Surface modification of vascular prosthesis and intracorneal lens polymers


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Surface modification of vascular prosthesis and intracorneal lens polymers
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xvii, 208 leaves : ill. ; 29 cm.
Widenhouse, Christopher William, 1967-
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Materials Science and Engineering thesis, Ph. D
Dissertations, Academic -- Materials Science and Engineering -- UF
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Thesis (Ph. D.)--University of Florida, 1996.
Includes bibliographical references (leaves 195-206).
Statement of Responsibility:
by Christopher William Widenhouse.
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University of Florida
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Copyright 1996


Christopher William Widenhouse

This Dissertation is dedicated to all the people who believed

I could do it more than I did myself; especially, Tammy,

whose love and support I could not do without;

and to the memory of Charles Bucaria.


I would like to express my deepest gratitude to my

advisor and doctoral committee chairman, Dr. Eugene P.

Goldberg, for his guidance, encouragement, and patience.

Sincere thanks are also extended for the advice and teaching

of the members of my supervisory committee: Dr. Christopher

D. Batich, Dr. Anthony B. Brennan, Dr. James S. Seeger, and

Dr. Richard Dickinson. I am also grateful to the research

group of Dr. James S. Seeger. The guidance and input of Dr.

Dinesh 0. Shah, an original committee member, is also


Special appreciation is also felt for the assistance and

encouragement from my colleagues during my graduate tenure.

These include Jesse Arnold, David Bennett, Dr. Sameer

Bhatia, Dr. Emmanuel Biagtan, Charles Bucaria, Scott Butler,

Don DePalma, Shannon Eggers, Kirk Foster, Penelope Kao, Dr.

Steven Kuo, Ingrid Leidermooy, Dr. Tung Liang Lin, Lili

Mateo, Dr. Khalid Mentak, Julie Miller, Tom Miller, Dr. Lynn

Peck, Mark Privett, Dr. Jeanne Quigg, Katsabu Rao, Dr.

William Toreki, John Wironen, Dr. Ali Yahiaoui, Stacey Zambo,

and Mike Zamora.

Special thanks are extended to James F. Kirk for his

wise and thoughtful input, Drew P. Amery for assistance in

learning the ropes of analytical techniques, Paul J. Martin

for his lively debates and invaluable microscopy assistance,

Dr. Anthony B. Brennan for always pointing out the not so

obvious and to his openness and fairness to all questions,

James S. Marotta for his enthusiasm and encouragement, and to

my wife, Tammy for her love, support, pessimistic optimism,

and encouragement.



ACKNOWLEDGMENTS ........................................... iv

LIST OF TABLES ............................................. ix

LIST OF FIGURES ............................................xi

ABSTRACT ................................................. xvi


1 INTRODUCTION ............................................. 1
1.1 Vascular Grafts ...................................1
1.2 Intracorneal Lenses ...............................8

2 BACKGROUND ...............................................14
2.1 Synthetic Vascular Grafts ........................14
2.1.1 Vessel Replacement Surgery and Vascular
2.1.2 Properties of the Natural Vessel..........18
2.1.3 Synthetic Vascular Graft Materials and
Properties ................................23
2.1.4 Advantages of PMMA, SEMA, and PDMS
Surfaces .................................31
2.1.5 Gamma Radiation Initiated
Polymerization ............................38
2.1.6 Polymer Solution Coatings and
Techniques ...............................45
2.2 Intracorneal Implants ............................48
2.2.1 Refractive Corneal Surgery and
Intracorneal Implants ....................48
2.2.2 Intracorneal Lens Materials and
Designs ..................................49
2.2.3 Surface Properties of Ocular
Biomaterial Implants ......................52
2.2.4 Surface Modification Techniques ...........53

3 MATERIALS AND METHODS .................................55
3.1 Materials .......................................55
3.1.1 Substrates................................55
3.1.2 Monomers and Reagents .....................57
3.2 Methods ..........................................57
3.2.1 Sample Preparation and Substrate

3.2.2 Surface Modification Methods .............60
3.2.3 Solution Dip Coating of PDMS onto
Dacron* ...................................63
3.2.4 Characterization ..........................65
3.2.5 Diffusion/Flow Cell Testing of ICLs ....... 73
3.2.6 Porosity and Coating Stability Testing
of Vascular Grafts ........................75
3.2.7 Ex Vivo and in Vivo Studies of Vascular

4 RESULTS AND DISCUSSION...................................80
4.1 Gamma Radiation Induced Polymerization of
Methyl Methacrylate on PET, PTFE, and PDMS ......80
4.1.1 Swelling of PET, PTFE, and PDMS in MMA
Solutions ................................81
4.1.2 Radiation Grafting of MMA on PET, PTFE,
and PDMS .................................95
4.2 Gamma Radiation Induced Polymerization of
Sulfoethyl Methacrylate on PET .................. 146
4.2.1 Swelling and Gravimetric Analysis of
PET Surface Modified with SEMA ...........146
4.2.2 XPS Analysis of Radiation Grafted SEMA
onto PET .................................147
4.3 Solution Dip Coating and Thermal Curing of
PDMS Coatings on Dacron* PET .................... 148
4.3.1 Conditions for Dip Coating Dacron with
PDMS ..................................... 148
4.3.2 Thermal Curing of PDMS Coatings on
Dacron ..................................149
4.3.3 Gamma Polymerization of MAOP-t-PDMS..... 150
4.3.4 Analysis of PDMS Coated Dacron*..........151
4.4 Results and Discussion of Fenestrated ICLs
with Hydrograft' Surface Modifications .......... 167
4.4.1 Determination of Presoaking and
Grafting Conditions ...................... 167
4.4.2 Analysis of Modified ICLs ................ 174

