POTENTIAL FOR ENHANCING BIOCOMPATIBILITY THROUGH

MICROSTRUCTURING AND ANTI-COAGULATING BIOMOLECULAR COATINGS

by

TINA TALWAR

(Under the Direction of William Kisaalita)

ABSTRACT

To demonstrate the potential for increasing implantable sensor life span by sensor surface

microstructuring, microwells were fabricated into SU-8(an epoxy based negative photoresist

material). The microstructured surfaces were modified with heparinized medical grade

polyurethane and the biocompatibility of the surfaces was evaluated with respect to protein

adsorption and porcine platelet adhesion. With a commercial implantable sensor surface

(MINIMED), as a control, the microstructured/ heparinized surfaces offered more resistance

(p=0.05) to protein and cell attachment, suggesting potential for increasing implantable sensor

lifespans through microstructuring.

INDEX WORDS: Biofouling, Biosensor, Cell-material interaction, Microstructuring, SU-8

POTENTIAL FOR ENHANCING BIOCOMPATIBILITY THROUGH

MICROSTRUCTURING AND ANTI-COAGULATING BIOMOLECULAR COATINGS

by

Tina Talwar

B.E., Dr. Babasaheb University, India, 1998

A Thesis Submitted to the Graduate Faculty of The University of Georgia in

Partial Fulfillment of the Requirements for the Degree

MASTER OF SCIENCE

ATHENS, GEORGIA

2012

© 2012

Tina Talwar

All Rights Reserved

POTENTIAL FOR ENHANCING BIOCOMPATIBILITY THROUGH

MICROSTRUCTURING AND ANTI-COAGULATING BIOMOLECULAR COATINGS

by

TINA TALWAR

Major Professor: William S. Kisaalita

Committee: Yiping Zhao

William Tollner

Electronic Version Approved:

Maureen Grasso

Dean of the Graduate School

The University of Georgia

May 2012

iv

ACKNOWLEDGEMENTS

I express gratitude to Dr. William Kisaalita, Dr. Robert Brown, Ju Rong and the Department of

Veterinary Science for their support in my academic career.

v

TABLE OF CONTENTS

CHAPTER 1: Implantable Glucose Sensors................................................................................... 1

1. 1 Introduction .......................................................................................................................... 1 1.2 Objectives ............................................................................................................................. 3 References ................................................................................................................................... 4

CHAPTER 2: Biocompatible Polymers and Anticoagulants .......................................................... 6

2. 1 Literature Review................................................................................................................. 6 2.2 Biocompatible Polymers - Advantages and Disadvantages .................................................. 9

2.3 Heparin and PU Mechanism ............................................................................................... 12 2.4 Anticoagulant ...................................................................................................................... 12

2.5 PU Synthesis ....................................................................................................................... 13 References ................................................................................................................................. 14

CHAPTER 3: Polymer Selection with Quartz Crystal Microbalance (QCM) ............................. 18

3.1 Abstract ............................................................................................................................... 18 3.2 Introduction ......................................................................................................................... 19

3.3 Polymers evaluated ............................................................................................................. 20 3.4 Polymer selection ................................................................................................................ 22 References ................................................................................................................................. 25

CHAPTER 4: Potential for enhancing Biocompatibility through microstructuring and anti-

coagulating biomolecular coating ................................................................................................. 27 4.1 Abstract ............................................................................................................................... 28 4.2 Introduction ......................................................................................................................... 29

4.3 Materials and Methods ........................................................................................................ 30 4.4 SEM and X-ray analysis ..................................................................................................... 34

4.5 Results and Discussion ....................................................................................................... 34 4.6 Concluding remarks ............................................................................................................ 37

References ................................................................................................................................. 38 Acknowledgements ................................................................................................................... 42

CHAPTER 5: Conclusion and Future Directions ......................................................................... 48 5.1 Discussion of results ........................................................................................................... 48

5.2 Future Directions ................................................................................................................ 49 APPENDIX 1: PU synthesis and modification with pendant acetylated thiol groups ................. 51 APPENDIX 2: PU and heparin binding........................................................................................ 53

APPENDIX 3: SU-8 Coating and Micro patterning ..................................................................... 55 APPENDIX 4: PPP, PRP extraction and Cell Counting .............................................................. 60 APPENDIX 5: Cell Adhesion on SU-8 micro wells .................................................................... 62 APPENDIX 6: PU spin coating parameters.................................................................................. 63

1

CHAPTER 1: Implantable Glucose Sensors

1. 1 Introduction

The design of in vivo glucose biosensors for clinical purpose remains a significant challenge

mostly due to poor biocompatibility. When a sensor comes into contact with blood, it provokes a

defensive reaction of the blood. The human body regards any sensor implanted either

subcutaneously or placed in the blood as foreign and tends to reject it by fouling the sensor

surface. Implantable sensors are fouled by protein adsorption followed by platelet adhesion that

change the cells‟ morphology and activate them; leading to formation of thrombus [1]. Thrombin

effectuates conversion of fibrinogen into fibrin, an insoluble mass of strands. The activated

platelets and fibrin form thrombus emboli. When these thrombi detach from the sensor it causes

infection and form a biofilm or a scar tissue. This fouling of sensor negatively impacts the long-

term utility of the sensor by reducing glucose diffusion to the sensor which results in decreased

current output. This leads to either partial or complete malfunction of the sensor calling for

frequent replacement, i. e. every 2-7 days [2, 3]. Although, reliable sensor performance in vitro

has been reported, a continuous glucose monitoring in vivo is still in experimental stages.

In all the studies mentioned here sensor performance has been achieved mainly by polymer

coatings such as polyurethane and anti coagulants such as heparin. Amperometric sensors based

on glucose oxidase have come a long way since their introduction in 1962 by Clark and Lyons.

In 1995, needle type sensors with a „long‟ life span of 25 days was reported. The longest lifespan

2

of an implantable electrochemical type of glucose sensors under clinical trials is about 4 -7 days

only and in vitro stability stretching to about 56 days. The long term stability of such

polyurethanes (PU)-epoxy in bovine serum is reported to be 6 months [4]. In 2000, Yang et al.,

reported a PU as a diffusion limiting membrane with stable output in bovine serum for 70 h and

linearity up to 50 mM. Another study reported the use of bulk and surface heparinization with in

vivo efficacy ranging from a week to a few months [5]. Polyurethanes (PU) containing Ag have

shown lowered foreign body reaction for up to 19 days [6], which is still not sufficient for long

term usage in vivo. Another factor compromising life of the sensor is a decline in enzyme

activity over time. For long-term application, a fully implanted glucose sensor that works

reliably for a month to at least six months is desirable. A possible approach in increasing the

lifespan of the sensors may be achieved by coating them with anti-biofouling and anti-

coagulating materials [7] and providing a constant supply of enzyme [8]. All the above

mentioned studies have tackled either protein or cell attachments issues separately,

demonstrating potential for increasing in vivo life of sensors, but have not achieved the desired

long life span of the sensor yet. The present study to explore increasing the life span of

amperometric type glucose sensors is based on using multi layered structure of protein resistant

PU polymer and anti thrombin, heparin. Of the many available biocompatible polymers, we

chose to use PU for its unique ability to combine with heparin with ease. To achieve the above

mentioned in vivo life span we proposed to couple PU polymer and antithrombin heparin coating

with SU-8 surface microstructuring. This unique combination showed resistance against both

protein and cell attachments, in turn demonstrating the potential (in vitro) of this combination

with respect to increasing the in vivo life span. Further, follow up studies will functionalize the

structures with GOD-conjugated gold nanoparticles to verify longer lifespan in vivo. The intent

3

behind this study is to provide SU-8 microstructured surfaces coated with biocompatible PU

polymer and anti thrombin heparin to demonstrate potential to increase the life spans of

subcutaneously implanted glucose sensors over 6 months.

1.2 Objectives

Glucose sensors are coated with different biocompatible copolymers to resist protein

adsorption and with anti coagulants such as heparin, chitosan, coumarine to prevent biofouling

and in turn to optimize the lifespan. In heparin coated sensors, heparin gets leached out and the

anti thrombic properties degrade over time. Thus the major problem of cell and protein adhesion

leading to sensor failure prevails. Several studies have focused on protein or cell attachments

issues separately but have not dealt with the two issues together. In our research we suggest

using heparin modified PUs and SU-8 surface microstructuring for enhancing the lifespan of the

sensors.

The SU-8 microstructure in the sensor combats the cell adherence problem while the

heparinized polyurethane (PU + Hep) coating on this microstructure avoids protein adsorption,

with heparin working as an anti coagulant. Thus with this novel combination we propose to

demonstrate the potential for increasing the lifespan of glucose sensor through the following

specific objectives:

Objective 1: Identify a biocompatible polymer that is not only protein resistant but also binds

well with heparin without running the risk of losing heparin over time.

Objective 2: Determine the optimal size of the microwells that is the diameter ratio of the

microwells and the cells. The micro well size should be such that the platelets (cells) will be

prevented from entering the microwell and attaching onto the sensor surface within. A photo

4

resist material- SU-8 is used as micro structure substrate in obtaining the optimal microwell

diameter.

References

1. Sung W J, Na K, Bae YH. Biocompatibility and interference eliminating property of pollulan

acetate/PEG/heparin membrane for the outer layer of an amperometric glucose sensor. Sensors

and Actuators. 2004. 99(2-3): 393-398.

2. Jin W, Brennan JD. Properties and applications of proteins encapsulated within sol-gel derived

materials. Analytica Chimica Acta. 2002. 461(1): 1-36.

3. Moussay F, Harrison DJ, O‟Brien DW, Rajotte RV. Performance of subcutaneously implanted

needle type glucose sensors for employing a novel trilayer coating. Anal.Chem. 1993. 65(15):

2072-2077.

4. Yu B, Long N, Moussy Y, Moussy F. A long term flexible minimally –invasive implantable

glucose biosensor based on epoxy-enhanced polyurethane membrane. Biosensors and

Bioelectronisc. 2005. 21(12): 2275-2282.

5. Michanetzis GPA, Katsala N, Missirlis YF. Comparison of haemocompatibility improvement

of four polymeric biomaterials by two heparinization techniques. Biomaterials. 2002. 24(4): 677-

688.

