PEG is combined with dextran in a buffer solution, this mixture canbe useful for the treatment of biological materials, especially for purificationpurposes [81].

3. PEG is able to fuse with cell membranes and can form protective layers oras an additional modifying layer covering transport vectors for drug deliveryapplications [79].

4. PEG is able to bind proteins, resulting in a nonimmunogenic and nonantigeniccomponent increasing the protein´s lifetime [82].

5. PEG surface modification rejects protein adsorption and acts as a non-foulinglayer [83].

This list was published in 1992 and should be longer today. This short section onlyprovides a brief introduction into polymer science focusing on Poly(ethylene glycol).

1.4.3 Surface layer proteins

Regularly structured surface layer proteins, also termed S-layer proteins, shape theexternal cell envelope of a variety of prokaryotic cells. The surface structure has beenobserved as an almost universal feature of archaea and on different species of walledeubacteria, on both gram-positive and gram-negative eubacteria and archaea. S-layerproteins present the outer surface of archaea, species with a remarkable evolution-ary background, that are able to exist in extreme habitats like extreme halophiles(Halobacteriales), sulphur-dependent extreme thermophiles (Sulfolobales & Thermo-proteales) and methanogens (Methanothermus). This leads to the assumption, thatthis protective S-layer adds to the microorganism’s selection advantages by actingeventually as a molecular sieve, molecule or ion trap, or an additional structure forsurface recognition and cell adhesion skills or as an auxiliary element for cell shape

37

1 Introduction

determination [84].

Protein structure

The S-layer proteins possess the ability to form regular lattice structures with identicalpore sizes, ranging between 2 to 8 nm, with a relative molecular mass of 40 to 200kDa. The pores themselves represent 30 - 70 % of the whole surface area. Thedistance between two centers varies between 3 - 35 nm and the layer thickness rangesfrom 5-20 nm. The exact structures of all S-layer proteins were not published yet,but reveal a rather similar composition of 20 % α helices, 40 % β sheets and between5 % to 45 % β turns [85, 86].Dependent on the microorganism and also conditions of cultivation S-layer proteinsdevelop oblique (p1,p2), square (p4), or hexagonal (p3, p6) lattice symmetries, as canbe seen in Figure 1.15.

oblique

hexagonal

squarep1 p4p2

p6p3

Figure 1.15: S-layer lattice types.

The notation p1 - p6 defines the point group symmetry and consequently theamount of subunits involved in the formation of a closed morphological unit. The de-pendency of cultivation parameters on the S-layer synthesis was studies by Sara and

38


coworkers in 1996. The model organism Bacillus stearothermophilus PV72/p6 wascultivated under oxygen-limited conditions under which the S-layer protein SbsA/p6,with a molecular weight of 130 kDa and a hexagonal lattice structure was synthe-sized. By increasing the oxygen rate the microorganism started to express SbsB/p2with an oblique lattice structure and a molecular weight of 97 kDa. This variationof S-layer proteins encoded by one single organism can be explained by the existenceof at least one silent S-layer gene, where only one is expressed at a fixed time [87].

Forces influencing the S-layer assembly

The investigation of the amino acid content of the S-layer proteins enabled a betterunderstanding of S-layer assembly processes. For instance sulphur containing aminoacids are quite rare or even nonexistent, which indicates that disulfide bonds do notcause subunit binding. In contrast 40 - 60 mol% are presented by hydrophilic aminoacids, indicating a high amount of hydrophobic interaction during assembly. Anothermajority, one quater of the S-layer protein amino acid content, is charged and there-fore generates ionic bondings. The N-terminal end of the S-layer proteins, a morehydrophilic region of the protein, provides the binding between the peptidoglycanof the microorganism and the surface layer, while the C-terminus holds most of thesurface-located amino acids. For the assembly process,the N-terminal domain is indis-pensable, while a truncated C-terminal domain has no influence on S-layer assembly[85]. The intensive examination of the binding forces occuring between single sub-units and subunits with the membrane surface gave rise to the assumption that thebonds between S-layer subunits are stronger than between the S-layer subunits andthe underlying layer [86]. Dependent on the assembly conditions as pH value, ionicstrength and ion composition the S-layer units can assembly into closed vesicles, openended cylinders and into flat sheets [34].

Surface properties of the S-layer

The crystallized lattice structure provides a neutral charge on the surface becausecharged carboxylic acid groups are neutralized by the free amino groups from lysineresidues, while the inner face shows a negative charge, caused by a high number ofcarboxylic acid groups [88]. The charge can also be influenced by changing the pHvalue of the surrounding solution. The Surface layer provides a stable structure andseemed to be very difficult to be removed from bacteria surfaces. Proper effective

39

1 Introduction

detergents are for instance highly concentrated hydrogen-bond-breaking agents likeguanidinium hydrochloride (GHCl), metal-chelating agents like EDTA, cation substi-tution agents and pH changes. To increase the surface stability, the S-layer can bemodified with glutaraldehyde, which causes a charge density of 1.6 carboxylic acidgroups per nm2 in case of square S-layer lattice from Lysinibacillus sphaericus CCM2120 and therefore a negative surface charge. A higher amount of carboxylic-groupsand amino-groups inside the pores prevent charged macromolecules to penetrate thelattice structure [68].

Non-fouling properties of S-layer proteins

The non fouling behavior of S-layers was investigated in context with membranefouling. In combination with microfiltration membranes the S-layer acted as activefiltration layer in nanometer range. The combination of membranes and S-layer wastermed S-layer ultra filtration membranes, abbreviated SUMs. Figure 1.16 shows theschema of the SUM fabrication method.

microfiltration membrane

Figure 1.16: Schema of the SUM preparation; a) microfiltration membrane b) S-layer protein

crystallization on the surface; c) Cross linking of the S-layer with glutaraldehyde

S. Weigert and M. Sara analyzed in 1996 the non-fouling properties of SUMs andthe influence of certain parameters. It turned out that S-layers covered microfiltrationmembranes perform a regular porous structure, with constant pore size in contrast tothe heteroporous structure of ultra filtration (UF) membranes and indeed offered non-fouling properties. By modifying the S-layer surface structure, the surface propertiescan be changed from negatively charged to positively charged, hydrophilic to hy-drophobic without changing the crystalline structure of the SUMs. The examination

40


of the SUMs lead to the assumption that many parameters influence the protein ad-sorption process, net charge and charge density, as well as the accessibility of chargedgroups and the hydrophobicity, beside other factors as pH and protein size. Charge tocharge interactions dominated in general hydrophobic bonds. The best SUM perfor-mance concerning nonfouling reactions is afforded by hydrophilic uncharged surfaces[68].

41

2 Materials and Methods

2.1 Micro total analysis system (μTAS)

2.1.1 Materials

2.1.1.1 Solutions

Table 2.1: Detailed overview of the solutions used for LOC experiments

solution composition/comment manufacturer

Polydimethylsiloxane

(PDMS)

1:10 dilution of the Sili-cone elastomer. Curingagent with the SiliconeElastomer Base

Sylgard 184 Silicone KitSylgard

Ethanol 70% purity VWR

2-propanol 99.5 % MERCK

bright and shiny dish

liquid concentrate

randomly Dreco Werke

Phosphate buffered

saline (PBS)

70.1 g Sodium chloride12.8 g di Sodium hydro-gen orthophosphate

BDH

4.4 g Sodium dihydrogenorthophosphate2.0 g Potassium chloride1:10 dilution in deionizedwaterpH 7.2

43


Table 2.2: Detailed overview of the solutions used for LOC experiments


Potassium chloride

(KCl)

- Sigma Aldrich

Potassium hexacyano-

ferrate (III)

insulation from light Alfa Aesar (A JohnsonMatthey Company)

D+ Glucose diluted in PBS with 1 MKCl

MERCK

Glucose Oxidase from

Aspergillus niger

diluted in PBS in a stan-dard concentration of 100 μ

g/ml

Fluka Bio Chemika

Tris(hydroxymethyl)-

amino methyl buffer

(TRIS)

diluted in distilled water,pH adjustment with HCl;pH 7.2

MERCK

Tris buffered salin

(TBS)

50ml TRIS150 ml NaCladjust to pH 7.4

MERCK

Anti-SbpA antibody produced in rabbit (affinityisolated antibody)

BAXTER

Anti-rabbit IgG (whole

molecule)-TRITC

- Sigma-Aldrich

Citrat buffer pH 4.2 MERCK

44


2.1.1.2 Equipment

Table 2.3: Detailed overview of the Equipment used for LOC experiments

instrument manufacturer

Plasma asher Femto Diener Electronic

Borofloat glass slides Thermo Scientific

Cover glass slides VWR

Two-component epoxy adhesive(Loctite: 9492)

2.5 ml glass syringes HAMILTON

Syringe pump kd Scientific

L-Valves and PEEK tubes:PEEK tubin 1/32" ID:130 μmTygon tube ID:0.51 mm, wall:0.91mm

Upchurch Scientific

Micro Static mixig TEE Upchurch Scientific

Software: EC-Lab V 9.98

Multi potentiostat VMP3 Princeton Applied Research

Light microscope MZ16 Leica

Fluorescence microscope TE 2000-S Nikon

45


2.1.2 Method

2.1.2.1 Microfluidic monitoring station

Cellmos 2 measurement station

The measurement station consisted of a pumping station, a chip, an aluminum chipfixture, a water pump, a connection box and a multichannel-potentiostat. A syringepump was used to simultaneously inject the samples into the chip using four differentsamples at the same flow rate. The chip fixture itself had various functions involvingtemperature regulation and contained electronic connectors to the amperometric sen-sor and the potentiostat. Temperature was adjusted using a water pump, that wasconnected via silicone tubes. The connection between the chip and the multichannel-potentiostat originated from the upper part of the fixture, two flexible blocks madeof polycarbonate, equipped with two plug-ins and gold tips, pressed onto the con-tact pads of the chip to get contact with the measurement chamber. The plug-insconnected the fixture with an extern connection box, that eventually connected withthe potentiostat. To get an idea of the whole setup, Figure 2.1 provides a detaileddescription of all components, an even more exact Figure of the chip placement isshown in Figure 2.2.

Figure 2.1: Setup of the Cellmos 2 measurement

station

1. syringe pump

2. syringe with the sample

3. valves and tubing system

4. chip

5. chip fixture and heating block

6. connection box

7. potentiostat connector

46


1. chip placed in the middle of the fixture

2. aluminum block of the chip fixture

3. polycarbonate block with copper wiresthat are connected with the gold tips

4. connector plug for the external box

Figure 2.2: Chip placed in the Cellmos 2 fixture

Cellmos 3 measurement station

For electrochemical measurements on LOC-deviced a Cellmos 3 LOC station wasused. Figure 2.3 illustrates a detailed configuration of the measurement station.

Figure 2.3: Scheme of the Cellmos 3 LOC station

1. microscope

2. experimental interface

3. syringe pump

4. syringe with probe


6. chip fixture

7. chip

47



2. chip

3. water pump inlet(heating system)

4. printed circuit boards (PCBs)

5. heating plate

Figure 2.4: Scheme of the Cellmos 3 chip fixture

The Cellmos 3 LOC-station is an advanced Cellmos 2 LOC-station and thereforeregarding the setup the two are quite similar. The Cellmos 3 LOC-station consisted of7 major components. First the optical microscope (light microscope) placed above theLOC fixture for a visual contact with the measurement chamber, to check electrodeconditions and bubble formations. Second the interface between experimental stationand the operator, a PC, where the essential software EC Lab, necessary for electro-chemical measurements was installed. Third the pumping station, a syringe pumpwith adjustable pump velocities. Fourth the syringe containing the measurement so-lutions. Fifth the valves and the tubing system to control flow directions. Sixth thechip fixture consisting of an aluminum block with integrated heating plate containingthe micro total analysis system itself placed on the heating plate and fixed in printedcircuit boards (PCB). The PCBs are equipped with plug-in connectors, necessary toconnect the LOC device with the multichannel potentiostat. Seventh the multichan-nel potentiostat suitable for various electrochemic measurement methods and able tomeasure on 14 channel parallel and independently.

2.1.2.2 Microfabrication of platinum chips

Principle

The platinum chip was fabricated using UV lithography. In UV-lithography a pos-itive photoresist is spinned onto a substrate such as a silicon or glass wafers andprovides a sacrificial stencil layer for a further lift off process. So after the spinningstep, the photoresist-layer needs to be covered with a negative mask, representing the

48


chip design, while exposed to UV light. The unprotected area is removed by washingthe wafer with a solvent. The metallic structure itself is fabricated by sputtering andlift off techniques. A sputtering machine sputters a metallic layer onto the preex-isting pattern on the silicon wafer with a defined thickness. The lift off procedureis necessary to remove the sacrificial and stencil photoresist layer. A solvent is usedto wash away both the sacrificial layer and the metal layer on the top of the sacrifi-cial layer. What remains is the metal layer that had direct contact with the substrate.

Procedure

The fabrication of the platinum chip commenced using pretreated glass substrates.The glass was dry baked (170◦C for 5min) and primed (TI-Prima, MicroChemicals).On the pretreated glass an initial nonphotosensitive resist layer (LOR 3A,MicroChem)was spun on the wafer at 3000 rpm for 35 s. Secondly a positive photo resist (positiveresist AZIMIRTM 701) was applied and soft baked at 110 ◦C for 60 s. By using aMJB 3 mask aligner from SUESS Microtec (350 W mercury lamp, exposure time6.35 s) the resist was patterned and afterwards developed for 25 s (AZIMIF 726). Inan additional sputtering step a 10 nm titanium and a 200 nm platinum layer wassputtered onto the pre-patterned glass substrate and placed into a solution of NMP(1-methyl-2-pyrrolidinon) at a temperature of 70-80 ◦C for at least 60 min. After thetreatment with NMP the required platinum structure of the chip remained, namelythe structure of the contact pads and the electrodes. Figure 2.5 shows the schematicdescription of the UV-lithography and the lift off process.

49


Figure 2.5: Schematic description of the UV lithography:a)glass wafer, b)positive photoresist

spun onto the waver, c) UV treatment after the coverage of the wafer with a mask,

d) only the photoresist beneath the transmissible film stabilizes and remains;

sputtering with platinum, e) treatment of the surface with a solvent, reacting

with the photoresist, f) metal film beneath the uncovered wafer surface remains

50


2.1.2.3 Fabrication of Microfluidics

First of all the microfluidic was designed via software, mainly TARGET 3001, at theend also CLE WIN 7.0. Both programs are specialized on exact drawing for masks ofmaster molds, which were then printed on transparencies. The mold was fabricatedvia UV-lithography, followed by a wet etching step. The fabrication procedure tookplace in a clean room. After the pretreatment of the silicon wafer with a dry bake step(120 ◦C 5 min) the TI-Prime, an adhesion promoter was spun onto the wafer (3000rpm for 20 s) and then soft baked at 120 ◦C for 2 min. In the next step the structuredetermining hard polymer SU-8 was spun onto the silicon wafer at 3000 rpm for 35s. The thickness averaged 30 μm equivalent to the height of the microchannel ofthe microfluidic. To conserve a homogeneous thickness, the photoresist was hardenedby a soft bake step (65 ◦C 60 s and then 95 ◦C 60 s). Afterwards the mask wasclamped onto the photoresist and exposed to UV-light (7 s) to selectively crosslinkthe polymer. In a post exposure process (65 ◦C 60 s and then 95 ◦C 60 s) the SU-8was stabilized to prevent melting in the developer. The mold was then developed in aSU-8 developer until the unexposed parts of the SU-8 were completely removed (60 -70 s), cleaned with isopropanol for 10-15 s and then hardened through a hard bakingstep (65 - 160 ◦C 15 s). After the hard baking, the master mold cooled down to roomtemperature. This mold was used as a negative mask of the microfluidic structureand was reused. The soft polymer PDMS was spilled onto the wafer in a petri dishand air bubbles were removed by putting the glass petri dish into an desiccator andusing a vacuum pump for at least 5 minutes until no more bubbles appeared. ThePDMS was cured at room temperature, but since this takes more than one day, thebaking of the polymer was also performed at 70 ◦C for 2 hours.