5 SUMMARY AND CONCLUSIONS ................................179
5.1 Surface Modification of Vascular Graft
Substrates ..................................... 180
5.1.1 Gamma Radiation Induced Polymerization
of MMA on Vascular Graft Substrates ......180
5.1.2 Solution Dip Coating and Thermal
Polymerization of PDMS on Dacron*, with
and without Pre-modification with MAOP-
t-PDMS ...................................184
5.2 Surface Modification of Fenestrated PMMA
Intracorneal Lens Substrates ..................186
5.2.1 Gamma Radiation Induced Polymerization
of NVP on PMMA ICL Substrates ............ 186

6 FUTURE WORK ................................... ......... 189
6.1 Surface Modification of Vascular Graft
6.1.1 PMMA Modified PET ........................ 189
6.1.2 PMMA Modified PTFE .......................190
6.1.3 PMMA Modified PDMS .......................191
6.1.4 SEMA Modified PET........................ 192
6.1.5 Solution Dip Coatings on PET.............192
6.2 Surface Modification of Intracorneal Lenses .....193

LIST OF REFERENCES ......... ........................ .....195

BIOGRAPHICAL SKETCH ......................... .................. 207



Table page

2.1 Initial and two week post implant dynamic
compliance values for canine vascular graft

3.1 Molecular weight data for PMMA ocular implant

4.1 Solubility parameters and H-bonding groups for
selected solvents and polymers .......................81

4.2 Percent weight increases of Mylar* D-1000 films in
selected solvents at room temperature ................82

4.3 Percent weight increase of DMSO-MMA solutions by
PET (Mylar* D-1000) films and woven Dacron"
fabrics at 60C. Swelling time is 24 hours for
all samples................. ........................90

4.5 Percent weight changes of Teflon* PTFE irradiated
in MMA-acetone and MMA-DMSO solutions as a
function of radiation dose and solution

4.6 Percent weight changes of GORE-TEX' ePTFE
irradiated in MMA-acetone and MMA-DMSO solutions
to 0.11 Mrad as a function of monomer
concentration ...................................... 103

4.7 Percent weight changes of GORE-TEX* ePTFE
irradiated in 100% MMA as a function of radiation
dose ................................................104

4.8 Contact angle of Teflon PTFE irradiated in MMA-
acetone and MMA-DMSO solutions as a function of
radiation dose and solution concentration ...........111

4.9 Average carbon, oxygen, and fluorine
concentrations for Teflon*, GORE-TEX*, and PMMA as
determined with XPS.................................115

4.10 Weight increase of PDMS irradiated in MMA-DMSO
solutions as a function of monomer concentration
and radiation dose..................................127

4.11 Contact angle data for PDMS irradiated in MMA-DMSO
solutions as a function of solution concentration
(left hand 2 columns) and total radiation dose
(right hand 2 columns) ..............................132

4.12 Percent weight increases of Dacron" dip coated
with PDMS or gamma radiation irradiated in MAOP-t-
PDMS as a function of modification conditions....... 152

4.13 Carbon, oxygen, and silicon atomic concentrations
for Dacron* modified with dip coatings of PDMS and
gamma irradiated in MAOP-t-PDMS as determined with

4.14 Silicon concentrations in solutions from the
vascular graft delamination analysis as measured
by ICP.............................................. 158

4.15 Pressurized porosity analysis of unmodified and
PDMS dip coated Dacron*. Flow rate is reported as
ml/min-cm2 normalized to the sample surface area. ...159

4.16 Platelet counts from ex vivo AV shunt experiments
for unmodified, PDMS dip coated, and MAOP-t-PDMS
modified Dacron* fabric. Platelet counts are
reported as counts/mm2 sample surface area ..........169

4.17 Penetration depths of 10 and 20% NVP into
fenestrated PMMA disks as determined by optical

4.18 Flow rate of water through modified and unmodified
ICLs at room temperature and constant pressure
(atmospheric pressure plus 1 inch of water). Lens
type column gives fenestration hole size and
surface percentage of fenestrations................. 176


Figure page

1.1 Schematic diagram illustrating the anatomy of the
human eye and relavent features .......................8

1.2 Schematic diagram illustrating a cross section of
the cornea with an ICL in place, and diffusion and
flow processes of fluids and nutrients ................9

2.1 Cumulative patency plotted against time for four
types of vascular graft materials ....................17

2.2 Cut away view of an artery and vein showing the
three distinct layers ................................19

2.3 Typical curve for vessel compliance measurements.
Compliance is given as % radial change per
millimeter Hg X 10-2 .................................21

