potential for enhancing biocompatibility through
TRANSCRIPT
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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15
7. Sung WJ, Na K, Bae YH. Biocompatibility and interference eliminating property of pollulan
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microdevices. Applied Surface science. 2003. 206(1-4): 218-229.
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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
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biocompatibility and cell patterning selectivity. 2004. Biosensors and Bioelectronics. 20(9):
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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
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16
16. Gretzer C, Gisselfalt K, Liljensten E, Ryden L, Thomsen P. Adhesion, apoptosis and
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17
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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: [email protected], [email protected], [email protected] ,
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: [email protected], [email protected], [email protected] ,
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
and Actuators. 2004. 99(2-3): 393-398.
2. Jin W, Brennan JD. Properties and applications of proteins encapsulated within sol-gel derived
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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