6. Chou CW, Hsu SH, Chang H, Tseng SM, Lin HR. Enhanced thermal and mechanical

properties and biostability of polyurethane containing silver nanoparticles. Polymer degradation

and stability. 2005. 91(5): 1017-1024.

7. Wickramasinghe Y, Yang Y, Spencer SA. Current problems and potential techniques in In

Vivo glucose monitoring. Journal of Fluorescence, 2004. 14(5): 513-520.

5

8. Abel PU, Woedtke T, Schulz B, Bergann T, Schwock A. Stability of immobilized enzymes as

biosensors for continuous application in vitro and in vivo. Journal of molecular Catalysis B:

Enzymatic. 1999. 7(1-4): 93-100.

6

CHAPTER 2: Biocompatible Polymers and Anticoagulants

2. 1 Literature Review

Diabetes mellitus is a disease characterized by endocrine metabolic disorder. The body‟s

inability to produce sufficient amount of insulin that regulates blood sugar leads to elevated

glucose levels. Diabetes affects about a million people each year in the US alone, no wonder it

has become a major health concern around the world. The normal blood sugar physiological

range is 110 +/-25 mg/dL, anything above or below this range is considered abnormal.

Hypoglycemia (low sugar level) can cause mental confusions, convulsions, coma and even

death. Whereas Hyperglycemia, a condition characterized by high sugar levels, can cause many

long term neuropathic and micro vascular disorders including blindness due to high levels of

protein [1, 2].

There are various types of glucose sensors based on different measurement principles. They

are broadly classified into invasive and non-invasive. The non-invasive type based on NIR

spectroscopy has low sensitivity and poor selectivity caused by NIR absorption by body

chemicals other than glucose and the study [1] focuses on semi-invasive, potentially implantable

type of enzymatic electrochemical glucose sensor, where determination of blood glucose is a

direct consequence of the chemical reaction taking place at the transducer-analyte interface (as

represented by the equations (1)-(6) below). The construction of amperometric needle-type

glucose sensor is a simple three-electrode system, consisting of the working enzyme electrode, a

platinum counter electrode and a silver/silver chloride reference electrode. In the amperometric

type of glucose biosensor, the enzyme glucose oxidase (GOD)

7

catalyses the oxidation of β-D-glucose by molecular oxygen producing gluconolactone and

hydrogen peroxide. During enzymatic oxidation of glucose by GOD, the cofactor flavin-adenine

dinucleotide (FAD) is reduced to FADH2, followed by oxidation of the enzyme co-factor

(regeneration of the bio-catalyst) with formation of H2O2. The reactions are expressed as follows:

β-D-glucose + GOD (FAD) → glucono-δ-lactone + GOD (FADH2 )…………….. (1)

GOD (FADH2) + O2 → GOD (FAD) + H2O2 ……………………………………. (2)

Glucono-δ-lactone + H2O → gluconic acid …………………………………….. (3)

The gluconolactone (reaction 1) is hydrolysed (reaction 2) in aqueous media to gluconic acid

(reaction 3). The complete reaction can be summarized as:

β-D-glucose + O2 + H2O → gluconic acid + H2O2 …………………………….. (4)

The amount of glucose is determined by measuring anodic current of oxidation of hydrogen

peroxide produced as:

GOD

Glucose + O2----------------- Gluconic acid + H2O2 ……………………………. (5)

The formation of hydrogen peroxide is determined by the amperometric current during electrode

oxidation: H2O2 = O2 + 2H+

+ 2e- …………………………………..…………………. (6)

The conversion of glucose to gluconic acid involves transfer of two protons and two electrons

from substrate to the enzyme. The electron transfer from redox cofactor (GOD) to the sensing

electrode is facilitated by the polymer coatings.

The amperometric sensor measures current that results from the oxidation or reduction of

electroactive compounds (H2O2). The current is linearly proportional up to 500 mg/dl of the

concentration of glucose. The combined advantage of linearity and high selectivity makes

amperometric type of sensors widely preferred [1, 3, 4, and 5]. However, a difficulty

8

encountered in these types of sensors along with cell and protein attachment, is the direct electro-

oxidation of organic body chemicals like ascorbic acid, uric acid, p-acetaminophen, that give rise

to interference signal. This interfering signal compromises the selectivity of the sensor. The

enzyme activity decreases over a period of time and leads to low output signal. This drawback

combined with poor biocompatibility of the sensor still remains an area of research and a

solution to the drawback is pre-requisite for long-term sensor use. Here in an effort to improve

the long term performance of the sensor using different biocompatible polymers by correcting

the problem of electro-oxidation of physiological fluids other than glucose, an anti-coagulating

and permselective biomaterial such as heparin is employed.

It is well known that surfaces of the implanted sensor are fouled by protein adsorption,

platelet activation and adherence and formation of thrombus causing scar tissue formation at the

site of implantation Proteins rapidly accumulate to available solid-liquid interfaces. This

adsorption changes protein‟s conformation, which causes formation of gel-like layer of

denatured protein. This gel-like layer causes membrane fouling and fibrous encapsulation. This

is followed by inflammation, wherein the leukocytes from the blood stream enter tissues and

activate the macrophages which further trigger cytokines to develop chronic inflammation and

foul the implant site. It is also well known that more cells adhere on a protein layer. These

deleterious processes lead to membrane biodegradation and finally sensor failure in vivo. Earlier,

steps were taken to minimize this effect by encapsulating the sensor with various cellular

components, but that led to alteration of mass transport of the glucose to the sensor surface and

increase of the lag time response. Also it was observed that due to protein clogging the surface

permeability was reduced [6]. Thus a need to create anti-fouling and anti-coagulating membranes

to address the problem of signal drift and permeability change arose. To address this need,

9

sensors are now surface modified and coated with special biodegradable anti-biofouling

copolymers that reject protein adsorption, but still fail to give practical long term usage in vivo.

Our work is geared toward tackling the problem of protein adsorption and clot formation using

multi-layered polymer coated sensor along with permselective and anti-coagulant biomaterial,

heparin, and bringing in micro structuring to lengthen the life span of implantable glucose

sensors. We envision that this study of biocompatible, protein resistant and anti coagulant

polymers would further benefit continuous glucose monitoring systems wherein our sensor could

be connected to automatic insulin pumps mimicking pancreatic functions.

2.2 Biocompatible Polymers - Advantages and Disadvantages

The following protein resistant polymers were considered for sensor coating. A review of the

advantages and disadvantages of the short listed polymers are discussed below:

Polyethylene glycol (PEG): PEG has been extensively used as a biocompatible coating due to

its non- immunogenic and nontoxic and non-antigenic behavior. It is known that PEG repels both

protein and cells and thus inhibits acute thrombosis by controlling protein fouling. The efficacy

of ultra thin PEG (MW = 1000 Da) films used on silicon based micro devices is satisfactory up

to a period of 4 weeks [1, 7]. But stability is an issue in ultra thin PEG films (5.10 + 2.21 A0 for

0.5% PEG concentration), which can be overcome by increasing the film thickness and

immobilization time (32.5 + 1.41 A0 for 1% concentration and 120 min immobilization time).

The literature suggests that a low (0.5%) PEG concentration fails to provide a well defined film

[8, 9]. But again, if the concentration and film thickness are increased, the desirable feature of

nano/ micro devices to have homogeneous and ultra thin coatings is compromised.

The thickness,„d‟ of PEG film is calculated by the following equation:

10

= d ρo ……………………………………………………………………… (7)

Where, (in g/cm2) is the surface concentration of PEG film and ρo is the density of

crystalline film.

PEG tends to hydrate even under controlled dry conditions and affects the film thickness.

Ellipsometric measurements show that PEG thickness is stable at room temperature (25 0C) but

decreases at body temperature (37 0C). It also poses as a restriction of being minimally exposed

to oxygen and light during long term use. Moreover, the PEG coupling methods are lengthy

complex synthesis routes needing expensive equipments. Although alkanethiol terminated (O-

EGn) PEG modified surfaces significantly improve the protein resistance [10, 11], it is limited to

gold substrates only and cannot be used for BioMEMS applications where silicon platforms are

used.

Polypyrrole (Ppy): The versatility of another conducting biodegradable and biocompatible

material, polypyrrole (Ppy) has been exploited to the fullest since the early 90s. Apart from its

ability to abort protein adhesion and block interfering electroactive anions responsible for

background current, polypyrrole can be easily polymerized electrochemically with various

dopents and deposited on any electrode surface [12]. Such doped pyrrole (GOD/Ppy/Pt)

improves the electron transfer and sensitivity (330 nA/ mM cm2) of the sensor and gives a

response time of only 20 seconds [13]. Ppy allows well fixed immobilization of bio-catalyst

(GOD) for enzyme based sensors and does not hinder in its bioactivities. Such GOD based Ppy

have been used in continuous glucose monitoring. Although this type of sensor is simple in the

sense that the pyrrole is unsubstituted, reproducible, small with precise localization of enzyme, it

still faces the problem of instability. The carboxyl and hydroxyl based pyrrole (3-(1-pyrrolyl)

propionic acid (PPA) and 3-(1-pyrrolyl) propanol) too face limited life- spans [14].

11

Polyurethane (PU): Literature analysis reveals that biodegradable fibers made of PU/

(polyetherurethane) PEU/(polyurethane ureas) PUUR with latent thiol groups have high tensile

strength and high modulus (3MPa) and so Tecoflex and Tecothane with pendant acetylthio

groups were used in our research. PUUR are multi block copolymers composed of repeated soft,

usually a polyether or a polyester diol and hard segment based on reaction of diisocyanante and a

chain-extended diamine. These elastomers have the semi crystalline hard phase dispersed in the

soft viscous matrix. By incorporating hydrolysable linkages into the polymer, purposely

degradable PUs can be synthesized in a variety of ways. They exhibit a wide range of properties

from being very brittle and hard materials to soft tacky ones, due to the options available in

selecting the chemistry and molecular weights of the various components, and the ratios in which

they are induced in the polymer. These possibilities together with the fact that polyurethanes

have excellent physical and mechanical properties and excellent biocompatibility in a variety of

applications have favored their use and development as biomaterials in various blood contacting

devices [15, 16 and 17].