2.4 Possible mechanisms of gamma radiation induced
radiolysis and free radical formation for PET,
PTFE, and PDMS....................................... 44

2.5 Reaction mechanism for thermal curing of Shin-Etsu
PDMS .................................... ............47

2.6 Possible gamma radiation reactions and products
for PMMA............................................54

3.1 Chemical structures of polymer substrates used for
surface modification .................................56

3.2 Chemical structures of monomers used for surface

3.3 Schematic diagram of the irradiation chamber used
for polymerization ...................................62

3.4 Photograph of motorized carousel used to provide
uniform exposure within the irradiation chamber......62

3.5 Schematic representations of the captive bubble
technique and angles measured for contact angle
goniometry showing (a) sample analysis chamber,
and (b) bubble at interface ..........................67

3.6 Schematic representation of the IR beam, crystal,
and sample for FT-IR/ATR spectroscopy ................69

3.7 Schematic illustration of diffusion and flow cell
used for evaluating permeability of ICL

3.8 Schematic illustrations of the setups used for
leakage and stability analysis for vascular
grafts. (a) Leakage analysis setup and (b) flow
system for stability analysis ........................77

4.1 Percent weight increase with time of Mylar* D-1000
PET films in MMA-chloroform solutions as a
function of solution concentration...................83

4.2 Mt/Mm vs. time1/2/1 for MMA-chloroform solution as
a function of solution concentration .................84

4.3 Diffusivity of MMA-chloroform solutions in PET as
a function of solution concentration .................85

4.4 Maximum percent weight uptake by PET of MMA-
chloroform solutions as a function of solution
concentration ......... ...........................87

4.5 Solubility parameter for polyisobutene and
polystyrene as determined by intrinsic viscosity
measurements in a series of solvents................. 88

4.6 Variation of D (diffusion coefficient) and (De)p
(polymer-fixed diffusion coefficient of the
diluent) with volume fraction of benzene for the
natural rubber-benzene system.......................88

4.7 Percent weight increase with time of Mylar* and
Dacron* in MMA-DMSO solutions at 60"C as a
function of solution concentration................... 92

4.8 Cls binding peak of polystyrene showing the n-n*
shake-up peak....................................... 98

4.9 SEM micrograph of Mylar* D-1000 film showing the
dispersion of silica on the surface..................100

4.10 SEM micrograph of Meadox woven Dacron* fabric.......100


4.11 Percent weight increase of GORE-TEX (ePTFE)
irradiated in MMA-acetone and MMA-DMSO solutions
to 0.11 Mrad as a function of monomer
concentration ...................................... 105

4.12 Percent weight increase of GORE-TEX* (ePTFE)
irradiated in 100% MMA monomer as a function of
radiation dose...................................... 107

4.13 Contact angle of Teflon" (PTFE) irradiated in MMA-
acetone and MMA-DMSO solutions to 0.11 Mrad as a
function of solution concentration................. 112

4.14 FT-IR/ATR absorbence spectra of PMMA, GORE-TEX*,
and GORE-TEX* surface modified with MMA.............113

4.15 XPS Cls spectra for Teflon*, GORE-TEX*, and PMMA
showing differences in the chemical shifts due to
various carbon bonds ................................115

4.16 Sample calculation to determine the surface
concentration of PMMA and PTFE on Teflon" and
GORE-TEX' following modification with MMA............116

4.17 Surface concentration of PMMA on GORE-TEX* and
Teflon* irradiated in MMA-acetone and MMA-DMSO
solutions as a function of monomer concentration.
Concentrations of PMMA determined with XPS..........118

4.18 Surface concentration of PMMA on GORE-TEX" and
Teflon" irradiated in 100% MMA to 0.11 Mrad as a
function of radiation dose ..........................119

4.19 XPS Cls spectra for Teflon* irradiated in MMA-
acetone solutions showing changes in the spectra
with increasing surface concentrations of PMMA as
a function of solution concentration ................120

4.20 SEM micrographs of unmodified GORE-TEX* .............122

4.21 SEM micrographs of GORE-TEX* irradiated in 3%
MMA-DMSO to 0.11 Mrad................. ................123

4.22 SEM micrographs of GORE-TEX* irradiated in 20%
MMA-DMSO to 0.11 Mrad. Deformation and stretching
of nodule structure is visible ......................124

4.23 SEM micrographs of GORE-TEX* irradiated in 3%
MMA-acetone to 0.11 Mrad. Breaking of the nodule
structure is typical for samples modified in
acetone solutions..................................125

4.24 Percent weight increase of PDMS irradiated in 10%
MMA-DMSO as a function of radiation dose............128


4.25 Percent weight increase of PDMS irradiated in MMA-
DMSO to 0.11 Mrad solutions as a function of
solution concentration ..............................131

4.26 Contact angle of PDMS following polymerization of
10% MMA-DMSO solutions as a function of radiation
dose ................................................133

4.27 Contact angle of PDMS irradiated in MMA-DMSO
solutions as a function of monomer concentration
polymerized to 0.10 0.13 Mrad.....................133