2.1.2.4 Chip assembly

The Chip consisted of 3 layers. The upper layer, a glass slide where in and outlettubes (ports) were fixed. A middle layer, the microfluidic, and a lower layer of thelithographic machined glass slide with the contact pads and the electrodes.After the polymerization of the PDMS the microfluidic was carefully cut out of thepetri dish using a scalpel. With a metal tube of less than 1 mm in diameter, smallholes were pitched into the in- and outlets of the microfluidic to connect the in- andoutlets of the glass slide with the microchannels. The glass slide of the same size as

51


the fluidic was drilled at the exact positions of the in and outlets of the microfluidic,using a micro drill. Before bonding, the layers of the LOC-device (platinum chip,fluidic and upper glass slide) were cleaned using different solutions, such as ethanol,isopropanol and dish liquid concentrate. Another possibility to clean the fluidic wasthe treatment in a ultrasonic bath at the lowest intensity for 10 minutes. The exacttreatment varied, according to the chip condition.To bond the glass slides to the PDMS microfluidic and enable the fabrication of a sealLOC device the surfaces were treated with oxygen plasma, using a plasma cleaner.The oxygen plasma (oxygen ions) was created by collision ionization of oxygen ions inan electromagnetic field and subsequent bombardment of the surface of interest. Theoxygen plasma activated groups on the treated surfaces, for instance the Si−CH3 onsilicon surfaces , that reacted with the oxygen and formed Si-OH. The Si-OH groupsstarted a condensation reaction when they got in contact with other activated groups.After the surfaces were activated and bearded against each other, covalent bonds wereformed. The activation reaction was performed under low pressure of 0.8 mbar for 30s at 40 W and 22 ◦C. After the activation the slides were pressed against each otherand left in this state over night to ensure complete bonding resulting in a seal device.The in- and outlet tubes, also called ports were stuck onto the chip by using a two-component adhesive. The tubes consisted of two different sizes, a 1

32” tube in a

116” tube and were stuck into the drilled hole of the glass slide and the holes of the

PDMS microfluidic, using the epoxy adhesive. The two-component adhesive was usedfor fixation and for sealing purposes. The whole procedure is illustrated in Figure 2.6.

52


Figure 2.6: Schematic description of the chip assembly process split in 3 steps. Firstly the

plasma bonding of PDMS onto the glass surface secondly the fixation of the

tubes into pre-drilled holes in the glass top layer using an epoxy adhesive and

thirdly again the plasma bonding step of the upper layer onto the lower layer.

2.1.2.5 Surface modification with S-layer protein

The concentration of the S-layer protein solution averaged 1 mg/ml and was storedat 4 ◦C. The crystallization reaction was performed under room temperature afterdiluting the S-layer protein in a buffer solution. In case of SbpA 0.5 mM TRIS with10 mM CaCl2 pH 9 was used in a ratio of 1:10, the same ratio in case of SbsB butwith Citrate-buffer pH 7.2. The solution was pumped through the chip at a flowrate of 0.7 mm/min over night. After the crystallization the chip was cleaned withdeionized water for 1 h at a flow rate of 20 mm/min to remove non crystallizedproteins. Then the chip was stored at 4◦C for 2 hours to stabilize the S-layer proteinlattice structure. Important was that the S-layer did not dry out during the wholeprocedure.

2.1.2.6 Imaging by Atomic force microscopy

Principle

Atomic force microscopy (AFM) belongs to scanning probe microscopy, an imaging

53


tool able to achieve molecular resolution. Many biological materials were studied us-ing AFM such as proteins, membranes, membrane bound proteins, nucleic acids andmany more. Atomic force microscopy measures forces acting between a fine tip placedon a cantilever and the sample of interest caused by attractive or repulsive forces re-sulted from interactions between tip and sample and leading to positive or negativebending of the cantilever itself. The specimen stage can be moved by a piezoelectricscanner. Piezo-crystals are amongst other things ceramics expanding or contractingin the presence of voltage gradients and work also in the opposite direction, meaningthat they create a voltage gradient when they are forced to expand or contract. Piezo-ceramics are used in AFM for a precise three dimensional movement of the specimen.A micromachined cantilever and a special tip together form the probe, which is thepart of the AFM that interacts with the sample. For the experiments a silicon ni-tride (Si3N4) tip on a V-shaped cantilever was used. Important characteristics ofcantilevers are the force constant, the resonant frequencies and the spring constant.An optical lever acts as detection system. In atomic force microscopy a laser beam isreflected off the back of the cantilever to a position sensitive photodetector (PSPD)consisting of four photodiodes. The position of the laser beam on the diodes canbe converted into the angle of deflection of the cantilever, its exact status and sothe height of the sample. The tip scans over the sample and the deflection of thecantilever has to be recorded and translated into heights to produce a three dimen-sional surface of the specimen. There are in principle two types of scanning methods,either the tip has to be scanned over the probe as described, or the scanning of thepiezo-electric specimen stage while the tip remains constant. The three dimensionalmovement of the specimen stage can be converted into a height image.The operation of AFM imaging can be performed in different modes, three modes ingeneral: the contact mode, the non contact mode and the tapping mode. In contactmode repulsive forces, more concrete Pauli Exclusion Principle forces dominate. Thebehavior of the cantilever is determined by Hook´s law:

F = −k ∗Δx (2.1)

k denotes the spring constant of the cantilever. The operation can be performedunder constant height or constant force conditions, where the first condition meansthat the cantilever has a fixed height and the second means constant force/distancebetween tip and sample surface. The moving of the scanner in z-direction is recordedand contains the information about the surface topography of the sample. For the

54


contact mode, the choice of cantilever is highly dependent on the tip material, softtips protect the sample surface and enable high resonant frequency, avoiding vibra-tional instabilities. The advantages of contact mode are high atomic resolutions, highscan rates and the possibility of scanning rough samples. Disadvantages are capillaryand electromagnetic forces sometimes causing a drift of the tip towards the samplesurface resulting in material damage.In the non contact mode, the system operates in the distance region where attractiveforces dominate. The cantilever is stimulated to oscillate. The oscillation changes af-ter interactions between tip and sample surface, causing a phase-shift of the resonance-frequency. The phase shift of the resonance-frequency is detected and translated intosurface texture. In non contact mode the spring constant is higher than in contactmode, avoiding sticking to the sample surface at small amplitudes. The non contactmode is mainly performed in vacuum and used for extremely sensitive sample sur-faces.In the tapping mode the cantilevers oscillates close to its resonance frequency influ-enced by the forces between sample and tip, causing changes in the amplitude andtherefore a phase-shift. The resonance frequency remains constant and causes changesof the specimen stage in z-direction to conserve constant resonance frequencies. Themovement in z-direction transmits the height of the specimen-topography. In the tap-ping mode weak forces are active and cause less specimen damage. A disadvantageof tapping mode is the slow scan speed compared to contact mode.

Procedure

After recrystallization of S-layer proteins on various surfaces such as gold, PDMS,platinum and glass, the sample surface was covered with a drop of NaCl and placedon the specimen. For AFM imaging of the S-layer lattice the contact mode was used.Figure 2.7 (next page) shows the principle of atomic force microscopy.

55


Figure 2.7: Principle of Atomic Force Microscopy: A laser beam, emitted from a light source

launches onto the back of the cantilever and is reflected to a position sensitive

photodetector (PSPD). The photodetector consisting of four diodes marked as

A,B,C and D and transduces the position signal into the angle of deflection.

This information is used to create a height distribution of the specimen. The

force types occurring between tip and specimen are dependent on the operation

modes, in contact mode attractive and repulsive forces dominate. The feedback-

loop corrects the position of either the cantilever itself or the piezzo-electric

specimen stage.

2.1.2.7 Simulation of microfluidic behavior

The simulations presented in Section 3.1.1 were performed by using the softwareCOMSOL Multiphysics 3.4, a specialized software for microfluidic purposes and manymore. The program consists of about eleven modules, while only the MEMS Modulewas used:

• MEMS Module/Microfluidics/Incompressible Navier-Stokes

• MEMS Module/Microfluidics/Convection and Diffusion

COMSOL Multiphysics was programmed to deal with complex physical phenomena.After starting the desired module and drawing a representative geometry the sub-

56


domain and boundary conditions were defined. By creating a mesh of selectableaccuracy the indefinite space-coordinates become discretized and on every interceptpoint of the mesh the partial differential equations get solved using a numerical solver.An extra fine mesh size is shown in Figure 2.8.The flow conditions in the simulations are described by Navier-Stokes equations and

Figure 2.8: Extra fine mesh size in a microchannel.

the diffusion behavior by Fick´s law. The pressure of the outlet was set to zero, whichinduces a pressure drop in the microfluidic geometry. By using the post processingtab the calculated characteristics were displayed in different ways. In this work theslice and arrow plot were mainly used (see Figure 2.9).

57


Figure 2.9: Slice and arrow plot of a flow and diffusion simulation. The arrows describe

the velocity field in the channel, with dimensions proportional to the flow rate.

The slice plot (colors) characterizes the diffusion behavior within the cylin-

dric geometry. The colored bar shows the concentration maximum (red) and

minimum(blue).

The partial differential equations can only be solved, if the error converges againstzero. The parameters used for the simulations were matched to the parameters ofwhole blood. The values can be obtained from Table 2.4.

Table 2.4: Parameters in dependence of whole blood

symbol value unit

d 1 ∗ 10−4 mρ 1.055 ∗ 103 kg/m3

u 1 ∗ 10−4 per inlet m/s

η 5 ∗ 10−3 Pas

The equation necessary to calculate flow behavior is the Navier-Stokes equationat steady state conditions, as mentioned in Section 1.3.3 and is revisioned in Equa-

58


tion 2.3.

−∇η(∇u+ (∇u)T ) + ρ(u∇)u+∇p = 0 (2.2)

∇u = 0 (2.3)

where ρ denotes density in kg/m3, u the velocity in m/s, η the viscosity in Pas andp the pressure in Pa. An extremely low Reynolds number of about 0.01 describes theabsence of turbulences in the microchannel, so that mixing can only be achieved bydiffusion of the molecular components and can be tested via simulations.The mass flux in the reactor is given by convection and diffusion, resulting in a massbalance as cited in Equation 2.4.

∇(−D∇c + cu) = 0 (2.4)

D denotes the diffusion coefficient in m2/s and c the concentration in mol/m3.

59


2.2 Surface plasmon resonance (SPR)

2.2.1 Materials

2.2.1.1 Solutions

Table 2.5: Detailed overview of solutions used for SPR


Tris buffered saline

(TBS)

50 mM TRIS150 mM NaClpH 7.4

TRIS buffer 0.5 mM TRIS buffer with10 mM CaCl2

Bis(Sulfosuccinimidyl)

suberate solution BS3

1 mg/ml BS3 dissolvedin HEPES bufferadjust pH to 8.995

Pierce

Albumin from human

serum (HSA)

96-99 % purity Sigma Aldrich

Albumin from bovine

serum (BSA)

96-99 % purity Sigma Aldrich

Glutaraldehyde (GA) 2 % Glutaraldehyde in0.1 M Caco buffer % pu-rity in Caco buffer

Cacodylic acid

sodium salt (CACO

buffer)

100 mM

Natrium Borhydrate

solution

10 mM NaOH plus 10mM NaBH4

Sigma Aldrich

4-(2-hydroxymethyl)-

1-piperazinesulfonic

acid (HEPES) buffer

10 mM HEPES in waterat pH 8-9

MERCK

60


2.2.1.2 Equipment

• sputtering machine

• sample chamber made of PDMS

• peristaltic pump

• tygon tube ID:0.51 mm, wall:0.91 mm

2.2.2 Method

2.2.2.1 Principle

Surface Plasmon Resonance (SPR) Sensors are based on the measurement of propa-gating surface plasmons and count as refractometric sensing devices. Surface plasmons(SP) are electromagnetic surface waves, originating from coupled collective oscillationsof electron plasma in an electromagnetic field between metal-dielectric interfaces. SPsare transversal electromagnetic (TM) waves, with a magnetic field-vector H perpen-dicular to the electric field-vector E. Both field-vectors are perpendicular to thedirection of propagation. The propagation constant β describes the plasmon and isdefined in Equation 2.5.

ω

c

√n2mn

2d

n2m + n2

d

=2π

λc

√n2mn

2d

n2m + n2

d

(2.5)

ω denotes the angular frequency, c the speed of light in vacuum, nd the refractiveindex of the dielectric, nm the refractive index of the metal and λ the wavelength. Theintensity of the magnetic field reaches its maximum at the metal-dielectric interfaceand decays by a factor of 1/e into metal and dielectric. The intensity of the magneticfield is described by the penetration depth Lpen, the distance perpendicular to thesurface at which the field amplitude decreases by a factor 1/e. The propagation lengthLpro, the distance along the metal-dielectric interface necessary to drop the SP to 1/e

is another parameter describing the dissipation process.Light waves can only couple to surface plasmons if the wave-vector is parallel to themetal film and matches the real part of the SP propagation constant Re{β}. Asthe SP´s propagation constant on the metal-dielectric interface is larger than thewavenumber of the light, the wavenumber has to be increased to realize a couplingevent. The increase of the wavenumber can be reached by attenuated total reflection

61


and the aid of a coupling device. A prism coupler with Kretschmann geometry is acommon coupling device. In attenuated total reflection a high refractive index prismnp is used to lead a light beam to a thin metal film with the measurement chamberon its top, containing the dielectric solution (nd < np). At the base of the prism, thelight beam is totally reflected under the angle of incidence θ and penetrates the metalfilm via its evanescent field. The evanescent wave propagates along the metal surfacewith a certain propagation constant, that can be adjusted to match that of the realpart of the surface plasmon by varying the angle of incidence.