When polyurethane is used as the outermost protective layer, it shows extended linearity up to

50 mM and stability up to 70 hrs [18]. Tissue culture on PUUR fibers have shown a remarkable

compatibility for up to 2 years in vitro and showed no sign of chronic inflammation or foreign

body reactions [16]. Thus, based on the excellent glucose diffusion-limiting behavior, protein

resistant nature and long term biocompatibility, we proposed to use PU as the outermost layer in

our glucose sensor construction.

Poly Vinyl Butyral (PVB): PVB too satisfies the criteria of biocompatibility and prevention of

protein transfer, however it is water insoluble and leads to excessive swelling of the base

membrane.

12

2.3 Heparin and PU Mechanism

When blood comes into contact with a polymer surface plasma proteins, platelets get

deposited on the surface. This is followed by formation, growth and detachment of thrombi.

Thrombosis depends on surface energy charge, hydrophobicity or hydrophilicity, polarity,

surface roughness and composition of the polymeric surface. All these factors become crucial in

determining the design of any biocompatible biomaterial. All the previous work indicate that the

red blood cells, platelets, normal vascular endothelium carry heavy negative charge and as such

the natural blood vessel is resistant to thrombosis due to repulsion of these negative charges.

Thus polymeric layers with the introduction of negative-zeta potential, like heparin, are not only

blood compatible but also anticoagulant. The anticoagulant nature of heparin is due to sulphate

and aminosulphate groups on the heparin molecule. Heparin is a naturally occurring

polysaccharide that can ionically bind to a polymer (in our study, PU). It passivates the PU

surface and makes it biocompatible [19]. For blood contacting applications, polyurethanes have

indicated relatively better thromboresistance. Furthermore, incorporating ions of heparin in PU

result in increasing the tensile strength and toughness of PU and help in changing normally

hydrophobic PU, hydrophilic, yielding thromboresistant biomaterial system [19]. The PUs can be

surface modified by grafting or immobilizing heparin into it or by modifying the surface texture.

It is known that a rough surface is more blood compatible as it promotes neointima propagation

and further enhances hydrophobic nature of PU which resists protein adsorption at the surface.

2.4 Anticoagulant

Heparin: There are many biological anticoagulants available, like TM (thrombomodulin),

prostacyclin, prostaglandin, urokinase, heparin etc. Heparin was discovered by Mclean in 1916

and was called Antithrombin III (now simply AT), because it requires a plasma cofactor for its

13

anticoagulant activity. Only approximately one third of an administered dose of heparin binds to

AT, and this fraction is responsible for most of its anticoagulant effect. The remaining two thirds

has minimal anticoagulant activity at therapeutic concentrations, but at concentrations greater

than those usually obtained clinically, both high- and low-affinity heparin catalyze the AT effect

of a second plasma protein, heparin cofactor II. Anti-coagulant and biocompatible biomolecules

such as heparin that inhibit clot formation and other coagulating proteases are used as outer

membrane coatings for increasing the heamocompatibility. It is a naturally occurring potent

polysaccharide that interacts with antithrombin III to inhibit fibrin clot. The polysaccharide

chains compose of repeating units of D-glucosamine and either L-iduronic or D-glucoronic acids

[10, 17 and 20]. Earlier work [22] states that when heparin is immobilized into PU, it provides

long term antithrombogenicity vs. heparin releasing system, where although heparin‟s bioactivity

is at a higher level, it is suitable for short term only. The aim of resembling the

nonthrombogenicity of the endothelial cell (EC) layer lining the inner wall of the healthy blood

vessels can thus be achieved by heparanizing PU surface and enhancing the sensor‟s life.

Heparin is also known to maintain the complete blood count (CBC) [23].

2.5 PU Synthesis

The outer lipid layer of natural red blood cell membrane is highly biocompatible with other

blood components and it consists of about 80% of phosphorylcholine head groups. This puts

forward an idea of constructing a heamocompatible synthetic polymeric membrane that mimics

the natural cell membrane. An analysis of literature related to PU synthesis tells us that, 2-

methacryloyloxyethyl phosphorylcholine (MPC) copolymerized with n-butyl methacrylate (poly

(MPC-co-BMA)) in conjunction with polyurethane would significantly improve the sensor

performance [1, 18, 19 and 24].

14

Another driving factor in using PU and heparin composite is that PU is susceptible to

oxidative biodegradation; especially due to highly oxidative H2O2. Addition of heparin modifies

the chemical structure of PU and provides stability against oxidation [25].

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2. Kortesuo P, Ahola M, Kangas M, Yli-Urpo A, Kiesvaara J, Marvola M. In vitro release of

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Journal of Controlled Release. 1987. 5(1): 23-36.

6. Wickramasinghe Y, Yang Y, Spencer SA. Current problems and potential techniques in In

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15

7. Sung WJ, Na K, Bae YH. Biocompatibility and interference eliminating property of pollulan

acetate/PEG/heparin membrane for the outer layer of an amperometric glucose sensor. Sensors

and Actuators. 2004. 99(2-3): 393-398.

8. Sharma S, Johnson RW, Desai TA. Ultrathin poly (ethylene glycol) films for silicon-based

microdevices. Applied Surface science. 2003. 206(1-4): 218-229.

9. Menz B, Knerr R, Göpferich A, Steinem C. Impedance and QCM analysis of the protein

resistance of self assembled PEGylated alkanethiol layers on gold. Biomaterials. 2004. 26(20):

4237-43.

10. Zhu A, Lu P, Wu H. Immobilization of poly(caprolactone)_poly(ethylene oxide)-

poly(caprolactone) triblock copolymer on poly(lactide-co-glycolide) surface and dual

biofunctional effects. Applied Surface Science. 2007. 253(6): 3247-3253.

11. Lan S, Veiseh M, Zhang M. Surface modification of silicon surfaces for improved

biocompatibility and cell patterning selectivity. 2004. Biosensors and Bioelectronics. 20(9):

1697-1708.

12. Ramanavicius A, Ramanaviciene A, Malinauskas A. Electrochemical sensors based on

polymer- polypyrrole. Electrochimica Acta. 2006. 51(27): 6025-6037.

13. Ma M, Qu L, Shi G. Glucose oxidase electrodes based on microstructured polypyrrole films.

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14. Yasuzawa M, Nieda T, Hirano T, Kunugi A. Properties of glucose sensors based on the

immobilization of glucose oxidase in N-substituted polypyrrole film. Sensors and Actuators.

2000. 66(1-3): 77-79.

15. Gisselfalt K, Edberg B, Flodin P. Synthesis and properties of degradable Poly(urethane

ureas) to be used for ligament reconstructions. Biomacromolecules. 2002. 3(5): 951-958.

16

16. Gretzer C, Gisselfalt K, Liljensten E, Ryden L, Thomsen P. Adhesion, apoptosis and

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18. Yang Y, Zhang SF, Kingston MA, Jones G, Wright G, Spencer SA. Glucose sensor with

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20. Lin WC, Liu TY, Yang MC. Hemocompatibility of polyacrylonitrile dialysis membrane

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DOCA conjugate dispersed as nanoparticles in polyurethane film. Biomaterials. 2000. 22(3):

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23. Kung FC, Chou WL, Yang MC. In Vitro evaluation of cellulose acetate hemodialyzer

immobilized with heparin. 2006. Polym. Adv. Technol. 17(6): 453-462.

24. Alferiev IS, Fishbein I. Activated polyurethane modified with latent thiol groups.

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17

25. Hsu SH, Chou C W. Enhanced biostability of polyurethane containing gold nanoparticles.

Polymer Degradation and Stability. 2004. 85(1): 675-680.

18

CHAPTER 3: Polymer Selection with Quartz Crystal Microbalance (QCM)

Tina Talwar1,4

, William Kisaalita1,

*, Yan Geng2 and Yiping Zhao

3

1Cellular Bioengineering Laboratory, Faculty of Engineering, University of Georgia, Athens, GA

30602, USA

2Department of Chemistry, University of Georgia, Athens, GA 30602, USA

3Department of Physics and Astronomy, University of Georgia, Athens, GA 30602, USA

4Present Address: Applied Medical, 22872 Avenida Empresa, Rancho Santa Margarita, CA,

92688

*To whom correspondence should be addressed. Tel.: +1 706 542 0835; Fax: +1 706 542 8806.

E-mail: ttalwar@engr.uga.edu, williamk@engr.uga.edu, ygeng@chem.uga.edu ,

zhaoy@physast.uga.edu

3.1 Abstract

Biopolymer selection for the purpose of enhancing the biocompatibility of implantable sensors

was an important aspect of our study. Along with the requirement of high protein resistance, an

important criterion in the biopolymer selection was its ease of binding with heparin- an

anticoagulant. The principle of quartz crystal microbalance (QCM) wherein the change in

frequency directly relates to amount of protein deposited on quartz, was used to compare five

different polymers. A two way ANOVA was conducted on the polymers‟ result. The type of

polymer and the protein concentration were considered fixed and the frequency change was

considered as a dependent variable. Higher amounts of protein deposited resulted in higher

19

frequency change, applying the grade of polymer in terms of protein resistance. Since the t- test

performed on the data confirmed that all the polymers offer significant resistance to protein as

compared to no coating at all (α = 0.05); the choice of using one type of polymer finally

depended on ease of heparin immobilization. For this reason, polyurethane was picked.

Keywords: QCM, Polyehethylene glycol (PEG), Poly Vinyl Butaryl (PVB), Polypyrrole (Ppy),

polyurethane (PU)

3.2 Introduction

Implantable sensors are coated with polymers to resist protein adsorption and prevent

biofouling to optimize the sensors‟ life spans. Given the many possible polymers to choose from,

it was necessary to evaluate several candidates with respect to minimal protein adsorption. Five

polymers including: polypyrrole (Ppy, 600 nm thick); Ppy (30 nm thick); polyethylene glycol

(PEG); polyurethane (PU); and poly vinyl butaryl (PVB) were selected. A detailed review of

these biocompatible polymers is given in Chapter 2. The Quartz Crystal (QCM) technique was

used to assess protein adsorption. Sauerbrey (1959) pioneered the use of QCM technology in

calculating mass loadings. Since then QCM has been used as a mass and thickness monitors in

gas phase and thin film depositions. The QCM device is made up of built-in frequency counter,

resistance meter, and the crystal oscillator that operates with 5 MHz crystals. QCM works well

for uniform and thin film deposits for basic surface-molecule electrochemical studies. With

QCM the incremental change in mass (and hence frequency) comes from foreign film (here

protein) deposition and adds to the original thickness of the underlying quartz. The sensor crystal

is an inch in diameter, thin disk of 5 MHz, AT-cut (the quartz is a thin plate and is cut at 35° 15´

20

to the optic axis of the crystal), α-quartz with circular gold electrodes patterned on both sides.