4.28 FT-IR transmission spectra for unmodified PDMS and
PMMA, and PDMS irradiated in 10% MMA-DMSO to 0.11

4.29 Surface concentration of PMMA on PDMS as a
function of MMA-DMSO solution concentration
polymerized to 0.11 Mrad............................136

4.30 Surface concentration of PMMA on PDMS irradiated
in 10% MMA-DMSO to 0.11 Mrad as a function of
radiation dose......................................137

4.31 1 Hz frequency storage modulus (E') of PDMS
modified with MMA-DMSO irradiated to 0.11 Mrad as
a function of monomer concentration .................139

4.32 1 Hz frequency loss modulus (E") of PDMS modified
with MMA-DMSO irradiated to 0.11 Mrad as a
function of monomer concentration ................... 140

4.33 1 Hz frequency storage modulus (E') of PDMS
irradiated in 10% MMA-DMSO as a function of
radiation dose......................................141

4.34 1 Hz frequency loss modulus (E") for of PDMS
irradiated in 10% MMA-DMSO as a function of
radiation dose......................................142

4.35 1 Hz frequency tan 6 of PDMS irradiated in MMA-
DMSO to 0.11 Mrad as a function of monomer
concentration ..................... ............... 143

4.36 1 Hz frequency tan 8 of PDMS irradiated in 10%
MMA-DMSO as a function of radiation dose............ 144

4.37 Storage modulus (E') plotted against percent PMMA
by weight of PDMS irradiated in MMA-DMSO as a
function of monomer concentration and radiation
dose ................................................145

4.38 CIs spectra showing, (a) a comparison between 60*C
cured PDMS from chloroform solution and high
temperature cured PDMS, and (b) a comparison
between Dacron* and PDMS dip coated Dacron*......... 156

4.39 SEM micrographs of unmodified DacronO fabric
(Meadox) ............................................162

4.40 SEM micrographs of PDMS dip coated DacronO fabric
(Meadox) ............................................163

4.41 SEM micrographs of PDMS dip coated reinforced
Dacron" vascular prosthesis (Bard).................. 164

4.42 SEM micrographs of unmodified Dacron* fabric
(Meadox) after ex vivo AV shunt analysis ............165

4.43 SEM micrographs of PDMS dip coated Dacron* fabric
(Meadox) after ex vivo AV shunt analysis ............166

4.44 Optical micrographs of a fenestrated ICL after
being surface modified with NVP and stained with
silver nitrate ..................................... 170

Abstract of Dissertation Presented to the Graduate School
of the University of Florida in Partial Fulfillment of the
Requirements for the Degree of Doctor of Philosophy



Christopher William Widenhouse

May, 1996

Chairman: Dr. Eugene P. Goldberg
Major Department: Materials Science and Engineering

Synthetic vascular replacements are expanded

polytetrafluoroethylene (ePTFE) and polyethylene terephthalate

(Dacron* PET). Poor long term patency of small diameter

vascular prostheses is attributed to platelet adhesion and the

inability of the vascular endothelium to regenerate. Most

attempts to reduce thrombus also reduce endothelial cell

adhesion, and attempts to promote endothelial cell

proliferation simultaneously facilitate thrombus formation.

Surface graft polymerization of polymethyl methacrylate (PMMA)

and sulfoethyl methacrylate (SEMA) onto prostheses substrates

(PET, PTFE, ePTFE, and PDMS) using gamma radiation induced
polymerization (GRIP) in solutions of dimethylsulfoxide (DMSO)

and acetone produced stable surface grafts of PMMA and SEMA.

This was studied as a method to reduce thrombus formation and

encourage healing. Platelet adhesion to PMMA modified PET was

not significantly different than unmodified PET.

Relationships between polymerization reactions and both

radiation dose and monomer solution concentration were

examined. PDMS was solution dip coated onto PET and thermally

polymerized. Bonding of PDMS coatings onto PET was

accomplished by GRIP of methacryloxypropyl terminated PDMS

onto PET prior to dip coating. Both processes produced stable

PDMS coatings on PET, and inhibited platelet adhesion in ex

vivo canine arteriovenous (AV) shunt studies.
Surfaces were characterized by gravimetric analysis,
contact angle goniometry, Fourier transform infrared
spectroscopy (FT-IR/ATR), X-ray photoelectron spectroscopy

(XPS), inductively coupled plasma (ICP), pressurized flow
analysis, optical microscopy (OM), and scanning electron
microscopy (SEM). Blood compatibility was evaluated by ex
vivo AV canine shunt experiments.
The second part of this research involved intracorneal
lenses (ICLs), which are designed to correct myopia,
hyperopia, and astigmatism. Current designs prevent nutrient
migration through the implant, leading to stromal necrosis and
complications. Fenestrated PMMA ICLs were surface modified
with an N-vinyl pyrrolidone (NVP) monomer presoak, followed by
GRIP providing a lens surface of polyvinylpyrrolidone (PVP).
The hydrophilic PVP surface is designed to facilitate nutrient
flow through the lens, and to provide a tissue protective
layer which is less tissue damaging and cell adhesive.
Modified ICLs were analyzed by contact angle goniometry, XPS,
OM, and diffusion and flow analysis.