Re{β} =2π

λnp sinΘ (2.6)

By varying the thickness of the metal film, the coupling strength of the evanescentwave to the SP is enhanced. The SPR Sensors measure changes of refractive indicescaused by sample solution influencing the evanescent field. A variation of nd leads to achange of the SP’s propagation constant. The detection of the signal shift is achievedby using angular, wavelength, intensity and phase modulations. In angular modu-lation a transverse magnetically polarized light is launched into the prism coupler,that is matched to the sensor surface and mounted on a motorized rotation stage.A photodiode acts as a detector and measures the intensity change of the reflectedlight wave that is dependent on the angle of incidence as shown in Figure 2.11. Themonochromatic wave is used to excite a surface plasmon on the metal surface and thecoupling strength of the SP is observed in a range of angles of incidences Θ resultingin a angular reflectivity spectrum. The intensity of the reflected light δR is relatedto the shift of the resonant angle of incidence δΘres. The definition of δΘres is shownin Equation 2.7.

δΘres = δR/(δR

δΘ) (2.7)

( δRδΘ) denotes the slope of the SPR reflectivity curve described in 2.10. The changes

of the spectral position of the SPR dip minimum is determined using a centroid,polynomial or Lorentzian curve fitting.

62


angle of incidence θ [°]

refle

ctiv

ity [%

]

critical angle

δθ

δR

SlopeδR/δθ

Figure 2.10: Reflectivity spectrum. The reflectivity % is plotted against the angle ◦C.The

angle of incidence Θinc is at the minimum dip the curve describing the reflected

light beam. The shift of the dip refers to the change of the propagation constant

after the apply of a sample solution with a higher refractive index

SPR affinity biosensors translates the binding events of a molecule on the sensorsurface into an output signal. After the crystallization of for instance SAMs on thegold sensor surface the thickness of the layer can be determined via measurement ofthe angular spectrum. This binding event causes an increase of the refractive indexand depends on the concentration/amount of molecules attached to the sensor surface,resulting in a sensor response as defined in Equation 2.8.

Δn = (dn

dc)Γ

h(2.8)

(dn/dc) denotes the refractive index increment ranging in case of proteins between0.16 - 0.25 and Γ denotes the surface coverage in mass/area. By polynomial curve fit-ting the thickness of the attached molecule layer on the gold surface can be calculatedand in further calculations the surface coverage.

2.2.2.2 Setup of the SPR sensor

The setup of the SPR affinity biosensor in shown in Figure 2.11. The measurement

63


Figure 2.11: Schematic description of surface plasmon resonance (SPR): s) First the light

beam passes a polarizer. The chopper induces a polarized beam, launched into

a prism coupler with Kretschmann geometry; The beam is reflected on the

sensor surface, and detected by a photo-diode. b) View of the sensor surface in

the measurement chamber: The SP propagates along the metal-sample solution

interface; The the amplitude of the magnetic wave decreases perpendicular to

the sensor surface and the propagation direction.

flow cell was assembled, using a PDMS flow cell chamber positioned on a flow coverglass slide with fixed in and outlet. The chamber was fixed on the low refractive indexglass slide covered with a thin gold layer (30 - 40 nm), sputtered onto the glass. Arefractive index matching oil was applied between the sensor chip optically and theprism coupler. The sensor plus prism coupler was placed into the setup and adjustedby changing the angle of the motorized rotation stage of the detector and the separateone of the measurement chamber. As shown in Figure 2.10 the angle of incidence wasdetermined using an angular spectrum, where the angle of the lowest point of thelinear relationship between reflectivity and angle with the linear slope δR

δΘwas used

64


for the kinetic measurement. The first angular scan was used to calculate the thicknessof the gold layer, by using the software WINSPALL 3.1, specially programmed forSPR measurements. By using metric such as centroid, polynomial or Lorentzian curvefitting parameters of interest, such as the layer-thicknesses were calculated by iterativemethods. During kinetic measurements the thickness of layers were calculated byadding the sample of interest until the signal reached a plateau and remained stable.After changing back to a starting solution the angular spectrum was measured andthe shift of the angle of incidence was determined. By further curve fittings newsurface coverages were detected. After the experiment the tubes and the flow cellwere flushed with a special cleaning solution for at least 10 minutes, separated andthe components of the measurement flow cell were then cleaned with ethanol.

65


2.3 Standard electrochemistry

2.3.1 Materials

2.3.1.1 Solutions

The solutions used for standard electrochemistry were the same as used in Section 2.1.

2.3.1.2 Equipment

• standardized electrodes Pt and Au

• reference electrode

• polishing kit from Basi

• Potentiostat Galvanostat from AUTOLAB

• GPES (General Purpose electrochemical systems) Version 4.9 from AUTOLAB

• hand brace

2.3.2 Method

2.3.2.1 Setup of the standardized measurement cell

Standardized electro-chemistry was used to determine the influence of Surface layerproteins on electrode surfaces during electrochemical measurements of components incomplex solutions. The working station for standard electrochemistry comprised ofa sample tube, the measurment cell, the electrodes with the connection wires to thepotentiostat and the computer providing the necessary software for the electro chem-ical measurements as shown below. The gold and platinum surfaces of the workingelectrodes had defined sizes and were disc-shaped. A Ag/AgCl electrode was usedas reference electrode and stainless steel as counter electrode. The electrode foul-ing event was measured by using cyclic voltammetry. The scan rate was set to 50mV/s, the voltage range averaged between -400 to 600 mV and the number of scanswas set to 10. After every measurement the electrode surfaces were polished using apolishing kit with defined polishing solution and a drill where the working electrodewas chucked. The reference and counter electrodes were washed with water and thereference electrode was stored in a 0.5 M KCl solutions.

66

2.3 Standard electrochemistry

Figure 2.12: Setup of the electrochemical mea-

surement cell: 1) potentiostat; 2)

measurement cell; 3) working elec-

trode either platinum or gold elec-

trode; 4) counter electrode made of

stainless steel; 5) Ag/AgCl refer-

ence electrode

Figure 2.13: Electrochemical measurement cell

with the three electrodes: 1) work-

ing electrode; 2) reference elec-

trode; 3) counter electrode

67

3 Results

3.1 Development of a glucose monitoring

Lab-on-a-Chip

The aim of this work was the developement of a Lab-on-a-Chip device for the contin-uous measurement of blood sugar levels of whole blood samples.The microfluidic chip used for initial experiments contained various electrochemicaldetectors including IDES structures for impedance measurements and band electrodesfor amperometric measurements. The chip was originally designed for multiparame-ter analysis of biofilm growth. Figure 3.1 shows the design of the first Lab-on-a-Chipdevice, the electrode pattern, the design of the fluidic and the assembled device.

Figure 3.1: Components of the chip: a) glass wafer with gold structure, showing the electro-

chemical detectors and the contact patches; b) microfluidic; c) assembled LOC

device.

The electrode configuration in Figure 3.1 a) consisted of three gold electrodes permeasurement position, on two positions of one microchannel. The microfluidic shown

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3 Results

in b) contains a serpentine shaped micromixer and microchambers before and afterthe mixer. The chip shown in Figure 3.1 c) was used for preliminary experiments todetermine if integrated electrochemical measurement are possible on a LOC deviceand to examine online enzyme reaction measurements.

Detection of enzyme-reaction on chip

The first experiment on the Cellmos 2 measurement station was performed to checkif enzyme reactions are detectable on the miniaturized measurement station usingdifferent enzyme and redox mediator concentrations. Furthermore the experiment wasuseful to determine the ideal mediator concentration at optimal enzyme quantities forthe enzymatic assay. The enzyme concentrations valued 10 - 50 - 100 μg/ml. At everyenzyme concentration a measurement series of the electron accepting redox mediatorferricyanide (0.5 - 5 - 10 - 20 - 30 - 40 - 50 - 60 - 70 - 100 mM ferricyanide dilutedin PBS buffer) was measured by adding the substrate glucose in excess (100 mM).The incubation time of every measurement averaged 30 min and was detected usingan external Ag/AgCl counter and reference electrode at a voltage of 350 mV and ameasurement time of 30 s.

Figure 3.2: Current-concentration traces of ferricyanide concentrations. Every curve de-

scribes one measurement series of ferricyanide at a fixed enzyme concentration.

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3.1 Development of a glucose monitoring Lab-on-a-Chip

The graph of Figure 3.2 shows that the measurement system is able to detect glu-cose in a buffered solution using an external counter and reference electrode. Thecurves obtained from the measurement show a sigmoidal current signal, reaching amaximum current at a ferricyanide concentration of 60 mM . The highest glucoseoxidase concentration of 100 μg/ml provided the highest current. An enzyme con-centration of 100 μg/ml and a ferricyanide concentration of 60 mM was used forfurther experiments.

Enlargement of the electrode surface by electrode plating

Electrode plating is a method used to change the surface metal of an electrode andto enlarge the electrode surface resulting in an increase of the measurement system´ssensitivity. The electrode plating experiment was performed using a special elec-trode plating solution (autopring solution) necessary to cover the electrode surfacewith gold. An external gold wire functioned as a working electrode. The platingwas performed for 30 min at a current of 500 nA and a flow rate of 1 μl/min. Achronoamperometric measurement was performed before and after the electrode plat-ing using a 5 mM ferrocyanide solution.

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b)

Figure 3.3: a) Current-time traces of the chronoamperometric measurement of electrode

plating approaches. b) LM-micrograph of the gold chip after the electrode plat-

ing experiment.

The plot of Figure 3.1 a shows that the amperometric signal did not improve afterelectrode plating, indicating that the electrode surface may have been compromised.The examination of the gold chip under the microscope showed that the gold elec-trodes degraded by the plating process, as indicated by the white circles of Figure 3.1b. The red arrow marks the flow direction of the plating solution in the microchannelto point out that the gold degradation occurred on the electrodes and on for the mea-surement independent gold structures that are only exposed to the electrode platingsolution.

Determination of the potential optimum of the band electrodes

Cyclic voltammetry was used to investigate the voltage optimum of the redox reactionof ferri-/ferrocyanide on golden band electrodes. Working, counter and reference elec-trodes consisted of gold. A saline solution with the oxidized and reduced form of the

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redoxsystem (K3Fe(CN)6/K4Fe(CN)6) inclusive 1 M KCl was used as CV solution.The scan rate was chosen of 50 mV/s at a voltage range between -300 to 600 mV .The reactivity of a redox-species can be determined measuring the current at variousvoltages to examine the optimal voltage induction necessary to reduce or oxidize theredox system. Figure 3.4 shows the voltammogram of the reversible electrochemicalreaction.

an. Ep

ip

potential optimum

chosen potential

Figure 3.4: Cyclic voltammogram of golden band electrodes: Current-voltage plot

The fast signal drop after the peak maxima indicates the diffusion dependency ofthe measurement after the discharge of the electroactive molecules near the workingelectrode. The cyclic voltammogram in Figure 3.4 shows that the anodic peak po-tential averages 50 mV and most of ferrocyanide near the electrode is oxidized after150 mV , resulting in a potential optimum of 250 mV , to enable a complete and fastoxidation of ferricyanide to ferrocyanide.

Investigation of the sensitivity and signal stability

Chronoamperometric measurements of ferrocyanide concentrations were performedto determine the stability of electrochemical measurements without an external silverwire as reference electrode. After the inability to cover the gold electrode with silverto create a Ag/AgCl reference and counter electrode, the measurement was performed

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3 Results

using gold as metal of the working, counter and reference electrode. By measuringferrocyanide concentrations of 1 - 0.5 - 0.1 - 0.01 mM , fed by a syringe pump thesignal stability and the sensitivity of the LOC were investigated. Figure 3.5 showsthe graph obtained from the measurement.

Figure 3.5: Current-time traces of a chronoamperometric measurement detecting different

concentrations of ferrocyanide.

The pulsed current-time traces at every concentration in Figure 3.5 show a clearconcentration dependent signal decrease. The current signal showed surface reactions,that can not easily be explained thus indicating a microfluid concentration profile.Concerning the sensitivity of the measurement, current and concentration show nolinear relation. While the measured ferrocyanide concentrations were decreased ten-fold, the current signal did not. The concentrations of 0.5 and 0.1 mM showed highcurrents, approximately 480 nA and were in the same range (only about 50 nA higherthan a concentration of 0.01 mM). The measurement showed high current values forlow concentrations but a low sensitivity concerning different concentrations ( signalof 0,01 - 0.1 - 0.5 mM).

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3.1.1 Simulations with fluidic designs

The chip design was altered for further measurements. The chamber geometries wereremoved and the reactor-channel was narrowed and prolonged to increase the reactiontime between enzyme and substrate. Instead of two separate measurement chambers,the amount was doubled to four separate reactors with a length of 8.4 mm and awidth of 100 μm.The design of the lithographically produced chip was adjusted to its purpose ofchronoamperometric measurements, meaning that the IDES structures were removedand replaced by three band electrodes surrounded by a ground to reduce electroniccross talk. The length of the electrodes is equivalent to the width of the reactor,thus 100 μm and the width of the electrodes averages around 20 μm. The electrodematerial of the microchip used for further materials is platinum. Figure 3.6 showsphotographs of a)the chip containing the electrochemical detectors, b)the microfluidiclayout and c)the assembled microfluidic biochip.

Figure 3.6: Lab-on-a-Chip device designed for glucose monitoring. a) electrochemical detec-

tors and contact pads; b) microfluidic design; c) assembled LOC;

To examine the flow behavior in new fluidic designs, simulations using the softwareCOMSOLE Multiphysics were performed. One characteristic of microfluidic systemsas mentioned in Section1.3.1 are laminar flow-conditions in microfluidics. Laminarflow conditions influence enzyme reactions on chip, most notably if enzyme and tar-get molecule are not premixed outside the chip. The challenge of the LOC device

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is to provide a certain geometry to enable the reaction between glucose oxidase andglucose, resulting in a representative measurement signal. To increase the mixingprofile in the microsystem a serpentine shaped microchannel was used as a passivemicromixer. Figure 3.7 gives a detailed impression of the reactor geometry.

Figure 3.7: Serpentine shape of the microreactor.

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The reactor of Figure 3.7 has a length of 84 mm a width of 100 μm, a height ofabout 30 μm and a total reactorvolume of 0.252 μl. Parameters as the length of thereactor, beside low flow rates provide high reaction times between enzyme and targetmolecules. At a flow rate of 0.1 mm/s per inlet, the velocity distribution is describedby Figure 3.8a.

a)

b)

b)

c)

Figure 3.8: a) b) Velocity field within the microreactor. The color-scale reaches from red

(high velocity) to blue (low velocity); c) Distribution of Re within the microchan-

nel (red indicates high Re and blue low Re)

The blue areas in the rear of the microchannels are caused by coarse mesh sizes andtherefore imprecise calculations. In large blue areas the differential equations weresolved near the channel wall and averaged over larger regions. The representativeresults are the ones of the first four to five channels and can be averaged over the restof the reactor. The window b) enlarges the velocity distribution at the channel turnsand illustrates, that velocity is higher at sharp edges, increasing mixing conditions,which confers the reactor its mixing capacity. The average or mean flow velocity inthe micro channel turned out to be 0.2 mm/s. The Reynols number (Re) behaveslike the fluid velocity and is direct proportional to the velocity-field which can beobtained from Figure 3.8c.The Reynolds number in microreactors averages approximately 0.01 and induceshighly laminar flow conditions. The optimal concentration of K4Fe(CN)6, the me-diator for the amperometric detection of the enzyme reaction was identified to be

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60 mM at an glucose oxidase concentration of 100 μg/ml in the channel, which waschosen for further experiments and checked with additional simulations.The diffusion behavior of K4Fe(CN)6 is illustrated in Figure 3.9a, where the convec-tion and diffusion behavior of fluids with various diffusion constants was calculated.The data for the solvent of ferrocyanide was chosen to equal the parameters of blood(see Table 2.4).