These types of crystals are known for their excellent mechanical and piezoelectric properties.

The external film deposition is considered rigid and thin enough not to be affected by shear

forces during vibration [1], [2] and [3].

The extremely sensitive response of a piezoelectric device toward mass changes at the surface

of the QCM electrodes was applied and the piezoelectric response was noted for mass changes at

different concentrations of Bovine Serum Albumin (BSA) for the quartz coated with different

polymers. With increasing concentrations of BSA in direct contact with quartz crystal, a higher

mass change per unit area at the QCM electrode surface was observed that was signaled as a

change in oscillation frequency of the crystal. Thus, the Δf (frequency change) translated into

corresponding protein deposition on quartz coated with polymers. A Student t-test was used to

compare polymer performance with no coating as the control. All polymers tested performed

satisfactorily.

3.3 Polymers evaluated

All the chemicals were obtained from Sigma-Aldrich, unless otherwise stated. QCM and

gold quartz crystals, flow cell, were obtained from SRS (Stanford Research Systems, Sunnyvale,

CA). The sensor crystals were handled carefully and in a way that did not harm the electrode

pads and the coatings. The crystals were coated (Fig. 4.1) with the mentioned polymers and

although the sensors are reusable, traces of consumption and wear were checked for prior to each

use. The crystals were discarded for any signs of cracks, metal peeling or discoloration due to

buffer, as the measurements become erroneous.

21

Pyrrole: Pyrrole (0.1M Pyrrole in 0.1M KCl at 7.45 pH), at 30 and 600 nm thickness was

coated using electrochemical method. The crystals were washed thoroughly with DI water and

ethanol and dried under N2 (g) prior to each use. A current (347.62µ amps) for 10 s and 200 s

was applied to produce 30 nm and 600 nm of Pyrrole thickness respectively. We used

polypyrrole (Ppy) film prepared in KCl aqueous solution instead of acetonitrile solution as

normally done, to obtain a larger surface area of the polymer [4].

PEG 5000: PEG (SH = 114, MW = 5000) was purchased from Sigma Aldrich. Before

surface treatment, the gold crystals were cleaned with DI water at least 3 times and then dried

under nitrogen. The crystal was immersed and kept immersed in the solution (0.1 mM PEG 5000

Da in 10 ml ethanol) for 3 days. Then it was washed with ethanol and dried under N2 (g) before

each use [5]. This technique forms a thin and uniform layer of the polymer.

Polyurethane: For preliminary experiments, 10 mg of Tecoflex PU (Thermedics) was

dissolved in 10 ml of chloroform and stirred for an hour to dissolve the polymer completely.

Two drops of the solution were spin coated on quartz crystal for 2 minutes at 5000 rpm. This

procedure was repeated twice for a more effective layer-by-layer electro deposition. The crystal

was washed with DI water prior to each use.

PVB: A dense 2 % w/vol PVB solution was prepared by dissolving PVB in ethanol and

spin coated on quartz crystal for 2 minutes at 5000 rpm, two times for better anchorage of the

polymer on quartz crystal, ensuring smooth and uniform coverage on the gold surface.

All the QCM experiments were conducted at room temperature. PBS (phosphate buffer

saline) was allowed to flow (0.1ml/ min) over crystal for at least an hour to stabilize and achieve

a stable baseline of frequency for at least 2 min. Care was taken to remove all air bubbles prior to

22

running the flow meter. Bovine Serum Albumin (BSA) at different concentrations (0.3 mg/ml,

3.0 mg/ml, and 30 mg/ml, details in Table 4.1) was injected into the flow cell to flow over

crystals coated with different polymers (as described in Chapter 2). Depending upon the

concentration of BSA and the type of prepared surface; change in resonance frequency (∆f) as a

function of protein adsorption was noted.

3.4 Polymer selection

The evaluated biocompatible polymers, PEG, PU, PVB and Ppy (at two thicknesses of 600

nm and 30 nm), were uniformly coated on the quartz crystals using spin coating and electro

deposition methods as described above. The coated crystals were inspected for any scratches or

peeling prior to each use. Figure 4.1 show examples of SEM images illustrating acceptable

uniform (a) and unacceptable nonuniform (b and c) coatings. The frequency drop (∆f) in

response to different concentrations of BSA at was translated into the amount of protein

adsorbed on the crystal (Table 4.1). Crystal samples were exposed to BSA at concentrations of

0.3 mg/ml, 3 mg/ml and 30 mg/ and the corresponding change (drop) in frequency with protein

mass deposition was recorded. At the lower most BSA concentration of 0.3 mg/ml, all polymers

resisted protein equally. At 3.0 mg/ml, Ppy (600 nm thickness) offered more resistance than PU.

However, 600 nm coating is mot practical in light of sensor miniaturization trends. Table 4.1 also

shows that a higher thickness of pyrrole (600 nm) resists protein adsorption more than the lower

thickness (30 nm) of pyrrole. A Student t-test was performed and at 0.3 mg/ ml of protein

concentration, all polymers (p < 0.05) but PVB (p = 0.08) significantly resisted protein. At 3.0

mg/ ml of protein concentration, all polymers (p < 0.05) with PU (p = 0.05) significantly resisted

protein. At 30 mg/ ml of protein concentration, all polymers (p < 0.05) but Ppy 30 nm (p = 0.06)

significantly resisted protein. This suggests that the selected polymers are equally good in terms

23

of protein resistance and it is a matter of picking the polymers that binds well with heparin (at the

time of short-listing the biopolymers, an anticoagulant “heparin” incorporation in the research

was not yet considered).

Thus, QCM analysis confirmed the selected biopolymers for their resistance to protein

adsorption- an important consideration for implantable sensor design. Studies with 2-

methacryloyloxyethyl phosphorylcholine (MCP) blended in PU, have reported for PU‟s surface

blood compatibility properties. Various MCP polymers copolymerized with cyclohexyl

methacrylate or 2-ethylhexyl methacrylate (EHMA) into Tecoflex60 using the same solvent have

revealed reduced platelet deposition on PU-MCP membranes [6, 7]. Thus, PU for its ease of

heparin immobilization and biostability seemed to be an obvious choice of all the preselected

biopolymers, as a protein resistant coating for sensor.

The simple relationship between changes in frequency (∆f) and mass (∆m) enables QCM to be

widely used in sensing applications. All the polymers evaluated are well known biocompatible

polymers used in many biosensors, BIOMEMS and implants. Since, QCM is very sensitive to

even minor depositions on quartz surface; care was taken to uniformly coat the polymers with no

defects. Different coating techniques such as electro deposition (Ppy), casting (PEG) and spin

coating (PVB and PU) were tried for effective coating on quartz. Different polymers showed

different affinity for protein and with higher protein concentration a higher change in frequency

(∆f) was noted indicating substantial protein mass adherence on the quartz surface.

Besides being biocompatible, all the selected polymers were easy to prepare and coat, were

stable and could be customized with additives or dopants. Since our study also involved an

anticoagulant, PU was selected for further study (Chapter 4), because of its excellent ability to

bind with heparin with minimum leaching allowance.

24

Figure 3.1 SEM images of three separate quartz crystals coated with PU (a), PVB (b) and PPy (c). a)

illustrated unacceptable nonuniform coating in comparison to acceptable uniform coatings in b) and

c).

a b c

25

Table 3.1 Frequency change at varying BSA concentrations with different polymers coated on

Quartz Crystal

References

1. Sauerbrey G. Z. Phys. 1959. 155(2): 206.

2. Lu C. Mass determination with piezoelectric quartz crystal resonators. J.Vac Sci Technol.

1975. 12(1): 578.

3. Kanazawa KK, Gordon II JG. The oscillation frequency of a quartz resonator in contact with

liquid. Analytica Chimica Acta. 1985. 175: 99-105.

4. Tamiya E, Karube I, Hattori S, Suzuki M, Yokoyama K. Micro glucose using electron

mediators immobilized on a polypyrrole-modified electrode. Sensors and Actuators. 2002. 18(3-

4): 297-307.

∆f (Hz)

for

Control

∆f (Hz)

for Ppy

30 nm

∆f (Hz)

for Ppy

600 nm

∆f (Hz)

for PEG

5000

∆f (Hz)

for PU

∆f (Hz)

for PVB

BAS

Concentration

mg/ml

0.3 8.8 0.9 1.2 1.5 2.2 6.6

0.3 9.0 1.0 0.1 1.1 3.5 0.8

0.3 8.0 0.8 1.1 1.0 1.7 0.9

0.3 0.4 1.5

0.3 2.4

3.0 15.0 6.0 2.7 3.5 6.6 2

3.0 12.8 5.5 2.0 3.1 6.9 2.5

3.0 10.2 5.3 2.6 2.7 6.9 3.1

30 20.0 17.0 11.5 16.8 10.7 12.7

30 23.0 20.0 11.3 11.5 11.1 10.0

30 22.4 16.0 12.2 11.6 11.0 9.5

26

5. Feldman K, Hahner G, Spencer ND, Harder P Grunze M. Probing resistance to protein

resistance of oligo (ethylene glycol)-terminated self assembled monolayers by scanning force

microscopy. J. Am. Chem. Soc. 1999. 121(43): 10134-10141.

6. Ishihara K, Shibata N, Tanaka S, Iwasaki Y, Kurosaki T, Nakabayashi N. Improved blood

compatibility of segmented polyurethane by polymeric additives having phospholipids polar

group. II. Dispersion state of the polymeric additive and protein adsorption on the surface.

Journal of Biomedical Materials Research. 1996. 32(3): 401-408.

7. Ishihara K, Tanaka S, Furukawa N, Nakabayashi N, Kurita K. Improved blood compatibility

of segmented polyurethanes by polymeric additives having phospholipid polar groups. I.

Molecular design of polymeric additives and their functions. Journal of Biomedical Materials

Research. 1996. 32(3): 391–399.