1.1 Vascular Prostheses

A vascular prosthesis is a synthetic or natural vessel

used to replace damaged or diseased sections of veins and

arteries. Vascular disease, complications from surgical

interventions, vascular trauma, and disorders which damage

the normal endothelial lining of vessels or impart

thrombogenic complications are often treated with synthetic

or natural vascular grafts (Brody et al., 1972, Epstein,

1988, and Gertler and Abbott, 1992). Currently, the primary

synthetic materials used for vascular grafts are woven or

braided Dacron* polyethylene terephthalate (PET) and expanded

polytetrafluoroethylene grafts, like GORE-TEX*. In the U.S.,

vessel replacements are used over 350,000 times per year

(Ratner, 1993). Improvements in the patency of vascular

grafts and other blood contact devices, such as arterial-

venous (AV) shunts (150,000 per year), heart valves (75,000

per year), and pacemakers (130,000 per year), through the use

of improved materials, would be a major health care advance

(Ratner, 1993).

The success of a vascular graft varies with the inner

diameter of the vessel replacement. Vascular grafts with an

inner diameter greater than 6 mm are considered large

diameter grafts, and those with a diameter less than 6 mm (3-

6 mm) are considered small diameter grafts. The majority of

complications arise from the use of small diameter grafts to

replace coronary and peripheral arteries (Underwood et al.,

1988 and Litwak et al., 1987). Although some intimal

hyperplasia (thickening of the inner vessel layer, especially

at the anastomosis) is observed in the large diameter

prostheses, the strong and sometimes turbulent blood flow

within these prostheses aids in the prevention of total

occlusion, and they remain patent for extended times, often

never necessitating repeat surgical correction (Underwood et

al., 1988 and Litwak et al., 1987). Small diameter grafts,

however, often have up to 50% or more occlusion within the

first 24 months of implantation, resulting in the need for

secondary surgical repairs, or ultimately causing patient

death (Underwood et al., 1988).

The ideal surface for blood contact is the normal

physiological endothelial lining of vessels, and attempts to

mimic this environment through surface modification with

polymers, proteins, cultured cells, and a host of other

approaches, both biologic and synthetic, have been, and

continue to be, explored (Greisler, 1991, Dale, 1978,

Gimbrone, 1987, Ratner et al., 1987, and Hoffman, 1984).

This research attempted to provide a synthetic polymer

surface which would have improved long-term patency and bio-

acceptance, especially for small diameter vascular graft

implants. Long term success depends on two primary factors,

which are often antagonistic. Modifications successful in

reducing platelet cell adhesion and thrombus formation also

discourage endothelial cell adhesion and growth. It is

necessary to inhibit thrombus formation to maintain

prosthesis patency, while endothelialization is necessary for

long term healing. The overall goal of this research is to

develop a vascular prosthesis that has reduced platelet

adhesion and thrombus formation as well as encourages long

term healing and re-endothelialization.

A careful review of vascular prosthesis literature and

experimental results within this University of Florida

research group suggested consideration of polymethyl

methacrylate (PMMA), anionic sulfonic acid containing

polymers such as polysulfoethyl methacrylate (SEMA), and

polydimethyl siloxane (PDMS) surfaces as less thrombogenic

than currently used vascular replacement materials. Surface

modification of currently used materials with PMMA, SEMA, or

PDMS may provide improved blood compatibility for vascular

prostheses. PMMA, SEMA, or PDMS surfaces, also may be

further modified with biological molecules that would serve

to mimic the performance (both chemistry and function) of the

endothelial surface. This second step, although beyond the

scope of the research presented here, is currently being

pursued within this laboratory.

The bulk properties of PMMA, SEMA, and PDMS are

unsuitable for vascular prostheses. PMMA is glassy at body

temperature, and is therefore too rigid--although the idea of

a graft woven from PMMA fibers is intriguing. SEMA is a

hydrogel with poor mechanical properties, and PDMS has a

modulus very close to that of the natural vessel, but has

poor tear strength. Silica fillers are often used to

reinforce PDMS, but the presence of filler decreases

hemocompatibility (Lim et al., 1994). All of these

materials, however, were found to have favorable surface

properties with respect to the vascular blood-contact

environment. This makes these materials excellent candidates

for surface modification onto currently available vascular

graft biomaterials. Surface modification by gamma radiation-

induced polymerization (GRIP) provides a chemically bound

surface layer with desirable physical and chemical properties

on a substrate biomaterial without altering the desirable

substrate bulk mechanical properties. Substrates were

modified in this research with PDMS by thermally curing a

two-part oligomer mixture. This enables surface modification

to be carried out after substrates have been coated with an

uncured PDMS mixture.