Figure 3.9: Convection and diffusion simulations: a)Diffusion behavior of ferricyanide in the

microchannel. b) Diffusion behavior of glucose oxidase in the microreactor. Red

indicates high concentration, blue low concentrations.

Figure 3.9a indicates a short yellow contact surface between the blue and the redsolution ending up in a green color (average concentration between red and blue),indicating that small molecules like ferricyanide with a diffusion-constant of approx-imately 8 ∗ 10−10 m2/s are immediately merged in the blood sample. A proteinlike glucose oxidase diffuses in contrast to ferricyanide much slower, due to its size,complex structure and hydrophilic and hydrophobic regions. The diffusion constantof glucose oxidase averages 4.5 ∗ 10−11 m2/s in the simulations, adduced from theprotein bovine serum albumin (BSA) with a comparable molecular weight, as listed

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in [89]. The performance of the enzyme in the microfluidic chamber indicates Fig-ure 3.9b. Except the parameters representative for proteins (c,D), the conditions werethe same as the ones of ferricyanide. The enzyme, with a molecular weight of 120kDa needs about 6 turns, which is equivalent to one third of the reactor-length untilit is completely diffused in the blood sample at a mean fluid velocity of 0.2 mm/s

in the reactor. This means furthermore, that the reaction between enzyme (glucoseoxidase), target molecule (glucose) and mediator (ferricyanide) has 4.7 min (velocityis 0.2 mm/s and two thirds of the reactor length 56 mm) to proceed and deliver arepresentative electric signal. In other words, the GOD has sufficient time to reactwith glucose using the present geometry.

3.1.2 Characterization of the microfluidic biochip

The new LOC design in the advanced Cellmos 3 measurement station was optimizedand tested for signal stability at various temperatures and flow rates. Before everymeasurement contaminations in and outside the chip were eliminated and the electriccontact between chip and potentiostat was tested.

Potential

For the fast determination of the potential optimum of the platinum microelectrodescyclic voltammetry was used. The CV solution was the same as the one used forthe examination of the potential optimum of the golden band electrodes: 1 mM

(K3Fe(CN)6/K4Fe(CN)6) diluted in PBS and an additional amount of KCl (1 M).The scan rate was set to 50 mV/s at a voltage range of -300 to 600 mV . Figure 3.4shows the cyclic voltammogram of the measurement.

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an Ep

chosen potential

potential optimum

Figure 3.10: Cyclic voltammogram of gold electrodes: the current is plotted against voltage

The graph in Figure 3.10 shows the fast electron transfer between microelectrodeand redox species, resulting in fast forward and backward reactions. The peak max-imum is reached at approximately 250 mV marked by the red line. After 250 mV

all the ferrocyanide near the electrode is oxidized resulting in a working potential of400 mV in the subsequent experiments to enable a fast and complete oxidation offerrocyanide.

Determination of flow rate influences

The flow rate influences on the measurement signal were examined to determine theoptimal flow rate for a stable current signal. The higher the flow rate the more of theelectroactive molecules are actively transported to the working electrode resulting inan improved chronoamperometric signal. The flow rate influence was tested using a2 mM ferrocyanide solution (diluted in PBS buffer) pumped through the microchipat changing flow rates at a voltage of 400 mV . The effect of flow rates on the sensorsignal is shown in Figure 3.11.

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Figure 3.11: Current-time traces of a chronoamperometric measurement, describing signal

increase by increasing the flow rates.

Increasing flow rates exhibited higher current signals in a non linear correlationbetween velocity and signal increase. The flow rate also positively influenced the sta-bility of the current signal. Higher flow rates caused more stable measurement condi-tions by reducing the diffusion dependency of the measurement and directly leadingferrocyanide to the electrode, thus reducing ionic depletion near the electrode surface.

The effect of temperature on chronoamperometric measurements

Since the physiological temperature of the human body averages 37 ◦C blood sampleshave to be measured at the same temperatures to decrease alternations of the bloodsample during measurements. The temperature stability in the measurement cham-ber and the current dependency on the temperature were investigated in followingexperiment. The temperature was adjusted to the required values using a circulatingwater pump. A 2 mM ferrocyanide solution was fed at a flow rate of 86.7 mm/s anda potential of 400 mV . The thermal water circulating pump was used to heat up thealuminum block to achieve stable temperature-levels. The temperatures investigatedin the experiment were 23 - 37 - 42 ◦C. Figure 3.12 shows the data obtained from

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the measurement.

Figure 3.12: Chronoamperometric measurement: The current is plotted versus time showing

the measurement signal at different temperatures.

The graph in Figure 3.12 shows that the current signals increase in the presence ofhigher temperatures and remained constant at every temperature level. The stablecurrent signal clearly demonstrates that the circulating water pump provides constanttemperature values and proved to be useful for stable temperature adjustment.

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Determination of a signal drift during measurements

During electrochemical measurements ionic depletion can cause signal drifts, disturb-ing the measurement signal and are therefore an object of extensive research [90, 91].To investigate the occurrence and behavior of signal drifts during long time measure-ments a ferrocyanide concentration of 5 mM was measured over 3 hours. Figure 3.13shows the chronoamperometric plot of the long time measurement.

Figure 3.13: Current-time traces of a chronoamperometric measurement of a 5 mM ferro-

cyanide solution.

The plot in Figure 3.13 shows the long time measurement of ferrocyanide. Thecurrent signal was measured every 5 s. The measurement signal shows a stable andreproducible current signal over three hours.

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3 Results

Testing of the sensor performance

To define the measurement range of the glucose sensor, the lower limit of detectionwas investigated by measuring a test series of ferrocyanide. The concentration rangewas chosen to value between 0 - 10 mM , namely 0 - 0.1 - 0.5 - 1 - 10 mM . For theapplication as a glucose measuring device of whole blood samples, the range between3 - 7 mM needs to be covered by the detection range. Every concentration was mea-sured 8 times to determine also the random error at a flow rate of 86.7 mm/s and avoltage of 400 mV . The calibration line is shown in Figure 3.14.

Figure 3.14: Calibration line of the sensor. The charge in nA/h is plotted vs. concentration

of ferrocyanide in mM .

The calibration line is described by a correlation coefficient of R2 = 0.95 and bythe following function:

y = 2.47x− 0.02 (3.1)

y denotes the charge in nA/h and x the concentration of ferrocyanide in mM . Fromthe calibration line can be derived, that ferrocyanide concentrations in the range of0.5 to 10 mM are detectable with a high reproducibility, proven by the coefficient ofvariation of 0.03 to 0.09. The concentration 0.5 mM describes the lower detectionlimit, the upper limit has not been determined.

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3.2 Characterization of the enzymatic reaction


The aim of this section is the testing of the enzyme reaction and the determination ofoptimal reaction conditions, after the examination of the LOC device and the systemsetup. The experimental series start with the determination of the proper mediator,the ideal concentrations and flow rates.

Identification of a proper mediator

Important characteristics of a proper redoxmediator is the reversible oxidation andreduction reaction at rather low electrode potentials to prevent the oxidation of un-wanted interfering species. The mediator needs to posses a high affinity to the enzymeglucose oxidase to receive the remaining electron of the oxidation reaction. The elec-tron transfer-rate has to be high enough to enable an increase of the current signal. Inthe course of this experiment four mediators were chosen, with diversifying numbersof transfer electrons and solubilities. Table 3.1 lists the analyzed mediators and theirperformance during the experiments. The mediators used for the experiment wereextremely heterogeneous molecules, from ferro complexes over aromatic to alicycliccompounds, namely hydroquinone, ferricyanide, 8-hydroxyquinoline, anthraquinoneand menadione. Hydrochinone and anthroquinone offer two transfer electrons andpossess therefore two redox centers, providing an advantage concerning signal trans-duction over the other mediators with only one redox center. Table 3.1 lists the resultsof the mediator performances testings.

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Table 3.1: List of mediators tested for the possibility of further applications and their

properties.

8-hydroxyquinoline and menadionen were not soluble in the selected concentrationrange and were therefore excluded, whereas anthroquinone dissolved partly in thePBS and was furthermore tested. To determine the optimal mediator for glucosemonitoring an open electrochemical system was used, which means that an openPDMS chamber was positioned on the electrodes and formed the open measurementchamber. The concentration of glucose oxidase during the measurement was 100μg/ml and the glucose concentration 10 mM . The chronoamperometric signal wasmeasured after 10 min of incubation for 5 min. The current values of the 5th min

were compared. Hydroquinone had a signal output of 957 nA with a backgroundsignal of 308 nA. Ferricyanide delivered a signal of 297 nA and a background signalof 3.2 nA and anthroquinone 390 nA and a background of 4.8 nA. The influence of themediators on the glucose oxidase activity was tested storing GOD separately in PBSand in the respective mediator solution for 24 h. The next day the enzyme reactions

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were compared again using the open fluidic system. By comparing the signal heightof the premixed enzyme mediator solution and the enzyme mediator sample mixedfor 5 min before the measurement the remaining enzyme activity was calculated.After hydroquinone storage 90 % of the enzyme activity remained, after ferrocyanidetreatment 96 % and after antroquinone storage 79 %. The graph of the experimentis shown in Figure 3.15.

Figure 3.15: Chronoamperometric measurement of enzyme activity. The current is plotted

versus the three mediators on the x-axis. The black quadrangle shows the

measurement signal of the enzyme stored in PBS and the red triangle marks

the signals of the enzyme stored in the respective mediator solution.

The experiment showed that the mediators ferricyanide, anthroquinone, hydro-quinone decrease glucose oxidase activity, but do not cause inactivation after longtime exposure. Beside the semisolubility of anthroquinone in PBS the strong in-activation of the enzyme caused the abdication of anthroquinone as mediator forfurther measurements. Hydroquinone caused the highest current signal but also ahigh background, which outweigh the advantage, because of the high noise-possibilityoccurring with hydroquinone. Although ferricyanide induced the lowest measurementsignal, the low enzyme inactivation and background signal give ferricyanide excellent

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mediator characteristics for glucose monitoring, resulting in the utilization of ferri-cyanide as mediator of the following experiments.

To determine vmax and Km the standardized electrochemical measurement cell wasused. Km also called the Michaelis Menten constant is defined as the substrate concen-tration at which the reaction rate reaches half of its maximum value and is a measurefor the affinity of a certain substrate to the enzyme. The higher the Km value the lessspecific the affinity between enzyme and substrate because more substrate concentra-tion is needed to reach the half maximum reaction rate. vmax defines the substrateconcentration at which the enzyme reaction is at its maximum velocity, meaning thatthe enzyme is constantly reacting and the active center as often as possible occu-pied. These two values are the most important constants for the characterization ofan enzyme and can be obtained via measurement of increasing concentrations of thesubstrate at constant enzyme concentrations. To determine vmax and Km of glucoseoxidase a dilution series of ferricyanide was measured at concentrations of 0.1 - 50 -100 - 200 - 300 - 400 - 641 mM . The enzyme concentration used was 100 μg/ml ata glucose concentration of 700 mM . After an incubation time of 5 min the currentwas measured for 5 min at a voltage of 400 mV . For the kinetic curve the substrateconcentrations were chosen in a broad range. In the kinetic curve the concentration isplotted versus the current of the chronoamperometric measurement, which is shownin Figure 3.16.

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Figure 3.16: a) Plot of the kinetic curve of the enzyme reaction. The current is plotted

versus mediator concentration. b) Hanes plot where substrate concentration in

mM is plotted versus cImM/μA. Km and vmax can be directly obtained from

the straight line.

The plateau of the curve determines the maximal reaction-velocity vmax and theenzyme reaction constant Km at the half-maximal reaction velocity. The Km value inthis particular experiment was calculated to be 79.24mM and vmax as 4.57 μA. Thecorrelation coefficient R2 of the kinetic curve averaged 0.95 and the one of the Hanesplot 0.93 which is an accurate value for curve fittings. By calculating the average Km

and vmax values by comparing the results with subsequent experiments the averagevmax was calculated of approximately 1.929 μA, resulting in an average Km value of60.25 mM .

Optical examination of the micromixer

To determine the ability of mixing in the integrated micromixer, the performanceof the serpentine shaped microfluidic reactor was investigated using an optical ex-periment. Therefore two dyes with different colors (blue and yellow) were used andpumped into the chip at increasing flow rates, to detect the distance until the twocolors were entirely merged. The pictures were taken at every flow rate after achievingflow equilibrium. The pictures taken during the experiments are shown in Figure 3.17.

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a) b) c)

e)d)

Figure 3.17: Mixing of two colors of ink (yellow and blue) in the microfluidic mixer at in-

creasing flow rates.a)-e): 0.07 - 0.7 - 2 - 7 - 20 mm/s

Figure 3.17 illustrates the highly laminar flow conditions in microchannels. a) showsthe mixing at a flow rate of 0.07 mm/s, where the two dyes were mixed after one turn.At a flow rate of 0.7 mm/s the merging of the dyes needed approximately four turnsuntil the two fluids were merged and at a flow velocity of 2 mm/s it took about onethird of the reactor, thus six turns until complete merging was achieved. At a flowrate of 7 mm/s the dye merged after twelve turns and at a flow rate of 20 mm/s thefluids did not merge at all. The information obtained from the mixing experiments iscomparable with the diffusion behavior of glucose oxidase in the microchannel, whilethe diffusion coefficient of the dye in water is comparable to ferricyanide in water, thusabout 8 · 10−10 m2/s proteins like glucose oxidase in water have diffusion coefficientsof about 4.5 · 10−11. Both molecule solution needed the same time to merge in themixing fluid, glucose oxidase in the simulation and the dye in the optical experiment.

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External enzyme reaction

Initial experiments involved an external enzyme reaction to ensure complete substrateconsumption by the GOD and to clearly maximum the current signal. Increasing con-centrations of ferricyanide 0.1 - 0.5 - 1 - 2 - 5 - 10 mM , 10 mM glucose and 100 μg/ml

glucose oxidase were externally incubated (10 min) and then injected into the LOCdevice at a flow rate of 2 mm/s. Figure 3.18 shows the the calibration line exhibitinga range over two orders of magnitude.

Figure 3.18: Current-concentration traces of the chronoamperometric measurement of exter-

nal enzyme reaction at increasing mediator concentrations.

The straight line of the linear fit is defined as:

y = 8.6 ∗ x+ 2.66 (3.2)

with a correlation coefficient R2 of 0.99, which shows a good reproducibility and ac-curacy of the on chip measurement.

Evaluation of the mixing performance of the integrated laminar mixer

A further experiment was performed to investigate the integrated serpentine shapedmicroreactor. Mixing on the LOC device was compared to the mixing performanceof an additional commercial available micromixer. The concentration of enzyme andmediator were again 100 μg/ml and 60 mM , using a glucose concentration of 10 mM .