27

CHAPTER 4: Potential for enhancing Biocompatibility through microstructuring and anti-

coagulating biomolecular coating

1 Tina Talwar, William Kisaalita, Yan Geng and Yiping Zhao. Submitted as a chapter in

“Biosensors for Health, Environment and Biosecurity”, ISBN: 978-953-307-155-8, Austria,

Vienna: In Tech-Open Access Publisher.

28

Potential for enhancing the biocompatibility through microstructuring and anti-

coagulating biomolecular coating

Tina Talwar1,4

, William Kisaalita1,

*, Yan Geng2 and Yiping Zhao

3

1Cellular Bioengineering Laboratory, Faculty of Engineering, University of Georgia, Athens, GA

30602, USA

2Department of Chemistry, University of Georgia, Athens, GA 30602, USA

3Department of Physics and Astronomy, University of Georgia, Athens, GA 30602, USA

4Present Address: Applied Medical, 22872 Avenida Empresa, Rancho Santa Margarita, CA,

92688

*To whom correspondence should be addressed. Tel.: +1 706 542 0835; Fax: +1 706 542 8806.

E-mail: ttalwar@engr.uga.edu, williamk@engr.uga.edu, ygeng@chem.uga.edu ,

zhaoy@physast.uga.edu

4.1 Abstract

Implantable glucose biosensors have short life spans due to poor biocompatibility, resulting

from protein and cell attachment. To demonstrate the potential for increasing implanted sensor

life spans, we coated microstructured (microwells) SU-8 (epoxy-based negative photoresist

material) with heparinized medical grade polyurethane (Tecothane). The "sensor" surface

performance was evaluated by porcine platelet adhesion in vitro with a commercially available

implantable glucose sensor (MINIMED) surface as the control. The microstructured (5 - 7.5 µm

wells)/ heparinized SU-8 surfaces exhibited a lower cell attachment in comparison to the

MINIMED surface (P = 0.05). Results from separate protein adsorption experiments revealed

that the above said microstructured/ heparinized SU-8 surfaces offer maximum resistance to

29

biofouling and show promise to increase lifespan of a biosensor. This short communication

provides preliminary evidence in support of further exploration of the microstructured/

heparinized surfaces for increasing implantable sensor life spans.

Keywords: Polyurethane, SU-8 microstructure, heparin, implantable glucose biosensors

lifespan.

4.2 Introduction

Implantable glucose biosensors are fouled by protein adsorption followed by platelet adhesion

that activates the cells leading to formation of thrombus [1]. The sensor fouling negatively

impacts the long-term utility of the sensor by reducing glucose diffusion to the sensor resulting

in decreased current output, leading to either partial or complete malfunction of the sensor. These

circumstances call for frequent sensor replacement, about every 2-7 days [2, 3].

To overcome biofouling, numerous steps have been taken to improve biocompatibility. One

of these steps has included polymers coatings such as polyurethane (PU) and anti coagulants

such as heparin. Examples include using a charged PU [4, 5], blending and copolymerizing

segmented PU with 2-methacryloyloxyethyl phosphorylcholine and with cyclohexyl

methacryalate [6, 7], using hydrophilic PU with polylysine [8], and PU containing gold [9] or

silver nanoparticles [10]. Furthermore, anti coagulant/anti thrombin heparin has been

incorporated on sensors surfaces. For example, Moon et al. [11] dispersed heparin-DOCA

(deoxycholic acid) in PU, Michanetzis et al. [12] modified surface by ionic and covalent

heparinization, and Aldenhoff et al. [13] used heparinized coiled tubular polymeric structures.

Sung et al. [1] reported a life span of 30 days using a PA/ PEG/ heparin membrane. Lin et al.

30

[14] used Chitosan/ heparin polyelectrolyte complex onto polyacrylonitrile (PAN) membrane. In

other studies, collagen [15] and hydrogel [16] coatings have been attempted.

PU and heparin have also been combined in several studies. For example, Wan et al. [17]

combined heparin and PU by immobilizing heparin in PU with ester groups. Zhou and Meyerhof

[18], developed biomimetic trilayer polymeric (PVC or PU) coatings with heparin and a

controlled nitric oxide release.

Beyond biocoatings, tissue engineering studies have demonstrated that micro-pore shape,

size, and interconnectivity are important factors that influence cell ingrowth and proliferation

[19, 20]. An interesting observation in our laboratory is that for fibroblasts and nerve cells, there

is a minimum pore size (usually larger than the cell diameter) below which the cells avoid

growing and proliferating in the pore [21]. Based on this result, we reasoned that

microstructuring a sensor surface might have potential to enhance the sensor in vivo life-span

especially if combined with heparinized PU coating. To establish proof-of-concept, we

microstructured SU-8 photoresist material by introducing microwells and coated the structure

with heparinized PU, and compared platelet adhesion on the structure with adhesion on a

commercial implantable glucose sensor surface as a control. In this short communication, we

report preliminary evidence in support of the potential to enhance sensor in vivo life span

through microstructuring.

4.3 Materials and Methods

SU-8 microstructuring

31

Procedures previously described by Wang et al. [22] were followed. Briefly, microwells were

fabricated on 25-mm cover slips (Fisher Scientific, Pittsburgh, PA, USA). Before fabrication, the

cover slips were cleaned with 20% sulfuric acid and then baked at 110°C for at least 3 h. SU-8

(2025, MicroChem, Newton, MA, USA) was first spun onto the glass substrate at a speed of

2000 rpm for 30 s to obtain flat surfaces. 69% (w/v) SU-8 was spin coated on cover slips

resulting in coating thickness of about 40 µm. The SU-8 coated cover slips were soft baked, first

at 65°C for 3 min and then at 95 °C for 9 min. After baking, SU-8 was exposed in soft contact

mode with a Karl Suss MJB 3 HP Mask Aligner using 365 nm UV at 10 mW for four 15 s

intervals, interrupted for at least 20 s in between, which corresponds to a total UV exposure

intensity (400 mJ/cm2

). The flat SU-8 coated substrates were subjected to chromium mask on

photolithography to build micro wells like patterns. Microwell patterns and their nominal

structure dimensions used in this study included: 5, 7.5, 15, 30, and 50 µm diameters.

Heparinized polyurethane (PU) coating

Tecothane was first derivatized to obtain pendant 7-carboxy-5-thiaheptyl groups, following

procedure described by Alferiev and Fishbein [23]. Tecothane (15.5 g) was kept in toluene (150

mL, 60 hrs) and then dissolved in 300 mL of N, N- dimetheylacetamide (DMAc). Then, a 15

mL, 126 mmol of 1,4-dibromobutane was added. Further, 1.0 M of lithium tert-butoxide in

hexane (7.6 mL, 7.6 mmol) diluted with dry DMAc (20mL) was added at -5°C to -6°C. The

mixture was stirred at - 1°C under Ar for 1 hour. Then, 6.5 mL of acetic acid was added and the

solution was poured in cold (1200 ml, -55°C) methanol. Coagulated polymer was separated and

washed in methanol and iso-propanol and dried in vacuum at room temperature. This modified

polymer with pendant bromobutyl groups was dissolved in DMAc (220 mL) under argon and

cooled to -8°C. Thiolacetic acid (5072 mL, 80 mmol) and 0.25 M DMAc solution

32

oftetrabutylammonium tetraborate (80 mL, 20 mmol) were further added at 0°C. This modified

PU was again washed as described above and completely dissolved in DMF (5% wt/vol). The

filtered PU was then spin coated on cover slips.

A solution of polyallyamine hydrochloride (PAA.HCL Mw 15kDa, 3.9 mM),

hexadecylpyridium chloride (1.5 mM), N-hydroxysulfosuccinimide sodium salt (sulfo-NHS, 14

mM) and 1-ethyl-3-(3'-dimethylaminopropyl) carbodiimide hydrochloride (EDC, 0.11 M) was

first prepared in water and the solution pH was adjusted to 5.5 using KHCO3. Then the cover

slips were kept immersed in the solution for a day, and then rinsed thoroughly with water. The

solution was freshly prepared and the cover slips immersed the following day. After a 2-day

immersion, the cover slips were washed with 0.1 M of HCI and water. Then the cover slips were

put in 0.3% aq. K2CO3, rinsed and air dried. Another solution (pH 7.0) of sulfo-NHS (46 mM),

EDC (26 mM) and unfractionated heparin (0.5 mM) was prepared and put over the cover slips

under mild shaking at room temperature for 18-20 hrs and finally rinsed with water before use.

All chemicals were purchased from Sigma Aldrich.

Protein and cell adhesion

Porcine blood was collected in sterile BD EDTA vacutainer tubes (7 mL k3EDTA, Atlantic

Medical Supply, NY) and centrifuged at 836xg for 15 minutes at 4°C to obtain platelet rich

plasma (PRP). The PRP was separated and centrifuged again at 1000xg for 15 minutes at 4°C to

obtain cell pellet. The resulting supernatant was removed without disturbing the platelet pellet.

The cells obtained from the platelet pellet were resuspended in MEM (minimum essential

medium, Sigma Aldrich) and used for cell study, whereas PPP (no cells) was used for the protein

study.

33

Three types of non patterned surfaces were investigated for the protein study: 1) untreated

glass cover slip, 2) SU-8 coated cover slip, and 3) SU-8 coated cover slip treated with PU. These

surfaces were exposed to PPP for 6 hrs. The Bicinchoninic Acid Kit (BCA kit, Sigma Aldrich)

was used for desorbing and quantization of protein following procedures described in the kit and

those published by Allen et al. [24]. Briefly, the superfluous protein was removed and the

surfaces rinsed with warm (37°C) DI water. Adsorbed protein was desorbed by a protein

solubilizing solution containing sodium dodecyl sulphate (SDS, 3% (wt/vol)), 8 M urea, DL-

Dithiothreitol (1 mg/ mL) in PBS (phosphate buffered saline). The desorbed protein containing

solution was centrifuged (50 min at 5000 x g) using Centricon filters (YM-30, Millipore) and

resuspended in 500 III of PBS (pH = 7.0). Protein concentration was assayed in 96 well plates at

595 nm with an absorption plate reader (DYNEX MRX, Dynex tech., Chantilly, VA). The

amount of protein adsorbed on the four different surfaces was determined by comparing the

absorbance values against a calibration curve obtained with 0 to 100 µm/mL protein standards.