Surface modification through GRIP of polymers, monomers,

and other compounds has been studied and used since the early

1940s (Chapiro, 1962). Radiation-induced polymerization of

vinyl functional monomers involves the initiation of free

radicals on both the monomer and the substrate polymer. To

increase the probability of substrate radicals propagating or

terminating with graft and interacting monomer/polymer

radicals (providing true graft polymerization), an intimate

mixture of monomer and substrate is desired. The process is

diffusion controlled, and the use of organic swelling agents

that enhance monomer penetration into the substrate is a

method of achieving desirable monomer-substrate interaction,

and increases the probability of grafting. Another technique

is to increase radical initiation in the substrate by

increasing the radiation dose (Z 0.5 Mrad), and to include an

inhibitor to prevent excessive solution homopolymerization

(e.g., Mohr's salts) in the monomer solution. The addition

of potentially toxic substances, like inhibitors, is not

desirable for production of implantable biomaterials, and

high radiation doses may degrade the mechanical properties of

many polymeric substrates. Therefore, swelling agents were

used to enhance the penetration of the monomer into the

substrate surfaces as described by Yahiaoui's presoakk

method" (Yahiaoui, 1990). Although organic solvents were

used (i.e., chloroform, acetone, etc.), an effort to utilize

more bioacceptable solvents (i.e., dimethyl sulfoxide) was


Simultaneous irradiation of a monomer in solution and a

polymer substrate initiates free radicals in both the

substrate and monomer creating a grafted, branched,

interpenetrating network (IPN) surface region with new

physical properties. Swelling, or presoaking, the polymer

with a monomer solvent mixture prior to irradiation is

helpful to provide a monomer-rich surface region. Swelling

increases the probability of grafting if (i)

homopolymerization of the monomer is favored, (ii) initiation

on the substrate is difficult, or (iii) interactions between

the substrate and monomer are unfavorable for an intimate

mixture. If homopolymerization is favored, the result is

polymerization outside and away from the substrate polymer

chains, resulting in little or no bonding of the new polymer

to the substrate. Presoaking in this situation provides

monomer within the surface of the substrate, which becomes an

IPN or graft polymer when polymerized. If initiation of the

substrate is difficult, an IPN may still be created with pre-

swelling of monomer into the substrate, but little or no

bonding of the IPN to the substrate through graft

polymerization occurs. Solvents also aid in overcoming

repulsive forces between incompatible monomers and

substrates, facilitating the intimate mixture required for

grafting. For example, a hydrophilic monomer in an organic

solvent may wet or swell a hydrophobic substrate better than

the monomer itself. Solvents enhance monomer diffusion to

the surface region and to the active substrate radicals

during polymerization.

It was the goal of the first part of this research to

investigate the surface modification of currently available

vascular prosthesis polymers, namely PET, PTFE, and PDMS, in

an attempt to obtain a more hemocompatible material. The two

major approaches used for surface modification were (i)

simultaneous gamma radiation induced polymerization of

monomer, solvent, and substrate polymer with and without pre-

swelling at relatively low dose (ca. s 0.15 Mrad), and (ii)

solution dip coating of an uncured oligomer mixture followed

by thermally curing the surface polymer.

PDMS is more readily available as a two-component curing

system, and may be polymerized after being formed into almost

any desired device shape. For this reason, solutions of the

oligomer mixture were used to coat substrates by dipping.

Following solvent evaporation, the PDMS was thermally cured

forming a coating on the polymer substrate. To improve the

bonding of the PDMS coating with the substrate, a vinyl-

acrylic functional silicone polymer (methacryloxypropyl

terminated PDMS) was radiation grafted onto the substrate

prior to dip coating. During the curing process, the

silicone components crosslink to each other and to the

surface grafted silicone functional molecules, providing

covalent bonding of the dip-coated surface to the substrate.

The resulting physical, chemical, and mechanical

properties of the modified substrates were characterized by a

variety of techniques including FT-IR/ATR, XPS, SEM, UV-VIS,

optical microscopy, gravimetric analysis, and contact angle

goniometry. Biocompatibility was evaluated using a canine ex

vivo shunt method (c.f., section and in viva

implants (c.f., section, in collaboration with the

research group of Dr. James M. Seeger.

1.2 Intracorneal Lenses

In the late 1940s, Jose Barraquer (Barraquer, 1949)

began experiments on the implantation of synthetic materials

within the corneal stroma of the eye to correct numerous

irregular curvatures and damage caused to the cornea.

Investigations have advanced to clinical studies on a variety

of implant designs and materials, the most widely tested

being intracorneal or intrastromal rings and lenses (ICLs).

Figure 1.1 shows a schematic diagram of the human eye.

Figure 1.2 shows a schematic cross-section of the cornea

region illustrating the placement of an ICL, and outlines the

diffusion and flow of fluids and nutrients in the cornea.

Current designs being studied for refractive corrections to

the cornea include PMMA rings, PMMA lenses, fenestrated PMMA

lenses, and polysulfone (PSf) lenses.

Lens Vitreous



Chamber ~ i


Figure 1.1 Schematic diagram illustrating the anatomy of the
human eye and relavent features, modified from
Corel Draw computer art program.

I 14 ^ Meabolic Pump L tlo Glucose
W 1r% Water SU* ked If WHr removal
In by ImbIblllon Acid
by OP pressure
Figure 1.2 Schematic diagram illustrating a cross section of
the cornea with an ICL in place, and diffusion
and flow processes of fluids and nutrients.
(Taken from McCarey, 1990).