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Table 3.2 compares the mixing performance of the integrated laminar mixer and thecommercial micromixer. The mixing performance of the external mixer was definedas 100 %.

Table 3.2: Comparison between LOC mixer with and without superposed micromixer

flow rate

[mm/s]

ext. mixer

[nA]

int. mixer

[nA]

capacity of

LOC

[%]

2 29.1 17.7 60.87 28.5 7.0 24.620 19.2 10.1 52.6

Table 3.2 shows that the mixing performance of the integrated micromixer is at60.8 % at a flow rate of 0.2 mm/s which decreased at increasing flow rates. Thelower mixing capacity found at 7 mm/s points at rather low enzyme reaction.

Flow rate optimization for on chip glucose monitoring

For further glucose measurements the optimal flow rate for enzyme reactions, resultingin representative current signals was investigated. An experiment was designed foron chip glucose detection. At an enzyme concentration of 100 μg/ml and a mediatorconcentration of 60 mM a solution of 20 mM glucose was measured. Figure 3.19presents the current signal at increasing flow rates.

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Figure 3.19: Current-time traces of the chronoamperometric glucose measurement, using the

fluidic micromixer. The current was measured at increasing flow rates 0.2 - 0.7

- 1.3 - 2 mm/s.

The plot in Figure 3.19 shows that continuous glucose measurement showed a con-centration profile over the time resulting in the measurement of current peaks. Thepeak maxima decrease with increasing flow rate, caused by a reduction of reactiontime between glucose and glucose oxidase. At a flow velocity of 2 mm/s the peakformation is minimal, resulting in a stable measurement signal of about 11 nA andnearly no glucose detection (background signal are 4 nA). Compared to the currentsignal of the external enzyme reaction and according to the linear Equation 3.2 thecurrent value of a 20 mM ferrocyanide solution induces a current signal of 174.6 nA.The average current value of the last 3 peak maxima of Figure 3.19 at the flow rateof 0.2 mm/s values 177.9 nA, which corresponds to the previous experiment with acurrent difference of 3.3 nA.

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Electrochemical investigation of the analysis of complex protein solu-

tions

Complex measurement samples, containing proteins can lead to instable measure-ments signals during electrochemical measurements. To examine irreversible proteinadsorption on electrode surfaces a protein solution was used for cyclic voltammetricmeasurements. A defined protein solution (30 mg/ml) human serum albumin (HSA)diluted in PBS was used for the examination of protein fouling on the electrodesurface. The electrochemical method cyclic voltammetry was used in a standardizedelectrochemical measurement cell, using both platinum and gold electrodes. The HSAsolution contained a ferrocyanide concentration of 5 mM and was measured with ascan rate of 100 mV/s between -300 and 600 mV .

10th scan

2nd scan

5th scan

Figure 3.20: Current-voltage traces of a cyclic voltammogram, with an enlargement of the

oxidation peaks.

The cyclic voltammogram of Figure 3.20 shows the oxidation and reduction reactionof ferricyanide in a human serum albumin solution. Ten scans were made and thedata obtained from the cyclic voltammetry are listed in Table 3.3.

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Table 3.3: Comparison of anodic peak potential and current using a protein and protein free

solution

solution electrode Eap

[mV ]

iap

[μA]

protein free: pt 300 13.8au 293 13.8

protein: pt 304 11.7au 338 6.4

Table 3.3 shows that during the measurement of protein free samples the oxidationpeak current is at the same level for both platinum and gold working electrodes. Incase of platinum 85 % of the current signals remains after 10 scans, in case of thegold working electrode only 46 %. Platinum is chemically more resistent to electrodefouling compared to gold but in both cases occured protein adsorption, resultingin the necessity of an anti-fouling layer to conduct electrochemical measurements ofcomplex solutions.

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3.3 Anti-fouling strategies to maintain surfaces

In this section anti-fouling surface strategies were investigated to find the ideal surfacemodification preventing protein and electrode fouling. One possibility is the usage ofSurface layer proteins, known as S-layer proteins possessing the ability to form a selfassembled two dimensional monolayer on various surface materials. In this work twodifferent kinds of S-layer proteins, namely SbpA of the microorganism Lysinibacil-lus sphaericus CCM2177 and SbsB of Geobacillus stearothermophilus PV72/p2 wereinvestigated for their ability to form a non-fouling layer in the microchannel of themicrochip and their influence on the sensor signal when crystallized on the electrodesurfaces. Also Polyethylene(glycol) and bovine serum albumine were tested for theirnon-fouling skills and all modifications were compared among one another.

3.3.1 Surface modification with S-layer protein monolayers

3.3.1.1 Crystallization of S-layer proteins under static conditions

To make sure that the surface layer can be formed within the whole reactor of the LOCdevice the S-layer protein was crystallized on every kind of surface material involvedin the μTAS microchannel. The materials of the microchip interfacing with the wholeblood during measurement are glass, the material of the bottom layer, gold and plat-inum the electrode materials and finally PDMS, the polymer used for the formationof the microfluidic. The first approach was the examination of the S-layer proteinSbpA of L. sphaericus CCM2177, that is able to form a lattice structure with squareunits, where each unit consists of four protein subunits assembled to each other. Theexperiment was performed using a SbpA stock solution, consisting of isolated proteinsstored in MilliQ water at 4 ◦C. The 1 mg/ml stock solution was diluted 1 : 10 in0.5 mM TRIS buffer with 10 mM CaCl2. At a final protein concentration of 0.1mg/ml, a high concentration of CaCl2 was required to activate the inducible crystal-lization behavior of SbpA. The probe was kept over night at room temperature andafterwards stored in water at 4 ◦C until AFM imaging. The AFM imaging processwas performed under water with KCl and scanned in constant height contact modeto achieve high atomic resolution. The AFM images are demonstrated in Figure 3.21.

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Figure 3.21: AFM imaging of crystalline SbpA monolayers on various surface-materials,

a)gold, b)platinum, c)glass and d)PDMS

The images of Figure 3.21 were not taken under the same zoom factors, resultingin different dimensions of the developed crystalline structure. Figure 3.21a shows theSbpA crystallization on the gold surfaces. The entire gold surfaces is covered with theS-layer lattice. Figure 3.21b shows the crystallization of SbpA on a platinum (Fig-ure 3.21b), glass (Figure 3.21c) and PDMS (Figure 3.21d) surface. The formation ofsmaller patches is caused by various nucleation sites of crystallization and a subse-quent patches growth until the patches get in contact to the neighboring ones. The

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crystallization process ends with the final fusion of the patches yielding a coherentlayer. The crystalline structure of SbsB on various surfaces is shown in Figure 3.22,the crystallization conditions were the same as with SbpA, but with citrate bufferinstead of CaCl2 in TRIS.

Figure 3.22: AFM imaging of the crystalline lattice structure of SbsB monolayers on a)

PDMS and b) glass surfaces.

Figure 3.22 shows the oblique lattice structure of SbsB on a PDMS (Figure 3.22a)and glass (Figure 3.22b) surface. On platinum and gold surfaces the lattice struc-ture of SbsB could not be identified, indicating that SbsB shows unstable assemblybehavior on this metallic materials.

3.3.1.2 Crystallization of SbpA under fluid flow conditions

After the indication of S-layer formation on various surfaces under static conditions,the assembling process of SbpA was also investigated under fluid flow conditions, toensure an enclosed layer formation in the whole microchannel. The S-layer crystal-lization under fluid flow conditions was investigated using fluorescence labeled anti-bodies. After the crystallization of SbpA in a microchannel at a flow rate of 0.67mm/s, SbpA was marked with anti SbpA IgG. A microchip with an borofloat glassbottom layer and a PDMS microfluidic was fabricated for the antibody labelling ex-

98


periment. Figure 3.23 illustrates the description of the labeling layer arrangement inthe microchannel during the fluorescence labelling.

Figure 3.23: Schematic representation of the antibody labeling experiment: The binding of

an anti SbpA IgG (violet antibodies) after the SbpA crystallization step (yellow

structure on the glass and PDMS surfaces) and the binding of a fluorescence

labeled anti IgG TRITC conjugate (in black with red labels).

After over night crystallization of SbpA, the protein layer was blocked with 3 %lactose PBS buffer for 1 h and afterwards washed with PBS buffer pH 7.2 for 20 min.Firstly the SbpA-layer was incubated with the first antibody (IgG anti SbpA diluted1:10 in PBS buffer pH 7.4) for 1 h at room temperature. The antibody immobilizationon the SbpA was stopped with a subsequent washing step using PBS buffer. Secondlythe IgG anti SbpA layer was incubated with anti IgG TRITC conjugate (secondantibody) for 1 h at room temperature. After gently washing the chip with PBS thefluorescent emission was investigated by inverse fluorescence microscopy. The pictureswere taken using an inverse CFI Achromat DL objective fluorescence microscope witha CFI plan fluor DL 4x objective as shown in Figure 3.24a and a ten fold enlargementshown in Figure 3.24b. To ensure that the S-layer is formed on every surface in themicrochannel, the focal point was altered to check the focal level on the bottom ofthe microchannel, in the middle at the walls of the microchannel and on the top.

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3 Results

a)

b)

increasing focal level

glass bottom PDMS walls PDMS top

Figure 3.24: Fluorescence images of SbpA monolayers. The focal point was varied at 2

different zooming factors presented as a) 4x and b) 10x.

The pictures in Figure 3.24 show various focal points of two enlargements. Thepicture row a) shows the focal point of the bottom (glass) the walls (PDMS) and thetop (PDMS) of the microchannel at a 4 fold enlargement. The picture row of b) showsthe same focal points at a 10 fold enlargement. In the images of Figure 3.24 the redfluorescence is very bright at every focal level, indicating that S-layer protein SbpAwas present on every surface in the microchannel. The labeling experiment provesthat the S-layer protein covers the microchannel surfaces and forms a closed layer.

3.3.2 Investigation of non-fouling properties of S-layer

proteins

3.3.2.1 Examination of S-layer protein assembly

The assembly process was investigated using Surface Plasmon Resonance (SPR), anoptic sensor device using surface plasmons and evanescent waves to examine surfacebinding events, as described in Section 2.2.2.1. SPR is a method to study surfacethicknesses or coverages after adsorption reactions on the sensor surface. For S-layer

100


proteins a refractive index of 1.45 was assumed as recommended in the work of Wolf-gang Knoll and his coworkers [92]. The experiments were performed under roomtemperature using the dilution buffers of the S-layer proteins SbpA and SbsB as ref-erence solutions, assuming a refractive index of approximately 1.333. For SbpA 0.5mM TRIS with 10 mM CaCl2 and for SbsB citrate buffer were used as referencesolutions.

a) b)

Figure 3.25: SPR measurement of SbpA crystallization a) Reflectivity-time traces of the

kinetic measurement of SbpA crystallization; b) Reflectivity-angle traces of the

angular spectrum, with the measurement of the angular spectrum before and

after SbpA crystallization;

Figure 3.25a shows the reflectivity-time plot (kinetic curve) of the SbpA crystal-lization process. The S-layer protein was induced after 10 minutes, resulting in arampant increase of the measurement curve. Most of the assembly process happenedduring the first 20 minutes, while the patch fusion process or the assembly on smallerregions to form a stable and closed S-layer required 50 min. Figure 3.25b shows theangular spectrum before and after SbpA crystallization. The shift of the angle of inci-dence was used to calculate the layer thickness of approximately 8 nm, the thicknessequivalent to the literature [34].Figure 3.26 shows the crystallization process of SbsB. SbsB was diluted in a ratio of1:10 in citrate buffer with a pH of 4.

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3 Results

a) b)

Figure 3.26: Assembly process of SbsB. a) Reflectivity-time traces of the SpsB assembly

process; b) Angular spectrum showing the angle of incidence before and after

assembly.

The reflectivity-time plot of SbsB in Figure 3.26a shows that the assembly processof SbsB on the sensor surface took roughly 180 minutes until the signal remained sta-ble. The main part of the assembly process took approximately 80 minutes and thefinal coverage required 95 min. The angular shift recorded in Figure 3.26b indicatesthat the thickness of SbsB averaged 4 nm, a value equivalent to the literature [34].

3.3.2.2 Investigation of the non-fouling properties of S-layer proteins

using surface plasmon resonance

Surface plasmon resonance was also used to study anti-fouling properties on modifiedsensor surfaces. SPR was used to ensure an enclosed S-layer protein crystallizationon the sensor’s gold surface and the non-fouling ability of modified gold surfaces wasinvestigated. The non-fouling properties of the S-layer proteins SbpA and SbsB wereexamined in the presence of human plasma and human serum albumin (HSA). Toimprove the non-fouling characteristics of SbpA, the S-layer was crosslinked with glu-taraldehyde and BS3. Glutaraldehyde is a crosslinking agent consisting of two highlyreactive aldehyde groups inducing high polymerization properties, by binding to theactive aldehyde group of another glutaraldehyde molecule or to other active groupspresent in the solution. The S-layer was incubated with glutaraldehyde for 30 min.

102


During the glutaraldehyde binding reactive imine groups were formed and needed tobe inactivated with NaBH4 for 50 min.The other crosslinking agent used to stabilize the SbpA layer was Bis(sulfosuccinimidyl)suberate also called BS3. BS3 possesses amino reactive Sulfo-NHS esters on both endsof the molecule. The S-layer was incubated with BS3 for 50 min. To quench unre-acted BS3 the surface was flushed for 30 min with a 1 M TRIS buffer (pH of 7.5).The nonfouling characteristics of S-layers were compared with known anti-fouling sur-face modifications Poly(ethyleneglycol) and bovine serum albumine. Table 3.4 liststhe surface coverages of modified surfaces after the treatment (50 min flushing) witheither human plasma or human serum albumin (30 mg/ml) solution.

Table 3.4: Surface coverages calculated using SPR after the treatment of the modified sur-

faces with protein solutions.

surface layer human plasma

[ng/mm2]

HSA

[ng/mm2]

Au surface 3.74 n.d.SbpA 0.18 +/- 0.16 n.d.SbsB 2.46 1.34SbpA/GA/NaBH4 0.18 n.d.SbpA/BS3 1.02 +/- 0.12 n.d.PEG 0.76 0.41BSA 1.92 +/- 0.16 n.d.

n.d. ... not defined

Table 3.4 lists the surface coverages of modified sensor surfaces after a 50 min

treatment with human serum albumin and with human plasma. The surface coverageof the golden sensor surface flushed for 50 min with human plasma was also measuredas basis value for maximum fouling. After SbpA assembly on the sensor surface, theprotein fouling was investigated and resulted in a 20 fold reduction of protein foul-ing compared to the gold surface. The flushing with human serum albumin solutionshowed no adsorption on the with SbpA modified sensor surface. The second S-layerprotein SbsB halved surface fouling in the presence of human plasma. Compared toSbpA SbsB showed a 14 fold higher amount of protein adsorption. SbsB also showed

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3 Results

HSA adsorption where SbpA did not. The crosslinking of SbpA with glutaraldehydedid not improve the non-fouling performance of SbpA and caused no additional pro-tein adsorption on the S-layer. The modification of the SbpA layer with BS3 evenincreased fouling in the presence of human plasma, resulting in a 6 fold higher amountof protein adsorption compared to SbpA.Poly(ethyleneglycol) as protective nonfouling layer provided a 5 fold reduction of pro-tein adsorption in the presence of human plasma compared to the unmodified goldsurface, but only measured 80 % of SbpA´s non-fouling capacity. PEG also showeda surface coverage of 0.41 ng/mm2 resulting from HSA, which is one third of theamount of SbsB. The modification of the sensor surface with BSA also provided anti-fouling characteristics, compared to the untreated gold surface only half of the proteinamount adsorbed on the surface and compared to SbpA the adsorbed amount was 11times higher.A brief summary of Table 3.4 is that both S-layer proteins SbsB and SbpA showanti-fouling characteristics. The surface modification with SbpA provided the high-est non-fouling performance of all tested non-fouling modifying surface treatmentsand provided higher non-fouling characteristics than with PEG treated surfaces inthe presence of both HSA and human serum. BS3 and glutaraldehyde modifica-tions of SbpA did not improve the non-fouling characteristics, even downgraded thenon-foulng performance (in case of BS3).