Cell attachment study was conducted with patterned structures: 1) SU-8 coated cover slips

and 2) and microstructured SU-8 coated cover slips. The human or porcine platelet size is

estimated to be between 2 to 4 µm [25]. Structures with 5, 7.5, 15, 30, 50 µm microwell sizes

were investigated. Experiments were replicated at least three times. Cells were stained with

Calcein (2.0 µM, Sigma Aldrich) prior to the microstructure exposure and then incubated at 37

°C in dark for 10 minutes. The cells "fallen" into the microwells were observed at 30 min, 1, 5,

and 19 h intervals.

Assessing lifespan enhancing potential

Cell attachment was compared on five surfaces: 1) untreated cover slip, 2) a 5 µm (microwell

diameter) patterned SU-8 substrate, 3) a 5µm SU-8 pattern with PU coating, 4) 5 µm SU- 8

34

pattern with PU + Heparin coating, and 5) MINIMED (Medtronic) glucose sensor surface. In

practice, MINMED is replaced every 2 days. Although MINIMED is a subcutaneous glucose

measurement it‟s interaction with platelets was considered a valid reference for our studies. The

MINIMED sensor surface was cut and used in the experiments. Cell pellet in MEM was used

and cell attachment was observed at 30 min, 1 h, 5 h and 19 h intervals. The experiments were

replicated at least 3 times.

4.4 SEM and X-ray analysis

SEM imaging was used to assess the quality of coating and to determine the dimensional

details of the microwells. Structures were sputter coated with gold for 60 s to achieve an Au

coating thickness of about 15.3 nm. SEM and X-ray images were captured with scanning

electron microscope (Zeiss 1450EP variable pressure SEM Carl Zeiss MicroImaging, Inc. One

Zeiss Drive, Thornwood, NY 10594, and Oxford Instruments X-Ray Technology, Inc., 275

Technology Circle, Scotts Valley, CA 95066). The X-ray peaks indicated the calcium content

present in Calcein, a dye used to stain cells, thus confirming cell attachment on PU and untreated

cover slips (non heparinized surfaces).

4.5 Results and Discussion

Protein adsorption

A nonthrombogenic, haemocompatible surface is mimicked by PU and Heparin combination

[18, 26]. Protein adsorption study was conducted for untreated cover slip (control), untreated SU-

8 and polyurethane-coated SU-8 (SU-8 + PU) in triplicate. Protein was measured as follows:

35

first, a baseline was first established using standard protein concentrations in the 0 µg/ ml -100

µg/ ml range using BCA and spectrophotometer. The absorbance values obtained for the

unknown protein concentrations for all the samples were compared against the baseline and the

corresponding protein concentrations were determined. Table 1 summarizes the effect of surface

coatings on protein adsorption. As expected, the untreated surfaces yielded the highest

absorbance values, indicating the highest protein adsorption. Protein concentration estimates

were: cover slip (1016 µg/ ml)>> SU-8 (113 µg/ ml) >> SU-8 + PU (10.9 µg/ ml). The total

protein average from the porcine protein sample (PPP) was 2.9 gm/dl (Roche 4/BMC, Hitachi

912 Analyzer).

The F-statistic applied to the log-transformed net absorbance (NA) values was 128.54, (p <

.0001), indicating that the null hypothesis that the three treatments were equivalently affected

was utterly rejected. Separate pair wise t-tests of the differences confirmed that the three groups

were significantly different with a P values < .0001. In particular, cover slip's mean NA value

was 9.09 times larger than SU-8's mean NA value, and SU-8's mean NA value, in turn, was

10.38 times larger than that of SU-8 + PU.

Cell Adhesion

Surface coating uniformity is an important aspect in experimentation. For coating quality

sake, a comparison of surface PU coating uniformity is shown in figures 1a and 1b. The cross

sectional view (SEM) in fig. 1c and top view in fig. 1d, confirm the microwell size of 5µm.

Fig.1e shows the micro well pattern and SU-8 coating depth. Different sized micro wells ranging

from 5µm to 50 µm were tested against cell attachment at set intervals (Table 2). Attachment of

cells inside microwells was time and well-size dependent, consistent with prior work [27]. With

36

microwell diameters less than 1.5 times (5 µm) the size of the platelets (2 - 4 µm), either cells

could not survive or failed to "fall" into the wells as reflected by the absence of cells in Figures

1f (5 µm), 1g (5 µm), and l h (15 µm); whereas in Figure 1i (50 µm), cells are clearly visible. In

general, independent of well size, cells tended to move away from well during the first hour of

incubation (also reported by Li. et al. [28]) on the different sized microwells. This cell behavior

is depicted in Table 2. With time, cells were observed to have "fallen" into larger size microwells

(30 and 50 µms) but not at all in 5 µm wells. Four samples including: SU-8 coated cover slip

(control), SU-8 + PU, heparinized SU-8 + PU and MINIMED were compared for cell attachment

at 0.5, 2, 4 hrs intervals. Each structure, with the exception of MINIMED was microstructured

with 5 µm wells. Figures 2a, 2b and 2c show cell attachment on non-heparinized SU-8 and PU

surfaces. Cell attachment 5 hrs post incubation for the microstructured and MINIMED samples

are shown in figures 2d and 2g respectively. A Fisher's Exact 1-sided test was conducted for the

5 µm SU-8 and MINIMED surfaces. Of the 6 trials, 3 resulted in cell attachment on MINIMED

and for the other 3 trials no cells attached on the microstructure substrate. According to Fisher‟s

Exact 1-sided test, there can be 4 possible groups (A to D, see Table 3) out of the 20 different

permutations (1, 9, 9, 1). These groups are arranged by how many successes (attachments) are in

the two groups.

The pattern observed was “D”, which is the most extreme in favor of the alternative

hypothesis that attachment is more common on MINIMED than on SU-8 at a P value of 1/ 20 or

0.05. MINIMED surfaces failed to resist cell attachment within the first hour (Fig. 2f) as opposed

to 5µm heparinized SU-8 + PU (Fig. 2d). To confirm that the fluorescent spots in the figures

presented were indeed cells, we conducted X-ray analysis of non heparinized PU coated surfaces

under the influence of calcein stained cells. Calcein is a complexometric indicator for titration of

37

calcium ions with EDTA, and enables flurometric determination of calcium. The acetomethoxy

group in calcein obscures the part of the molecule that chelates calcium. However, once inside

the cell, enzymes cleave off the group and the molecule binds to calcium within cell resulting in

green fluorescence. Thus the X-ray peaks in fig. (2h) are showing calcium content confirming

that the spots on the structure were cells.

4.6 Concluding remarks

This study reinforced the idea that mammalian cells respond differently when exposed to

microstructured substrates. Depending upon the pore size (5-15µm), the cells move away from

the microwell zone for at least the first hour of incubation, whereas for larger pores (30 µm and

50 µm) the cells immediately fall into the wells. The unique SU-8 + PU + heparin coating was

found to resist protein with heparin working as anti thrombin agent. Here, SU-8 was used for as a

coating and is found to be protein resistant, further enhancing the protein resistance feature of the

entire coating. The 5µm structure eliminated cell attachment. Thus the two major reasons (cell

and protein attachment) for sensor biofouling were tackled with this two pronged approach. Our

study shows promise for improved biocompatibility and offers support for more detailed studies

to confirm the potential for implantable glucose sensor in vivo life enhancement through sensor

surface microstructuring.

38

References

1. Sung WJ, Na K, Bae YH. Biocompatibility and interference eliminating property of pollulan

acetate/PEG/heparin membrane for the outer layer of an amperometric glucose sensor. Sensors

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42

Acknowledgements

We extend our thanks to Dr. Robert Dove, Animal and Dairy Science and Ju Rong,

Department of Chemistry for their technical support. This work was supported by NSF

(0304340) and UGA Engineering Grants

43

Table 4.1 Effect of surface treatments on protein adsorption.

SU-8 (3 samples) SU-8 + PU (3 samples)

Untreated Cover slip

(5 samples)

Net

Absorbance

nm

Protein

Concentration

µg/ ml

Net

Absorbance

Nm

Protein

Concentration

µg/ ml

Net

Absorbance

nm

Protein

Concentration

µg/ ml

0.060 86.2 0.008 12.1 0.9005 1500.88

0.071 101.0 0.010 14.8 0.489 816.44

0.116 165.2 0.005 7.346 0.5158 859.77

0.5102 850.33

0.6311 1051.83

44

Table 4.2 Cell attachment inside microwells

Size of the

Microwell

(µm)

Incubation Time (hours) Total no. of wells

(5.25" X 8"

snapshot section)

5% of total wells

in 5.25" X 8"

snapshot section

0.5 2 4 19

5 - - - - 242 12

7.5 - - + + 165 8

15 - + + ++ 30 2

30 - + + ++ 18 1

50 + ++ ++ ++ 9 1

“-“ means no cells in the 5.25" X 8" microscopic image section; “+” means cells in

microwells (5% to 50% wells of the total number of wells in the 5.25" X 8"microscopic image

section); “++” means cells in microwells (greater than 50% wells of the total number of wells in

the 5.25" X 8" microscopic image section).

45

Table 4.3 Fisher‟s Exact 1-sided test for the 5 μm SU-8 structure and MINIMED surfaces.