Hydrogel lens implants and PMMA ring implants do not

offer significant refractive correction, and therefore serve

to correct vision through manipulation of the corneal

curvature (Climenhaga et al., 1988). The major problem with

these implants occurs following implantation. The degree of

correction required is calculated prior to choosing a

specific implant, and following the surgery, the scaring and

healing of the cornea is unpredictable, often causing over or

under correction of the curvature (McDonald et al., 1993).

When solid lens implants are used, this problem is avoided by

providing refractive correction to the cornea instead of

relying on curvature corrections, and corneal healing has a

less dramatic effect on the overall success of the implant.

The refractive index of PMMA (1.49) and PSf (1.63) make

the polymers excellent candidates for lens materials, and

these lenses have had some success in recent research studies

(McCarey, 1990). The widespread use of PMMA as an

intraocular lens (IOL) material also makes it a material of

choice. The refractive index and curvature of the lens

design determine the refractive power, or diopter, of the

lens. PSf lenses have the advantage of being thinner and,

therefore, somewhat more flexible than PMMA lenses, and

because of the higher refractive index, a thin PSf lens may

have the same diopter as a thicker PMMA lens. PMMA lenses

have the advantage of being studied more thoroughly than PSf,

and the compatibility of PMMA in the ocular environment is

exceptional (Amon and Menapace, 1990). Both PMMA and PSf

lenses, however, block the flow of vital nutrients to stroma

anterior to the implant (Climenhaga et al., 1988).

Fenestrated lenses with holes of various dimensions, however,

may allow flow and diffusion of saline, oxygen, glucose,

proteins, and other metabolites to the stromal tissues

surrounding the implant.

As demonstrated previously in our laboratories by Osborn

(1985), Hoffmeister (1988), Goldberg et al. (1988 and 1989),

Yahiaoui (1990), Mentak (1993), and Lin (1995), surface

modification of PMMA with polyvinylpyrrolidone (PVP) provides

a hydrophilic surface with many distinct advantages. The

advantages of PVP modified PMMA (Hydrograft*) for IOL

applications include reduced corneal endothelial cell damage,

reduced iris abrasion, and a reduced adhesion and

proliferation of lens epithelial cells within the ocular

environment (Yahiaoui, 1990, Goldberg et al., 1988 and 1989).

A hydrophilic surface on a fenestrated lens should also

improve surface wetting and diffusion of nutrients to the

tissues surrounding the implant, particularly to the region

anterior to the lens.

This research investigated the surface modification of

fenestrated PMMA intracorneal lenses, and the effects on the

permeablity of the ICL to vital stromal nutrients. The

previously mentioned benefits of Hydrograft* modifications

with respect to tissue and cell damage were not investigated

in detail in view of prior research. Complete descriptions

and analysis of the tissue and cell compatibility of the PMMA

Hydrograft* are provided in Yahiaoui (1990), Goldberg et al.

1988 and 1989, Mentak (1993), and Lin (1995).

The cornea receives nutrients from the aqueous humor and

the limbal blood supply (McCarey and Schmidt, 1990), and the

driving force for the non-turbulent diffusion and flow

through to the anterior corneal tissues of the eye is

primarily the chemical potential of the solutions present.

The flow of water, saline, and glucose solutions through

fenestrated PMMA ICLs was studied by the use of a diffusion

flow chamber. Unmodified PMMA and Hydrografts (GRIP modified

PMMA with PVP) modified PMMA ICLs were compared to each other

and to hydrogel lens polymers. Process conditions suitable

for GRIP modification of the fenestrated ICLs also required


A diffusion chamber was used to evaluate the flow of

water, saline, and glucose through fenestrated PMMA ICLs

having different hole dimensions (10 to 80 pm) and surface

area coverage. The solution of interest was placed in one

side, and allowed to flow or diffuse through the

fenestrations of the implant. Differences between the

unmodified ICLs, HydrograftO ICLs, and hydrogel polymers were

observed and recorded.

Although studied in depth previously within our

laboratories, various GRIP modification parameters were

investigated to provide successful Hydrograft* modifications

of the fenestrated PMMA ICLs. The molecular weight of the

PMMA used for these lenses is significantly lower than

materials used previously (80,000 Mw vs. 2-4 million Mw) for

IOLs. Because of this, monomer presoak diffusion times and

temperatures needed to be investigated to determine

conditions which would afford Hydrograft" surfaces without

distorting or crazing the low molecular weight PMMA. The

diffusion of NVP solutions into the low molecular weight PMMA

as a function of time, temperature, and NVP solution

concentration was first investigated to optimize the

presoaking parameters. Once suitable conditions for

Hydrograft modification of the PMMA ICLs were determined,

wettability, graft penetration depth, and saline and glucose

permeability were studied. Finally, hydrogel lenses were

also tested to compare the permeability of fenestrated ICLs

to glucose and saline. The properties of the modified lenses


were determined using contact angle goniometry, optical

microscopy, and the diffusion apparatus previously mentioned.