3.3.2.3 Examination of S-layers on the electrochemical measurements

The influence of the S-layer protection on electrochemical measurement was investi-gated using a standardized measurement cell with gold and platinum working elec-trodes modified with S-layer protein SbpA of Lysinibacillus spaericus CCM 2177. Aprotein and a protein free solution was used to investigate mediator diffusion throughthe S-layer to the electrode surface and also to examine non-fouling characteristicsduring electrochemical measurements. 5 mM ferricyanide was diluted firstly in aHSA solution (30 mg/ml) and secondly in PBS. The fouling process was investigatedusing cyclic voltammetry at a scan rate of 50 mV/s, in a voltage range between -300and 600 mV . Figure 3.27 shows the cyclic voltammogram of the CV measurement.

104


PBS

HSA

background

Figure 3.27: Cyclic voltammogram of ferricyanide in PBS (black) and ferricyanide in HSA

(blue) using a with S-layer covered platinum electrode.

The same measurement as shown in Figuer 3.27 was performed using an Au workingelectrode. The data of the CV measurements of both electrode materials is summa-rized in Table 3.5.

Table 3.5: Comparison of anodic peak potential and peak current during measurements of

a protein and protein free solution using S-layer modified electrodes

solution material Eap

[mV ]

iap

[μA]

protein free: pt 301 13.8au 311 10.8

protein: pt 305 12.9au 318 9

Table 3.5 shows that the platinum electrode covered with SbpA had similar anodicpeak currents and peak potentials as without modification. Both the anodic peakpotential and the anodic peak current remained constant compared to the measure-ment of unmodified electrodes (compare with Table 3.3). Comparing the electrodeperformance during the ferricyanide detection in the HSA solution and the protein

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3 Results

free solution, the cathodic peak potential showed no significant increase, while theanodic peak current decreased from 13.8 to 12.9 and showed a current change of about0.9 μA. The anodic peak current shift of untreated platinum electrodes averaged 2.1μA (decrease from 13.8 to 11.7 μA), which means that ferricyanide detection in com-plex protein samples can be improved using an additional SbpA-layer. The squarestructure of SbpA induced a reduced fouling process and permitted the diffusion ofthe small mediator molecule to the sensor surface.By comparing treated and untreated golden working electrodes, the CV measurementof a protein free solution showed nearly stable anodic peak potentials. The anodicpeak current decreased from 13.8 to 10.8. indicating a decrease of the electrode sur-face induced by the S-layer. The protein adsorption on the modified gold electrodesis higher, compared to the platinum electrodes where the anodic peak potential in-creased from 311 to 318 mV . The anodic peak current decreased from 10.8 to 9μA, showing a decrease of approximately 1.8 μA, which is twice as high as the peakcurrent decrease on platinum electrodes. The S-layer on the gold electrode reducedthe protein adsorption during the measurement of the HSA solution. Without S-layerthe anodic peak potential rose from 293 to 338 mV (rise of 45 mV ), while the anodicpeak current increased from 311 to 318 mV (rise of 7 mV ) and caused a peak currentreduction of 1.8 μA instead of 7.4 μA in case of SbpA surface modification.Briefly summarized this means that SbpA crystallized on an electrode surface reducedthe protein adsorption during electrochemical analysis in the presence of protein so-lutions and showed a better non-fouling performance on platinum electrodes than ongold electrodes.

106

3.4 Application of the glucose sensing LOC device


The glucose sensing Lab-on-a-Chip device was designed to perform continuous glucosemeasurements during extra corporal blood purification processes. In the last experi-ments the lay out design was tested. The chip is intended to be used as described inFigure 3.28, including one measurement chamber for the blood sample, two reactorsfor a calibration curve using ferrocyanide in one chamber and ferricyanide and glucosein the second chamber. The last reactor is used to measure the background signal,such as blood plasma or blood only to reduce noises.

Figure 3.28: Lab-on-a-Chip device lay out design showing the measurement samples of the

four separate reaction chambers.

The microfluidic chip was designed to enable continuous online glucose measure-ments including auto-calibration and noise reduction. The final experiment was per-formed to proof the measurement principle and to check whether the measurementworks this way. At a flow rate of 0.67 mm/s and a voltage of 400 mV a 10 mM ferro-cyanide solution was measured beside 5 mM glucose, human plasma as background,blood, glucose spiked blood (1 mg/ml) and finally the glucose concentration of humanplasma. For the measurements using on chip enzyme reaction two liked chambers in-

107

3 Results

stead of one reaction chamber were used. The current signals of the measurement isshown in Figure 3.4.

a)

3.5

160.7

3.5

83.5

4.514.7

1 2 3 4 spikedblood

gluc. inplasma{

chamber

b)

Figure 3.29: a) Current time traces of the blood glucose measurement with autocalibra-

tion. b) Bar plot of the average current values plotted versus measurement

chamber/control.

The plot in Figure 3.4a shows stable current signals in a time range of 2 min,

108


in Figure 3.29b the measurement points are shown in a bar plot enabling a betteroverview over the current signal. The red current signal of the 10 mM ferrocyanidemeasurement on chip shows a current of approximately 160.7 +/. 0.3 nA , twicethe signal of the 5 mM glucose concentration with ferricyanide/glucose oxidase ofapproximately 83.5 +/- 1.5 nA. The human plasma background showed a stablecurrent signal at 3.5 +/- 0.09 nA while the glucose measurement in the plasma solutionshowed a mean value of 14.7 nA, which corresponds to a glucose concentration ofabout 0.4 mM glucose in the human plasma. Compared to the glucose concentrationmeasured by a commercially available glucose monitoring device (Accu check) thatmeasured a glucose concentration of approximately 502 mg/dl ≡ to 0.3 mmol/L, theconcentration difference averaged 0.1 mM . The current signals of the spiked andun-spiked blood are in the same current range with only 1 nA difference, showingthat the glucose detection in whole blood is inhibited, either by the inactivation ofthe enzyme or the consumption of the mediator during the measurement.

109

4 Discussion

In this work a glucose sensing Lab-on-a-Chip device was developed, able to contin-uously monitor glucose of human plasma samples. The microchip used for the firstexperiments consisted of two separate reactors both including microfluidic chamberand mixing structures. The electrochemical detectors were IDES structures and bandelectrodes, where the latter structures were ideally suited for preliminary experimentsbecause of the large electrode surface to prove the detection of enzymatic reactionsand the appropriate current range in a microfluidic device. The electrode platingexperiment was performed to increase the electrode surface and showed the oppositeeffect. The reason for the electrode losses is that cyano-complexes in the e-platingsolution reacted with golden structures and induced gold degradation of the litho-graphically unstable gold structures. One possible reason for this assumption is thatgold dissolved from the electrodes in the presence of the plating solution, from theplated electrodes and also gold from surfaces not involved in the plating experiment,only exposed to the plating solution. The gold-degradation is unusual for e-plating so-lutions, indicating weak gold-substrate adhesion properties. The experiment designedto test the sensitivity of the Cellmos 2 measurement station, where ferrocyanide con-centrations of 0.01 - 0.1 - 0.5 - 1 mM were measured showed high current signals butlow sensitivity, obtained from similar measurement signals of highly variating concen-trations. The result of the measurement indicates that large electrode surfaces as incase of band electrodes also cause higher noise ratios, which can be compensated bydecreasing the electrode surface.Generally the first chip design was not ideally suited for glucose measurements withelectrochemical enzymatic detection because chamber structures often trapped bub-bles, disturbing electrochemical detection and the three inlets were not necessary forglucose detection, only provided space for dead volume. The serpentine geometry ofthe first micromixer consisted of 6 turns and showed a lower mixing performance andthe band electrodes with the large surface increase the ratio of noise-detection. Be-

111

4 Discussion

side the inappropriate chip design, the Cellmos 2 measurement station often showedcontact problems leading to unstable measurement signals. The Cellmos 3 measure-ment setup in contrast enables constant measurement conditions, stable temperature,flow rates and measurement signals. The advanced microfluidic chip consists of fourseparate microreactors in serpentine shape with 17 turns and a length of 8.4 cm, idealfor passive mixing on chip. The electrochemical detectors consist of tree electrodesused as working, counter and reference electrode surrounded by a ground. The elec-trodes are made of platinum and the platinum ground surrounds the electrodes andacts as protective electrode shielding to prevent electrode cross talk. The reactor wasdesigned to enable the mixing of the protein glucose oxidase with ferricyanide andglucose under highly laminar flow conditions by the extraordinary reactor length of84 mm, enhancing the diffusion of slowly diffusing molecules as proteins. For highdiffusion capacities the contact surface between the mixable solutions has to be in-creased in the microchannel. Concerning mixing of two fluids in a microfluidic devicethe distance between slowly diffusing molecules (low diffusion constants D in a certainfluid) as proteins and their target molecules have to be decreased, while the contactsurface between two fluids needs to be increased to enable enzymatic reactions. Inrecent years there have been a lot of publications concerning different types of mi-cromixers, with extremely creative ideas of possible geometries, resulting in a highamount of inspiration-possibilities as referred in [93], [94].To enable long time measurements of glucose concentrations in complex measure-ment samples S-layer technology was used to overcome protein adsorption duringelectrochemical measurements. The S-layer with square lattice structures consistingof the S-layer protein SbpA proved themselves an ideal choice as non-fouling surfaces.In general electrode surfaces covered with surface layer proteins show an additionaldiffusion barrier for both small and large molecules, while smaller molecules as fer-ricyanide are able to diffuse through the lattice structure of the SbpA-layer to theelectrode surface. The SPR experiments show no protein fouling in the presence ofHSA only, and low fouling during human plasma treatment also proving that thelattice pores are small enough to prevent HSA diffusion, and allow only diffusion andadsorption of small components of the very complex human plasma. Also the negativecharge of HSA increases the rejection forces of the negatively charged inner surface(N-terminus) of the S-layer. Both crosslinking agents showed no significant improve-ment of the nonfouling characteristics of the S-layer, leading to the assumption that

112

SbpA is stable enough without further treatment. Glutaraldehyde is responsible forthe induction of a net negative surface charge and supports the formation of accessi-ble carboxylic groups, resulting in a slight protein adsorption process. BS3 possessestwo amine-reactive N-hydroxysulfosuccinimide (NHS) ester groups at each end of an8-carbon spacer. This means that the surface charge is not influenced, although thenon-fouling properties were not improved. The SbpA surface is uncharged, which isadvantageous for the non-fouling behavior.SbsB the S-layer protein of Geobacillus stearothermophilus PV72/p2 could not reallycompete with the S-layer protein SbpA from Lysinibacillus sphaericus CCM2177.Human serum albumin proved a high affinity to SbsB surfaces and human plasmashowed an even higher surface coverage of 2.457, which indicates that SbsB modifica-tion seems to be a suboptimal choice as non-fouling surface structure. PEG and BSAwere further surface modifying agents used in the experiments to compare the antifouling property of very popular substances with the ones of SbpA. Compared to theliterature the data obtained from the experiments showed proper high surface cover-age values, but by redoing the experiment, the surfaces coverages remained constant.One possible explanation is the choice of an impure product resulting in a fragilesurface structure. The negative charge of the BSA as non-fouling layer seems to beinappropriate for more complex solutions such as human plasma, because a lot ofsmall cations are present in the sample and form ionic bonds with the free carboxylicgroups. Comparing all non-fouling surfaces, SbpA seems to provide excellent non-fouling properties, resulting in the lowest surface coverage output after the treatmentwith HSA and human plasma, allowing and even improving electrochemical measure-ments.The usage of one reactor for on chip mixing and detection showed pulsed glucosedetection, induced by the vibration of the elastic PDMS polymer of the microfluidic.By using two reactors with a tubing bridge the pulsed detection was eliminated. Toimprove the signal quality the material of the microfluidic has to be changed to aharder polymer, as for example SU8. The first blood sugar measurements of wholeblood samples showed no glucose detection, of spiked (additionally 1 mg/ml glucose)and unspiked blood samples. By measuring the glucose content of a human plasmasample, the glucose concentration was comparable to the same measurement using anoften used glucose sensor for persons with type 1 diabetes (Accu check). The differ-ence between blood and plasma is the lack of erythro- and leukocyes in the plasma

113

4 Discussion

sample. The cells either cause glucose oxidase inhibition or the mediator ferricyanideis involved in an additional reaction and consumed during the measurement. Apartfrom the inability to use whole blood as measurement solution the Lab-on-a-Chipdevice enables accurate glucose measurement in human plasma, with a sensitive au-tocalibration system and background measurements.

114

5 Conlusion

In this work a continuously glucose sensing Lab-on-a-Chip device was developed. Thedevice was designed to be integrated into a dialysis device to measure blood sugarlevels before and after the dialysis process to prevent imprecise glucose administra-tion after the dialysis. The microfluidic of the micro Total Analysis System consistsof four separate microreactors ideally suitable for blood analysis with simultaneouscalibration and noise reduction measurements. The electrochemical setup consistedof the LOC device placed on an aluminum heating block connected with a circu-lating water pump for temperature adjustment, positioned under a light microscopeto accomplish visual contact with the measurement chamber and a syringe pump toinduce the sample solutions. The whole device was connected to a multichannel po-tentiostat. The measurement station proved stable temperature and flow conditionsenabling constant current measurement signals. The S-layer protein SbpA from Ly-

sisnibacillus sphaericus CCM2177 was integrated into the LOC device as non-foulinglayer to prevent protein and electrode fouling. The S-layer assembly in the microchan-nel was investigated to verify complete S-layer coverage, on the channel walls and onthe electrode surfaces, preventing protein fouling on all surfaces, which enabled longtime measurements of complex protein solutions. The glucose sensing device showedprecise glucose detections in plasma solutions with additional autocalibration andnoise reduction measurements. Measurements of whole blood samples induces lowglucose values, possible explanations are either the inactivation of the enzyme glucoseoxidase during the measurement or the reaction of ferricyanide with a blood compo-nent. With further experiments the glucose monitoring in whole blood samples needsto be investigated and optimized.