Substrate A B C D

SU-8 3 2 1 0

MINIMED 0 1 2 3

Possible

Permutations

1 9 9 1

46

Figure 4.1 SEM microscopy images showing nonuniformly (a) and uniformly (b) PU-coated

cover slips; crossection (c), top view (d), and coating thickness – approximately 51.33 μm – (e)

of SU-8 structures. Fluorescence of calcein stained cells floating on top of 5 μm well structure

after 1 (f) and 4 (g) hours and inside a 15 µm (h) and 50 µm (i) microwells after 4 hours of

plating.

e f

g h i

a b c

d

47

Figure 4.2. (a) Fluorescence image showing Calcein-stained cells attached on control (non

heparinized SU-8); X ray images of cell attachment (highlighted in inset) on SU- 8 (b) and PU

(c); SEM microscopy images showing no cell attachment on 5 μm heparinized SU-8 + PU

structure, 5 hr after incubation (d), MINIMED without cell incubation (e), cell attachment on

MINIMED after 1 (f) and 5 hrs (g) hr of incubation; X ray spectrum for SU-8+PU structure (h)

confirming the observed fluorescent spots to be cells.

d e f

g h

a b c

48

CHAPTER 5: Conclusion and Future Directions

5.1 Discussion of results

The specific aim of this study was to demonstrate the potential to enhance the life span of

implantable glucose sensors through sensor surface microstructuring. We hypothesized that a

biocompatible SU8 microstructure with microwells at size 5 µm (approximately 1.5 times the

platelet size) coated with anti protein PU (polyurethane) polymer, restricts the adherence of both

mammalian cells and protein onto the microstructured surface and prevents biofouling of the

sensor site of the glucose sensor. Our first step was to select the polymer material to use in the

study. Using QCM (Quartz Crystal Microbalance), we evaluated Ppy_30 (polypyrrole, 30 nm

thick), Ppy_600 (polypyrrole, 600 nm thick), PEG (polyetheylene glycol), PVB (poly vinyl

butaryl) and PU with respect to protein adsorption. As shown in Chapter 3, all polymers tested

were equally protein resistant and the ultimate choice of PU, for use in subsequent studies, was

based on ease of immobilizing heparin into the polymer for purposes of creating

antithrombogenic polymer. Heparin is a well known anti thrombic agent that prevents blood

coagulation. It‟s incorporation in a protein resistant polymer, PU and the final coating of PU +

Hep on Su 8 enhances the biocompatibility characteristic of the entire combination by helping

sensor surface stay free of blood cells and protein. However, heparin is mainly anti

thrombogenic, does not completely prevent cell attachment and is leachable. This limitation

directed to include a microstructured sensor site that prevents cell attachment completely.

49

To determine the best microwell size for microstructuring sensors surfaces, we used the photo

resist material SU-8 and photolithography techniques described in Chapter 4 to microstructure

with 5, 7.5, 15, 30 and 50 µm-diamter microwells. All microwells were 1.84 µm deep. By

exposing the microstructured SU-8 materials to porcine platelets for a range of 30 min to 4 hr,

we found that platelets refrained from entering the 5 µm microwells, but not in larger

microwells. The number of cells observed into the microwells increased with increasing

microwell diameter.

To demonstrate potential for sensor lifespan enhancement, we combined surface

modification and microstructuring by coating the 5 µm microwell structured SU-8 material with

antithrombin heparin immobilized PU to prevent both protein and cell attachment. With

MINIMED surface as control, this preparation was tested with respect to porcine platelet

attachment. As reported in Chapter 4, cells were observed to immediately populate MINIMED

surface while the 5 µm Su8-PU-Hep surface was able to resist cell attachment. The difference

between the rate of cell attachment suggested the combination to have potential for slowing

sensor biofouling and thus enhance the in vivo sensor life span.

Our study focused on decreasing the frequency of replacement (current continuous glucose

monitoring sensors are FDA approved for 3-7 days) of continuous glucose monitoring sensors by

subcutaneous implantation of glucose sensor and was compared to the only commercially

available implantable glucose sensor surface (MINIMED). Since cell attachment (irrespective of

the type) thrives under protein environment and plays a vital role in biofouling for any

implantable device, we dedicated a part of our research toward anticell and anti protein solutions.

49

With the available resources, platelets were picked for mammalian interstitial fluid cells, mainly

for their comparable size and the tendency to adhere. The idea was to design a better tolerated

biocompatible sensor surface that would mitigate the local immunogenic and thrombotic

response. Further research can address site specific tissue reaction including cell and protein

adherence at the site of implantation. Another limitation of our study was the in vitro

environment, where the pH of the plasma protein and high degree of protonation, could have

played a small role in preventing the protein attachment onto PU heparin surface. Further

research is needed to study protein absorption under true physiological conditions or in vivo.

5.2 Future Directions

Our results provide a compelling need to further investigate the combination of surface

modification and microstructuring in extending the life span of implantable sensors. Since the

proposed sensors are meant for subcutaneous implantation, wherein the sensor never comes in

contact with blood and platelets, it would be wise to address local tropism and cell and protein

adherence at the site of implantation by using laboratory cultured mammalian cells. In such cell

culturing methods, the media is changed about every 3 days and cells are passaged using trypsin-

EDTA solution as standard. It would thus also be worthwhile to study how EDTA interferes with

cell and protein inhibition on the sensor surface. In agreement to previous work, our study shows

protein resistant on PU + Hep surface with heparin as the anticoagulant, over a period of time the

entire surface can render as highly thrombogenic mainly due to heparin leaching effect. Thus, a

study that focuses on leaching prevention of additives either by plasma surface etching or some

other surface modification method can benefit or improve the performance of such implantable

sensors.

50

Indwelling medical devices get attacked by microorganisms and form microbial biofilms.

These bacteria are resistant to antimicrobial treatments and tenaciously adhere to the surface,

further forming a structural matrix and leading to increased infections at the implanted site.

Future research should also focus on incorporating antimicrobial agents that will reduce infection

risk and an early removal of the device due to infection.

51

APPENDIX 1: PU synthesis and modification with pendant acetylated thiol groups

Material(s): Techothane TT1074A, 1, 6 dibromohexane, polycaprolactone (PCL), Thiolacetic

acid, 4, 4`- diphenylmethane diisocyanante (MDI), 1, 3-diaminopropane (1, 3-DAP), lithium tert-

butoxide, tetrabutylammonium tetraborate, dimethylacetamide, methanol, 2- propanol,

dimethylformamide.

All the chemicals were obtained from Sigma Aldrich.

A „two step polymerization‟ method was used to synthesize polycaprolactone (PCL) as soft

segment and 4, 4`- diphenylmethane diisocyanante (MDI), 1, 3-diaminopropane (1, 3-DAP) as

the hard segment. Basically the synthesis was carried out in three major steps: bromoalkylation

to first achieve bulk carboxylation of PU, surface amination of PU and lastly heparinization.N-

bromobutylated polyether urethane was modified with pendant acetylated thiol groups, fig. (A-

1). This particular method resists the oxidative cross linking and is known to have high tensile

and mechanical strength which again contribute to the sensor‟s lifespan.

Procedure(s):

1. 15.5 g of Tecothane TT1074A was reacted with 7.1 mmol of 1, 6- dibromohexane (or 1, 4-

dibromobutane) in copious amounts of thiolacetic acid in the presence of thiacetate ion. This

provided PU with pendant bromoalkyl groups suitable for further derivatizations such as

covalently immobilizing with heparin.

52

1, 4-dibromobutane and lithium tert-butoxide as bases reduce the exposure of PU to strong basic

conditions thus minimizing PU degradation. These bromoalkyl side chains are then able to react

with thiol groups under mild conditions.

2. The modified PU was then dissolved in 220 ml of dry DMAc under argon flow.

3. The solution was cooled at -80

C.

4. Freshly vacuum distilled (at 115 mm Hg) thiolacetic acid (5.72 ml, 80 mmol) was added to

this solution.

5. A mix of 0.25 M DMAc solution and tetrabutylammonium tetraborate (Bu4N)2B4O7 (80 ml,

20 mmol) was prepared and kept at 00

C.

6. The mixture was stirred at low temperature for an hour under argon flow and poured into 1400

ml of cold (-600

C) methanol.

7. The polymer coagulate was separated, washed with methanol and 2-propanol and dried in

vacuum (0.5 mm Hg) at room temperature.

8. The PU obtained was kept at 40

C avoiding any light.

9. The resulting PU polymer (4% wt/vol) was then dissolved in DMF (dimethylformamide), and

stirred continuously for 4 hours for complete dissolution.

Fig A1.1 N-bromobutylated Polyether-Urethane

53

APPENDIX 2: PU and heparin binding

Material(s): PAA-HCl, hexadecylpyridium chloride, N- hydroxysulfosuccinimide sodium salt,

potassium bicarbonate. All these chemicals were purchased from Sigma Aldrich.

25mm cover slips (Fisher Scientific, Pittsburg, PA)

1. PU coated cover slips were prepared as described in Appendix 6.

2. The heparin and PU immobilization is shown in fig. (A-2). PU surfaces were immersed for 24

hours in an aqueous solution of PAA-HCl (58 mg/ml), hexadecylpyridium chloride (0.5 mg/ml)

and N- hydroxysulfosuccinimide sodium salt (sulfo-NHS -3mg/ml) EDC (22 mg/ml).

3. The pH was adjusted to 5.5 using KHCO3 (potassium bicarbonate).

4. The cover slips were then taken out and rinsed thoroughly with DI water.

5. Washed cover slips were again immersed in the same solution of PAA-HCl (58 mg/ml),

hexadecylpyridium chloride (0.5 mg/ml) and N- hydroxysulfosuccinimide sodium salt (sulfo-

NHS -3mg/ml) EDC (22 mg/ml) solution for 2 days.

6. The cover slips were washed with 0.1 M HCl and water.

7. The cleaned cover slips were then put in 0.3 % aq. K2CO3 for 4 hours.

8. The cover slips were taken out, rinsed and air dried.

54

9. The cover slips were then put in a solution of sulfo NHS (10mg/ml), EDC (5 mg/ml) and

heparin (10 mg/ml) at pH =7.0 with mild shaking at room temperature for 20 hrs.

10. Finally the cover slips were taken out and washed with lots of DI water. This scheme of

using amine group ensures that heparin is chemically bound onto PU and not just adsorbed

superficially.

Fig A2.1 Immobilization of Heparin onto the surface- aminated PU films

55

APPENDIX 3: SU-8 Coating and Micro patterning

Material(s): Photo resist SU-8 2025, SU-8 developer, MCC primer 80/20 (80% PM Acetate, 20%

HMDS) (Microchem, Newton, MA), 25 mm cover slips (Fisher scientific, Pittsburg, PA)

Equipment: UV Mask Aligner (Karl Suss MJB 3 HP), Masks (Advance Reproductions Corp.),

Vacuum Pump (Gardner Denver Thomas Inc., Welch Vacuum Technology, Niles, IL), Spin

Coater (Speciality Coating System Inc., Indianapolis, IN). This procedure was performed in a

clean room.