115

List of Figures

1.1 Dialysis device and position of the LOC device . . . . . . . . . . . . . 3

1.2 3 dimensional structure of glucose oxidase . . . . . . . . . . . . . . . 4

1.3 Flowchart of the enzyme reaction . . . . . . . . . . . . . . . . . . . . 6

1.4 Cyclic voltammogram . . . . . . . . . . . . . . . . . . . . . . . . . . . 8

1.5 Chronoamperometric diagram . . . . . . . . . . . . . . . . . . . . . . 10

1.6 List of state of the art of glucose sensors . . . . . . . . . . . . . . . . 12

1.7 Schematic cycle of electrochemical glucose sensing devices . . . . . . . 16

1.8 Concept of LOC devices . . . . . . . . . . . . . . . . . . . . . . . . . 17

1.9 schema of the continuity equation . . . . . . . . . . . . . . . . . . . . 28

1.10 Effect of shear forces on the flow behavior . . . . . . . . . . . . . . . 30

1.11 Parabolic flow conditions in microchannels . . . . . . . . . . . . . . . 31

1.12 Structure of albumin . . . . . . . . . . . . . . . . . . . . . . . . . . . 34

1.13 Cell patterning experiments . . . . . . . . . . . . . . . . . . . . . . . 35

1.14 PEG structure and applications . . . . . . . . . . . . . . . . . . . . . 36

1.15 S-layer lattice types . . . . . . . . . . . . . . . . . . . . . . . . . . . . 38

1.16 schema of the SUM fabrication . . . . . . . . . . . . . . . . . . . . . 40

2.1 Scheme of the Cellmos 2 measurement station . . . . . . . . . . . . . 46

2.2 Chip placed in the Cellmos 2 fixture . . . . . . . . . . . . . . . . . . 47

2.3 Scheme of the Cellmos 3 LOC station . . . . . . . . . . . . . . . . . . 47

2.4 Chip fixture . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 48

2.5 Schematic description of UV-lithography . . . . . . . . . . . . . . . . 50

2.6 Schema of chip assembly . . . . . . . . . . . . . . . . . . . . . . . . . 53

2.7 Principle of Atomic Force Microscopy . . . . . . . . . . . . . . . . . . 56

2.8 Mesh distribution using extra fine meshing . . . . . . . . . . . . . . . 57

2.9 Arrow and slice plot describing laminar flow behavior . . . . . . . . . 58

2.10 Example of a reflectivity spectrum . . . . . . . . . . . . . . . . . . . 63

117

List of Figures

2.11 Schematic description of SPR measurements . . . . . . . . . . . . . . 642.12 Setup of the electrochemical measurement cell . . . . . . . . . . . . . 672.13 Electrode configuration of the stand. electrodes . . . . . . . . . . . . 67

3.1 Design of the first microchip . . . . . . . . . . . . . . . . . . . . . . . 693.2 Determination of the mediator and enzyme optimum . . . . . . . . . 703.3 Electrode plating experiment . . . . . . . . . . . . . . . . . . . . . . . 723.4 Potential optimum of gold electrodes . . . . . . . . . . . . . . . . . . 733.5 Current-time traces of Cellmos 2 sensitivity . . . . . . . . . . . . . . 743.6 Lab-on-a-Chip device designed for glucose monitoring . . . . . . . . . 753.7 Serpentine shape of the microreactor . . . . . . . . . . . . . . . . . . 763.8 Velocity field within a microreactor . . . . . . . . . . . . . . . . . . . 773.9 Diffusion behavior of ferricyanide and glucose oxidase in a microchannel 783.10 Potential optimum of platinum electrodes . . . . . . . . . . . . . . . . 803.11 Velocity influence on sensor signal . . . . . . . . . . . . . . . . . . . . 813.12 Temperature influence on electrochemical measurements . . . . . . . 823.13 Signal drift chronoamperometric measurements . . . . . . . . . . . . 833.14 Determination of the sensor sensitivity . . . . . . . . . . . . . . . . . 843.15 Mediator influence on the enzyme activity . . . . . . . . . . . . . . . 873.16 Enzyme kinetic: kinetic curve and Hanes plot . . . . . . . . . . . . . 893.17 Optical investigation of mixing performance . . . . . . . . . . . . . . 903.18 Chronoamperometric measurement of an external enzyme reaction . . 913.19 Continuous glucose monitoring on chip . . . . . . . . . . . . . . . . . 933.20 Cyclic voltammogram of electrode fouling . . . . . . . . . . . . . . . . 943.21 AFM imaging of crystalline SbpA structures on various surfaces . . . 973.22 AFM imaging of the crystalline lattice structure of SbsB . . . . . . . 983.23 Schema of the SbpA fluorescence labeling . . . . . . . . . . . . . . . . 993.24 Fluorescence images of the SbpA monolayer . . . . . . . . . . . . . . 1003.25 Kinetic progression of SbpA crystallization . . . . . . . . . . . . . . . 1013.26 Investigation of SbsB assembly by SPR . . . . . . . . . . . . . . . . . 1023.27 Cyclic voltammogram of SbpA covered platinum electrodes . . . . . . 1053.28 Detailed description of the intended purpose of the chip . . . . . . . . 1073.29 Current time traces of the blood glucose measurement . . . . . . . . . 108

118

List of Tables

1.1 4 categories of mediating molecules for enzyme reactions . . . . . . . 71.2 Glucose detection methods . . . . . . . . . . . . . . . . . . . . . . . . 71.3 Lab-on-a-Chip applications . . . . . . . . . . . . . . . . . . . . . . . . 201.4 Lab-on-a-Chip applications part 2 . . . . . . . . . . . . . . . . . . . . 211.5 Lab-on-a-Chip applications part3 . . . . . . . . . . . . . . . . . . . . 221.6 Glucose monitoring on LOC devices . . . . . . . . . . . . . . . . . . . 231.7 Glucose monitoring on LOC devices 2 . . . . . . . . . . . . . . . . . . 241.8 Materials for microfabrications . . . . . . . . . . . . . . . . . . . . . . 251.9 Fluid dynamics in Lab-on-a-Chip systems . . . . . . . . . . . . . . . 271.10 Nonfouling surface modifications . . . . . . . . . . . . . . . . . . . . . 33

2.1 Detailed overview of the solutions used for LOC experiments 1 . . . . 432.2 Detailed overview of the solutions used for LOC experiments 2 . . . . 442.3 Detailed overview of the Equipment used for LOC experiments . . . . 452.4 Parameters of whole blood . . . . . . . . . . . . . . . . . . . . . . . . 582.5 List of solutions used for SPR . . . . . . . . . . . . . . . . . . . . . . 60

3.1 Output of the determination of the optimal mediator . . . . . . . . . 863.2 Comparison of the LOC mixer with an additional external mixer . . . 923.3 Electrode fouling after the measurement of a protein solution . . . . . 953.4 Surface coverages after treatment with protein solutions . . . . . . . . 1033.5 Investigation of electrode fouling on S-layer modified electrode surfaces 105

119

Bibliography

[1] WebMD. Medical Dictionary:third edition. WILEY Publishing, INC., 2007.

[2] Servos W. Breuch, G. Dialyse fuer Einsteiger. Elsevier, 2009.

[3] D. Mueller. Studies on the new enzyme glucoseoxidase. i. BiochemischeZeitschrift, (199):136–170, 1928.

[4] D. S Goodsell. Website, 2010. at http://www.pdb.org/pdb/education_discussion/molecule_of_th

[5] Hartree E.F. Keilin, D. Properties of glucose oxidase (notatin). BiochemicalJournal, 42(2):221–229, 1947.

[6] Ogura Y. Nakamura, S. Kinetic studies on the action of glucose oxidase. Thejournal of Biochemistry, 52:214–220, 1962.

[7] Bright H.J. Weibel, M.K. The glucose oxidase mechanism:interpretation of theph dependence. The journal of Biological Chemistry, 246:2734–2744, 1971.

[8] Turner A.P.F. Wilson, R. Glucose oxidase: An ideal enzyme. Biosensors and

Bioelectronics, 7(3):165–185, 1992. cited By (since 1996) 285.

[9] Bright H.J. Weibel, M.K. The glucose oxidase mechanism. interpretation of theph dependence. Journal of Biological Chemistry, 246(9):2734–2744, 1971. citedBy (since 1996) 112.

[10] Arnold M.A. Alexeeva, N.V. Near-infrared microspectroscopic analysis of ratskin tissue heterogeneity in relation to noninvasive glucose sensing. Journal ofdiabetes science and technology, 3(2):219–232, 2009. cited By (since 1996) 0.

[11] Optical diagnostics and sensing x: Toward point-of-care diagnostics. volume7572, 2010. cited By (since 1996) 0.

121

http://www.pdb.org/pdb/education_discussion/molecule_of_the_month

Bibliography

[12] Dattelbaum J.D. Der, B.S. Construction of a reagentless glucose biosensor usingmolecular exciton luminescence. Analytical Biochemistry, 375(1):132–140, 2008.cited By (since 1996) 8.

[13] Luo H.Q. Li N.B. Bian, X.Z. Electrochemical recognition for sugars onthe chitosan-poly(diallyldimethylammonium chloride)-nano-prussian blue/nano-au/4-mercaptophenylboronic acid modified glassy carbon electrode. Bioprocessand Biosystems Engineering, pages 1–8, 2010. cited By (since 1996) 0; Articlein Press.

[14] Li Q. Xu Y. Ding Y. Wang X. Deng D. Xu Y. Luo, L. Amperometric glucosebiosensor based on nife<sub>2</sub>o<sub>4</sub> nanoparticles and chi-tosan. Sensors and Actuators, B: Chemical, 145(1):293–298, 2010. cited By(since 1996) 0.

[15] Talary M.S. Mueller M. Dewarrat F. Klisic J. Donath M. Heinemann L. Sta-hel W.A. Caduff, A. Non-invasive glucose monitoring in patients with type1 diabetes: A multisensor system combining sensors for dielectric and opticalcharacterisation of skin. Biosensors and Bioelectronics, 24(9):2778–2784, 2009.cited By (since 1996) 3.

[16] Vielstich W. Hamann, C.H. Elektrochemie. WILEY VCH, 2005.

[17] Faulkner L.R. Bard, A.J. Electrochemical Methods. WILEY, 2001.

[18] Toumazou C. Cass A.E.G. Johnston D.G. Oliver, N.S. Glucose sensors: A reviewof current and emerging technology. Diabetic Medicine, 26(3):197–210, 2009.cited By (since 1996) 6.

[19] Silver D.M. Kaplan J. Saudek C.D. Boyne, M.S. Timing of changes in intersti-tial and venous blood glucose measured with a continuous subcutaneous glucosesensor. Diabetes, 52(11):2790–2794, 2003. cited By (since 1996) 76.

[20] Jerman J.H. Angell J.B. Terry, S.C. A gas chromatographic air analyzer fabri-cated on a silicon wafer. IEEE Transactions on Electron Devices, 26(12):1880–1886, 1979. cited By (since 1996) 318.

122

Bibliography

[21] Kurt E. Petersen. Fabrication of an integrated, planar silicon ink-jet structure.IEEE Transactions on Electron Devices, ED-26(12):1918–1920, 1979. cited By(since 1996) 16.

[22] Graber N. Widmer H.M. Manz, A. Miniaturized total chemical analysis systems:A novel concept for chemical sensing. Sensors and Actuators: B. Chemical, 1(1-6):244–248, 1990. cited By (since 1996) 1044.

[23] Loke W.K. Tan Y.T. Nguyen N.-T. Tan, H.Y. A lab-on-a-chip for detection ofnerve agent sarin in blood. Lab on a Chip - Miniaturisation for Chemistry and

Biology, 8(6):885–891, 2008. cited By (since 1996) 6.

[24] Ahn C.H. Do, J. A polymer lab-on-a-chip for magnetic immunoassay with on-chip sampling and detection capabilities. Lab on a Chip - Miniaturisation forChemistry and Biology, 8(4):542–549, 2008. cited By (since 1996) 11.

[25] Whitesides G.M. Weibel, D.B. Applications of microfluidics in chemical biology.Current Opinion in Chemical Biology, 10(6):584–591, 2006. cited By (since 1996)54.

[26] Verboven P. Vermeir S. Vergauwe N.-NicolaÃ¯ B. Lammertyn J. Atalay, Y.T.Modeling and optimization of a multi-enzyme electrokinetically driven mul-tiplexed microchip for simultaneous detection of sugars. Microfluidics and

Nanofluidics, 7(3):393–406, 2009. cited By (since 1996) 0.

[27] Meagher R.J. Albarghouthi M.N. Barron A.E. Doherty, E.A.S. Microchannelwall coatings for protein separations by capillary and chip electrophoresis. Elec-trophoresis, 24(1-2):34–54, 2003. cited By (since 1996) 142.

[28] Jobst G. Urban G.A. Moser, I. Biosensor arrays for simultaneous measurementof glucose, lactate, glutamate, and glutamine. Biosensors and Bioelectronics,17(4):297–302, 2002. cited By (since 1996) 65.

[29] Chatrathi M.P. Tian B. Wang, J. Microseparation chips for performing multien-zymatic dehydrogenase/oxidase assays: Simultaneous electrochemical measure-ment of ethanol and glucose. Analytical Chemistry, 73(6):1296–1300, 2001. citedBy (since 1996) 64.

123

Bibliography

[30] Lieberzeit P.A. Richter L. Schirhagl R.-Milnera M. Dickert F.L. Bailey A. Ertl P.Birnbaumer, G.M. Detection of viruses with molecularly imprinted polymersintegrated on a microfluidic biochip using contact-less dielectric microsensors.Lab on a Chip - Miniaturisation for Chemistry and Biology, 9(24):3549–3556,2009. cited By (since 1996) 1.

[31] Chen-Ta Ho, Ruei-Zeng Lin, Wen-Yu Chang, Hwan-You Chang, and Cheng-Hsien Liu. Rapid heterogeneous liver-cell on-chip patterning via the enhancedfield-induced dielectrophoresis trap. LAB ON A CHIP, 6(6):724–734, JUN 2006.

[32] Rasooly A. Herold, K.E. Lab on a Chip Technology; Volume 1: Fabrication andMicrofluidics. Caister Academic Press, 1992.

[33] Jensen H. Gamby J. Bai X.-Girault H.H. Wu, Z. A flexible sample introductionmethod for polymer microfluidic chips using a push/pull pressure pump. Lab on

a Chip - Miniaturisation for Chemistry and Biology, 4(5):512–515, 2004. citedBy (since 1996) 14.

[34] Mirkin C.A Niemeyer, C. Nanobiotechnology. Wiley-VCH, 2005.

[35] Enzelberger M. Quake S. Liu, J. A nanoliter rotary device for polymerase chainreaction. Electrophoresis, 23(10):1531–1536, 2002. cited By (since 1996) 145.

[36] Roussel Jr. T.J. Crain M.M. Bathlagunda-V. Jackson D.J. Gullapalli J. ConklinJ.A. Pai R. Naber J.F. Walsh K.M. Keynton R.S. Baldwin, R.P. Fully inte-grated on-chip electrochemical detection for capillary electrophoresis in a micro-fabricated device. Analytical Chemistry, 74(15):3690–3697, 2002. cited By (since1996) 87.

[37] Schild L. Renaud Ph. Gawad, S. Micromachined impedance spectroscopy flowcytometer for cell analysis and particle sizing. Lab on a Chip - Minituarization

for Chemistry and Biology, 1(1):76–82, 2001. cited By (since 1996) 184.

[38] Nojima T. Fujii T. Yamamoto, T. Pdms-glass hybrid microreactor array withembedded temperature control device. application to cell-free protein synthesis.Lab on a Chip - Miniaturisation for Chemistry and Biology, 2(4):197–202, 2002.cited By (since 1996) 50.

124

Bibliography

[39] Girault H.H. Rossier, J.S. Enzyme linked immunosorbent assay on a microchipwith electrochemical detection. Lab on a Chip - Minituarization for Chemistryand Biology, 1(2):153–157, 2001. cited By (since 1996) 99.

[40] Hayashi K. Fan X. Yamamoto K.-Kato T. Niwa O. Kurita, R. Microfluidicdevice integrated with pre-reactor and dual enzyme-modified microelectrodes formonitoring in vivo glucose and lactate. Sensors and Actuators, B: Chemical,87(2):296–303, 2002. cited By (since 1996) 31.

[41] Maekawa E. Okubo F. Mizutani F.-Yao T. Sekizawa R. Hisamoto H. Henares,T.G. Enzyme-release capillary as a facile enzymatic biosensing part for acapillary-assembled microchip. Analytical Sciences, 25(8):1025–1028, 2009. citedBy (since 1996) 0.

[42] Yun Seok Heo, Lourdes M. Cabrera, Jonathan W. Song, Nobuyuki Futai, Yi-Chung Tung, Gary D. Smith, and Shuichi Takayama. Characterization andresolution of evaporation-mediated osmolality shifts that constrain microfluidiccell culture in poly(dimethylsiloxane) devices. ANALYTICAL CHEMISTRY,79(3):1126–1134, FEB 1 2007.

[43] Michael W. Toepke and David J. Beebe. PDMS absorption of small moleculesand consequences in microfluidic applications. LAB ON A CHIP, 6(12):1484–1486, DEC 2006.

[44] R. Mukhopadhyay. When pdms isn’t the best. Analytical Chemistry, 79(9):3249–3253, 2007. cited By (since 1996) 7.

[45] Grodrian A. Kremin C. Hoffmann M.-Metze J. Schumacher, J.T. Hydrophobiccoating of microfluidic chips structured by su-8 polymer for segmented flow op-eration. Journal of Micromechanics and Microengineering, 18(5), 2008. cited By(since 1996) 1.

[46] Pai J.-H. Lai H.-H.-Sims C.E. Bachman M. Li G.P. Allbritton N.L. Wang, Y. Sur-face graft polymerization of su-8 for bio-mems applications. Journal of Microme-

chanics and Microengineering, 17(7):1371–1380, 2007. cited By (since 1996) 7.

[47] Franssila S. Tuomikoski, S. Free-standing su-8 microfluidic chips by adhesivebonding and release etching. Sensors and Actuators, A: Physical, 120(2):408–415, 2005. cited By (since 1996) 36.

125

Bibliography

[48] Tan K.L.-Hong G.D. Yang-L.J.-Gong H.Q. Zhang, J. Polymerization optimiza-tion of su-8 photoresist and its applications in microfluidic systems and mems.Journal of Micromechanics and Microengineering, 11(1):20–26, 2001. cited By(since 1996) 92.

[49] Reisner W. Stein-D. Del Bonis-O’Donnell, J.T. Pressure-driven dna transportacross an artificial nanotopography. New Journal of Physics, 11, 2009. cited By(since 1996) 0.

[50] Pidaparti R.M. Su, G. Transport of drug particles in micropumps through novelactuation. Microsystem Technologies, 16(4):595–606, 2010. cited By (since 1996)0.

[51] Dandy D.S. Lynn, N.S. Passive microfluidic pumping using coupled capil-lary/evaporation effects. Lab on a Chip - Miniaturisation for Chemistry andBiology, 9(23):3422–3429, 2009. cited By (since 1996) 1.

[52] Heinisch M. Hilber W. Jakoby-B. Lederer, T. Electromagnetic membrane-pumpwith an integrated magnetic yoke. pages 532–537, 2009. cited By (since 1996) 0.

[53] Yeung A. Nandakumar K. Jain, M. Efficient micromixing using induced-chargeelectroosmosis. Journal of Microelectromechanical Systems, 18(2):376–384, 2009.cited By (since 1996) 1.

[54] Xu G. Tng Y. Htet-K.Z.-Yang C. Ying J.Y. Lee, D. Large distance liquid pump-ing by ac electro-osmosis for the delivery of biological cells and reagents in mi-crofluidic devices. volume 6836, 2008. cited By (since 1996) 0.

[55] E. Schilling. Website, 2010. at http://faculty.washington.edu.

[56] Quadir M.A. Sharma S.K. Haag-R. Calderon, M. Dendritic polyglycerols forbiomedical applications. Advanced Materials, 22(2):190–218, 2010. cited By(since 1996) 4.

[57] Lee K.D. Pyo K.B. Park-S.Y.-Lee H. Lee, H. Catechol-grafted poly(ethyleneglycol) for pegylation on versatile substrates. Langmuir : the ACS journal of

surfaces and colloids, 26(6):3790–3793, 2010. cited By (since 1996) 0.

126

http://faculty.washington.edu

Bibliography

[58] Kinsel G.R. Peng, L. Improving the sensitivity of matrix-assisted laser desorp-tion/ionization (maldi) mass spectrometry by using polyethylene glycol modifiedpolyurethane maldi target. Analytical Biochemistry, 400(1):56–60, 2010. citedBy (since 1996) 0.

[59] Katira P. Hess H. Agarwal, A. Quantifying and understanding protein adsorptionto non-fouling surfaces. 2010. cited By (since 1996) 0.

[60] Li G. Xue H. Yang-W.-Zhang F. Jiang S. Gao, C. Functionalizable and ultra-lowfouling zwitterionic surfaces via adhesive mussel mimetic linkages. Biomaterials,31(7):1486–1492, 2010. cited By (since 1996) 2.

[61] Hou X. Holmberg, M. Competitive protein adsorption of albumin and im-munoglobulin g from human serum onto polymer surfaces. Langmuir, 26(2):938–942, 2010. cited By (since 1996) 0.

[62] J.M. Harris. Poly(Ethylene Glycol) Chemistry: Biotechnical and Biomedical Ap-

plications. Kluwer Academic Publisher, 1992.

[63] He Y. Bernards M.T. Jiang-S. Hower, J.C. Understanding the nonfouling mech-anism of surfaces through molecular simulations of sugar-based self-assembledmonolayers. Journal of Chemical Physics, 125(21), 2006. cited By (since 1996)11.

[64] Bernards M.T. Chen S. Tsao-H.-K. Sheng Y.-J. Jiang S. Hower, J.C. Hydrationof "nonfouling" functional groups. Journal of Physical Chemistry B, 113(1):197–201, 2009. cited By (since 1996) 0.

[65] Raghavan S. Tan J.L. Chen-C.S. Nelson, C.M. Degradation of micropatternedsurfaces by cell-dependent and -independent processes. Langmuir, 19(5):1493–1499, 2003. cited By (since 1996) 54.

[66] Xue H. Li G. Jiang-S. Cheng, G. Integrated antimicrobial and nonfouling hy-drogels to inhibit the growth of planktonic bacterial cells and keep the surfaceclean. Langmuir, 26(13):10425–10428, 2010. cited By (since 1996) 0.

[67] Chu W.-L. Chen W.-Y.-Zheng J. Liu-L. Ruaan R.-C. Higuchi A. Chang, Y. Asystematic spr study of human plasma protein adsorption behavior on the con-trolled surface packing of self-assembled poly(ethylene oxide) triblock copolymer

127

Bibliography

surfaces. Journal of Biomedical Materials Research - Part A, 93(1):400–408,2010. cited By (since 1996) 0.

[68] Sara M. Weigert, S. Ultrafiltration membranes prepared from crystalline bac-terial cell surface layers as model systems for studying the influence of surfaceproperties on protein adsorption. Journal of Membrane Science, 121(2):185–196,1996. cited By (since 1996) 17.

[69] Kane R.-Chen C.S. Ingber-D.E.-Whitesides G.M. Ostuni, E. Patterning mam-malian cells using elastomeric membranes. Langmuir, 16(20):7811–7819, 2000.cited By (since 1996) 147.

[70] Conconi M.T.-Gambaretto R. Pasquato-A.-Folin M. Bello-C.D. Parnigotto P.P.Dettin, M. In vitro generation of differentiated cardiac myofibers on micropat-terned laminin surfaces. Journal of Biomedical Materials Research, 60(3):472–479, 2002. cited By (since 1996) 45.

[71] Brunner S.-Dini T. Payet-V.-Galtayries A. Marcus-P. Frateur, I. Adsorption ofalbumin and fibronectin on model surfaces of biomaterials [adsorption d’albumineet de fibronectine sur surfaces modÃ¨les de biomatÃ c©riaux]. Materiaux et Tech-

niques, 98(1):41–48, 2010. cited By (since 1996) 0.

[72] Vu X.T.-Gilles S. Mayer-D.-Offenhaeusser A. Ingebrandt-S. GhoshMoulick, R.Impedimetric detection of covalently attached biomolecules on field-effect tran-sistors. Physica Status Solidi (A) Applications and Materials, 206(3):417–425,2009. cited By (since 1996) 1.

[73] Fane-A.G. Kim, K.-J. Performance evaluation of surface hydrophilized novelultrafiltration membranes using aqueous proteins. Journal of Membrane Science,99(2):149–162, 1995. cited By (since 1996) 31.

[74] Fullerton M.S.-Nielsen S.H. Jorgensen-L. Baldursdottir, S.G. Adsorption of pro-teins at the oil/water interface-observation of protein adsorption by interfacialshear stress measurements. Colloids and Surfaces B: Biointerfaces, 79(1):41–46,2010. cited By (since 1996) 0.

[75] J.R. Peters. Advances in Protein Chemistry. Academic Press INC, 1975.

128

Bibliography

[76] O’Sullivan B.-McArthur S.L. Wright-P.C. Salim, M. Characterization of fib-rinogen adsorption onto glass microcapillary surfaces by elisa. Lab on a Chip- Miniaturisation for Chemistry and Biology, 7(1):64–70, 2007. cited By (since1996) 9.

[77] Svenneby G. Wedege, E. Effects of the blocking agents bovine serum albuminand tween 20 in different buffers on immunoblotting of brain proteins and markerproteins. Journal of Immunological Methods, 88(2):233–237, 1986. cited By (since1996) 6.

[78] Michel-R. Voros J.-Textor-M. Hofer R.-Rossi A. Elbert-D.L. Hubbell J.A.Spencer N.D. Huang, N.-P. Poly(l-lysine)-g-poly(ethylene glycol) layers on metaloxide surfaces: surface-analytical characterization and resistance to serum andfibrinogen adsorption. Langmuir, 17(2):489–498, 2001. cited By (since 1996) 225.

[79] Wagner E. Boeckle, S. Optimizing targeted gene delivery: Chemical modificationof viral vectors and synthesis of artificial virus vector systems. AAPS Journal,8(4), 2006. cited By (since 1996) 13.

[80] Bradoo S.-Saxena R.K. Gupta, R. Aqueous two-phase systems: An attrac-tive technology for downstream processing of biomolecules. Current Science,77(4):520–523, 1999. cited By (since 1996) 3.

[81] Harris-J.Milton Dust, Julian M. Picryl ether of polyethylene glycol. preparation,characterization, and reaction with n-propylamine. Journal of Polymer Science,

Part A: Polymer Chemistry, 28(7):1875–1886, 1990. cited By (since 1996) 5.

[82] Herati Reza S. Harris, J.Milton. Synthesis of polyethylene glycol thiol. volume 32,pages 154–155, 1991. cited By (since 1996) 6.

[83] Yang W. Zhang Z. Cao-Z.-Cheng G. Piliarik-M. Homola J.-Jiang S. Vaisocherova,H. Ultralow fouling and functionalizable surface chemistry based on a zwitteri-onic polymer enabling sensitive and specific protein detection in undiluted bloodplasma. Analytical Chemistry, 80(20):7894–7901, 2008. cited By (since 1996) 14.

[84] Sleytr U.B. Sara, M. Crystalline bacterial cell surface layers (s-layers): Fromcell structure to biomimetics. Progress in Biophysics and Molecular Biology,65(1-2):83–111, 1996. cited By (since 1996) 31.

129

Bibliography

[85] Messner P. Pum D. Saara-M. Sleytr, U.B. Crystalline bacterial cell surface layers(s layers): From supramolecular cell structure to biomimetics and nanotechnol-ogy. Angewandte Chemie - International Edition, 38(8):1034–1054, 1999. citedBy (since 1996) 124.

[86] H. Plattner. Electron Microscopy of Subcellular Dynamics. CRC Press, Inc.,1989.

[87] Sleytr U.B. Sara, M. Biotechnology and biomimetic with crystalline bacterialcell surface layers (s-layers). Micron, 27(2):141–156, 1996. cited By (since 1996)20.

[88] Sleytr U.B. Sara, M. Comparative studies of s-layer proteins from bacillusstearothermophilus strains expressed during growth in continuous culture un-der oxygen-limited and non-oxygen-limited conditions. Journal of Bacteriology,176(23):7182–7189, 1994. cited By (since 1996) 28.

[89] Li P. Giddings J.C. Liu, M.-K. Rapid protein separation and diffusion coeffi-cient measurement by frit inlet flow field-flow fractionation. Protein Science,2(9):1520–1531, 1993. cited By (since 1996) 36.

[90] Gangopadhyay S. Temkin H. Pu-Q.-Liu S. Datta, A. Nanofluidic channels byanodic bonding of amorphous silicon to glass to study ion-accumulation andion-depletion effect. Talanta, 68(3):659–665, 2006. cited By (since 1996) 12.

[91] Yun J. Temkin H. Liu-S. Pu, Q. Ion-enrichment and ion-depletion effect ofnanochannel structures. Nano Letters, 4(6):1099–1103, 2004. cited By (since1996) 115.

[92] Naumann R. Friedrich M.Robertson J.W.F.-Loesche-M. Heinrich F.-McGillivrayD.J. Schuster-B. Gufler P.C. Pum D. Sleytr U.B. Knoll, W. Solid supported lipidmembranes: New concepts for the biomimetic functionalization of solid surfaces.Biointerphases, 3(2):125–135, 2008.

[93] McKnight T.E. Ramsey J.M. Jacobson, S.C. Microfluidic devices for electrokinet-ically driven parallel and serial mixing. Analytical Chemistry, 71(20):4455–4459,1999. cited By (since 1996) 150.

130

Bibliography

[94] Wu Z. Nguyen, N.-T. Micromixers - a review. Journal of Micromechanics and

Microengineering, 15(2):R1–R16, 2005. cited By (since 1996) 279.

131

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