The type of SU-8 used in our sensor construction is SU-8 2025, ideally suited for permanent

applications. A single coat can give a thickness in the range 0.5 – 200 microns. The exposed and

subsequently thermally cross-linked film is insoluble in SU-8 developers. SU-8 2025 has

excellent imaging characteristics and is capable of producing high aspect ratios and is a cost

effective solution to produce fine patterns. SU-8 can be processed with standard lithography

techniques, UV mask aligner and coated with either spin, spray or screen processes. SU-8 2025

(purchased from Microchem, Newton, MA) has 69 % solids, with a density of 1.219 gm/ml.

There is flexibility in the soft/ hard baking temperatures and spin coating parameters that

depend on type of substrate needed for an application. SU-8 microstructures of the sizes 5µm,

7.5µm, 15µm, 30µm and 50µm (no channels) were first prepared following the recipe.

Procedures:

1. Cover slips were immersed in 20% sulfuric acid (H2SO4) at room temperature for at least a

week prior to SU-8 patterning.

2. The cover slips were then thoroughly washed with DI water and baked at 110 0C overnight.

57

3. Vacuum pump was run at (30-40 in Hg) and nitrogen was turned on at 70 psig.

4. The photolithography machine was fired till it reached a temperature of 195 0C.

5. The cover slips were baked at 65 0C for 30 minutes.

6. The spin coater stage was covered with aluminum foil and the vacuum nozzle, chuck were

cleaned with acetone.

7. The cover slips were coated with MCC primer 80/20 (20 % HMDS or hexadimethylene

silicate and 80% PM acetate) for efficient SU-8 coating.

8. The cover slips were then spin coated with SU-8 (about 1 ml) The following spin parameters

were applied for spin coating.

RPM1: 500

RAMP1: 0005

TIME1: 0001

RPM2: 2000

RAMP2: 0005

TIME2: 0015

RPM3: 2000

RAMP3: 0005

TIME3: 0015

RAMP4: 0015

58

S: 0000

N: 0000

COATING: 0000

9. After coating with SU-8, the cover slips were baked at 65 0C for 3 minutes and then at 95

0C

for 30 minutes.

10. The UV lamp was turned on by pressing the „Start‟ button. This step can take 3-4 attempts

until it reads “rdy” for “ready”.

11. The mask aligner and microscope manipulator were turned on.

12. The mask is loaded and “Vacuum Mask” was pressed to „fix‟ the mask for exposure.

13. The cover slips were then exposed for 15 sec for 4 times, with 20 sec interval with differently

sized masks by pressing the „start‟ button.

14. Post exposure baking was done at 65 0C for 3 minutes and at 95

0C for 9 minutes.

15. The cover slips were then developed in SU-8 developer for 14 minutes with mild shaking and

were rinsed with IPA (isopropyl alcohol) and dried under N2 (g).

16. The cover slips were again baked at 65 0C for 3 minutes and at 95

0C for 9 minutes.

17. A final exposure was done without mask, 15 sec for 4 times, with 20 sec interval.

59

18. They were then developed again for 14 minutes in SU-8 developer with mild shaking, and

finally rinsed with IPA and dried under N2 (g).

19. At the end, the cover slips were soft baked at 65 0C for 3 minutes and at 120

0C for 30

minutes. All through the baking, care was taken that pattern side was kept up. The masks were

cleaned with acetone and dried with Nitrogen.

Thus microstructures of the sizes 5, 7.5, 15, 30 and 50 µm were prepared and analyzed for the

cell adhesion study.

60

APPENDIX 4: PPP, PRP extraction and Cell Counting

Material(s): Porcine blood collected by sinus orbital method, Swine facility, UGA. 4 ml BD

Vacutainer EDTA tubes, Minimum essential Medium (MEM) by Sigma Aldrich

Equipment: Centrifuger, Hemocytometer

1. Porcine blood was considered for all the biocompatibility tests, as pig is the closest model to

humans. Porcine whole blood (5ml, PIC crossbreed) was collected in EDTA (anticoagulant)

tubes

2. The tubes filled with fresh blood were centrifuged at 1840 rpm for 15 minutes to obtain PRP

(platelet rich plasma).

3. The PRP was separated and centrifuged again at 2200 rpm for 15 minutes to obtain PPP

(platelet poor plasma).

4. The platelet pellet remains at the bottom of the tube. The supernatant was carefully removed

without disturbing the pellet.

5. The pellet was then re-suspended completely in MEM media (MEM–nutrient mix F12 (1:1,

Gibco) with 2.5 mM -glutamine, 15 mM HEPES, 5.2 g/l glucose, 50 μg/ml gentamycin and

15% calf serum; Gibco), making sure that there were no cell clumps or agglomerates left.

6. 10 µl of this suspension was loaded on the two counting chambers on a hemocytometer

(Neubauer), (Figures A-3 and A-4).

60

7. The sample injection areas get filled with the sample by capillary action. The cells were

allowed to settle for at least 10 minutes under hydration (covered with a moist tissue stuck on the

inner side of a petri dish), before observing under a 40X magnification to determine the cell

density.

Fig. A4.1 C-CHIP (DHC-NO1) hemocytometer

Fig. A4.2 Counting grid

8. The counting grid consists of 9 large squares, each measuring 1 x 1mm, giving a total of

counting area of 3 x 3mm, with a depth of 0.1mm. For our experiments the cells were counted in

the 5 large squares (excluding the four corner squares and including the middle square) in both

the chambers and the cell count per µl was calculated.

61

Since the pig platelets varied in the wide range of 400 – 900 x 103

/ µl, the manual cell counts

were cross verified on coulter automatic cell counting machine.

62

APPENDIX 5: Cell Adhesion on SU-8 micro wells

Material(s): Calcein, MEM (Sigma Aldrich)

Equipment: Phase Contrast Microscope, SLR camera (Nikon)

Calcein (2.5 µl/ml) was added to the pellet and MEM solution, and incubated for 30 minutes at

37 0C in dark. 1.5 ml of this solution was put on SU-8 micro patterns (5 µm, 7.5 µm, 15 µm, 30

µm and 50 µm). These were then observed and photographed, under phase contrast microscope

(40 X and 60 X magnifications) to see the cell adhesion and the no. of cells entering in the micro

wells at 0 hr, 0.5 hrs, 2 hrs, 4 hrs and 19 hrs.

63

APPENDIX 6: PU spin coating parameters

Material(s): Non heparinized modified Techothane (Appendix 1), dimethyl fluoride (DMF),

BCA kit (Bicinchoninic Acid), PBS, sodium dodecyl sulphate-SDS, DL-dithiothreitol-DTT,

urea, All these chemicals were purchased from Sigma Aldrich.

Equipment: Spin Coater (Speciality Coating System Inc., Indianapolis, IN), Incubator,

Spectrophotometer (DYNEX MRX), Centricon YM-30 filters (Fisher scientific).

Procedure:

1. Clean the chuck, stage of the spin coater with acetone and cover the stage with aluminum foil.

2.The dissolved PU in DMF (50 -60 µl) was spin coated on cover slips pre treated with SU-8

coated (as described in Appendix 3) with the following settings:

RPM1: 500

RAMP1:0005

TIME1: 0001

RPM2: 2000

RAMP2: 0005

TIME2: 0060

64

RPM3: 5000

RAMP3: 0005

TIME3: 0120

RAMP4: 0015

S: 0000

N: 0000

COATING: 0000+

Protein Adsorption study

1. Porcine whole blood was centrifuged as described in the Appendix 4.

2. The supernatant protein serum PPP-solution of interest for protein adsorption study was

carefully separated.

3. Here, SU-8 coated, PU (polyurethane) coated on SU-8 and plain cover slips were exposed to

PBS (Phosphate buffer saline to maintain the pH) supplemented PPP (platelet poor plasma) for 6

hours at 370

C.

65

4. After 6 hours the protein solution was removed and the surfaces were gently rinsed with pre

warmed distilled water.

5. Protein solubilising solution (sodium dodecyl sulphate-SDS 3 % wt/vol in DL-dithiothreitol-

DTT 1mg/ml in 8M urea) was put on the cover slips to recover all the adsorbed protein and all

the cover slips were incubated overnight at 37 0C.

6. The desorbed proteins were then added in the sample reservoir of the Centricon ultra filters,

without touching the filter membrane.

7. The proteins were then concentrated using Centricon YM-30 filters (Fisher scientific) by

centrifuging (using fixed angle rotor) at 5000 g (4780rpm) for 45 minutes. The retentate in the

sample reservoir is the concentrated protein which was re-suspended in 100 µl of 1.0 M Tris

buffer at pH 7.4 (Allen et al., 2006).

8. Protein concentration was determined using BCA (Bicinchoninic Acid Kit and DYNEX

MRX). This method uses a reactive solution of bicinchoninic acid (BCA) and CuSO4 of green

coloration. Cu++

ions are reduced by proteins of the cell suspension in Cu+ ions, which form a

complex with BCA. The crimson coloration of this complex is directly proportional to the

protein concentration and the absorbance read in a spectrophotometer at 570 nm using an

ultraviolet spectrometer.

9. Protein concentrations of 100µg/µl, 80µg/µl, 60 µg/µl, 40 µg/µl, 20 µg/µl, 10 µg/µl, 1 µg/µl

and 0 µg/µl were prepared to obtain a standard curve on the absorbance vs. protein concentration

plot.

66

10. 25 µl of each of the known proteins and 25 µl of the unknown proteins recovered from the

cover slips were loaded in a 96 well plate.

11. The reagents A (Bicinchoninic Acid Solution) and B (4% (w/v) CuSO4• 5H2O Solution) were

mixed in the ratio 19 ml: 0.38 ml and 200 µl of this mixture was then added to the proteins in the

well plate.

12. Absorbance was noted on DYNEX MRX machine and the proteins were then determined

based on the principle that more the amount of protein in the solution, higher the absorbance.

Table A6.1: BSA preparation at different concentrations for polymer selection with QCM

For 50 ml of 30 mg/ml BSA 1.5 g BSA + 50 ml PBS

For 50 ml of 3 mg/ml BSA 5 ml of 30 mg/ ml BSA + 45ml PBS

For 50 ml of 0.3 mg/ml BSA 0.5 ml of 30 mg/ ml BSA + 49.5 ml PBS