2.1 Synthetic Vascular Grafts

2.1.1 Vessel Replacement Suraerv and Vascular Grafts

Atherosclerosis is the progressive deposition of plaque

in the arteries with resulting clogging and blood flow

restriction. It can lead to heart disease and stroke, and is

responsible for over 50% of the deaths in the US (Fox, 1987)

Reduced blood flow through the arteries ultimately results in

ischemic heart disease. Reduced flow through the coronary

arteries is a critical condition requiring immediate

attention to reduce the permanent damage of heart muscle by

reduced oxygen supply. This occlusion may also occur in

other areas of the body as well, and threaten permanent

damage to vital regions through reduced blood flow. The

severely affected patient often requires treatment of

coronary and peripheral arteries. Occluded arteries are

generally treated surgically today in one of two fashions,

either with balloon angioplasty or with bypass surgery

utilizing a vascular prosthesis. Balloon angioplasty

involves insertion of a catheter with a balloon which is

inflated in the occluded region. Inflation of the balloon

re-opens the artery, often leaving the thrombus and plaque

free to circulate as emboli, which may clog capillaries and

can cause stroke. Natural Cardiovascular Prostheses

Natural vascular prostheses are classified based on the

origin of the replacement vessel. Autografts are transplants

within the same individual, homografts or allografts are

transplants between different individuals of the same species

(typically a donor organ or vessel), and xenografts or

heterografts are transplants from different species (the most

common of which is the use of porcine heart valves in


Autografts and homografts are the most commonly used

natural prostheses, and include those taken from the

saphenous vein, the umbilical vein, and the mammary vein.

Veins are generally used instead of arteries because the body

has an ability to re-route blood flow through veins more

easily than through arteries without causing permanent damage

to vital regions of the body. Veins also have a slightly

different surface chemistry than arteries which render them

less thrombogenic than arteries. The valves inside the veins

typically are stripped away to eliminate "dead spots" in the

flow caused by the valves, which can lead to thrombus

formation. Synthetic Vascular Grafts

The primary synthetic vascular replacements are expanded

polytetrafluoroethylene (ePTFE), like GORE-TEX*, and woven or

braided Dacron* polyethylene terephthalate (PET) fabrics

(Greisler, 1991). The largest supplier of PET fibers in the

world is DuPont, supplying Dacron*, which was initially

developed for the textile industry. The standards regulating

the production of Dacron* fibers is still largely controlled

by textile industry demands. Dacron& fibers contain titanium

dioxide, which gives the fibers a bright white appearance

demanded by the textile industry, and fibers used for

vascular prostheses, therefore, contain titania as well. The

fibers are woven into a number of different patterns by

different manufacturers. The Dacron* used in this research

was either a non-velour woven Dacron* fabric from Meadox or a

reinforced velour woven knit Dacron* vascular graft from

Bard. GORE-TEX* ePTFE is manufactured by a high temperature,

high speed extrusion process which forms the polymer into a

foam like structure. GORE-TEX* is readily available with

different pore sizes depending on the properties required for

the specific application.

The majority of small diameter ePTFE (GORE-TEX) grafts

and woven Dacron* grafts occlude the vessel by more than 50%

within a 3-year period (Whittemore et al., 1981). A study by

Pevec et al. (1992) shows that all small diameter grafts

occlude within 5 to 10 years. Figure 2.1 shows a re-

tabulation of data presented by Abbott (1987) comparing the

success of natural and synthetic small diameter graft


- Saphenous Vein
--- Umbilical Vein
--0--- PTFE
----- Dacron

Patency (%) 75-




0 2 4 6 8 1012141618202224

Figure 2.1 Cumulative patency as a function of time for four
small diameter vascular graft materials. Data
from Abbott, 1987, and re-plotted for
presentation here.

Surface properties are the primary factors controlling

the acceptance of biomaterials, especially the blood-material

interactions of vascular grafts (Andrade et al., 1991,

Hoffman, 1987, and Mustard et al., 1987). Baier (1969)

showed that the immediate response of the biological

environment to a foreign material involves protein adsorption

within the first few seconds. The initial proteins adsorbed

by the surface and resulting conformational changes affect

the sequence of events that follow. The surface activity of

adsorbed proteins determines which peptide sequences,

clotting factor proteins, and hence, which cells adhere and

attach upon reaching the surface (Ratner, 1993). Control of

these initial events is critical in the development of a

surface which does not activate the complex sequence of

thrombogenic events (Miyauchi and Shionoya, 1988 and De Mol

van Otterloo et al., 1992). The physical and chemical

characteristics of the surface determine how proteins are

initially adsorbed and which factors control the interactions

between other blood proteins and circulating biological


Clotting and intimal thickening, or hyperplasia, is

often attributed not only to surface thrombogenicity of the

prosthesis, but also to the increased stresses resulting from

the mismatch in the radial modulus between the synthetic

replacement and the natural vessel at the anastomosis

(Brothers et al., 1990). (Most physicians label the radial

elastic properties of vessels as "compliance." This is not

to be confused with the engineering definition of compliance

which is the inverse modulus.) Changes in the modulus of the

prosthesis may alter the flow pattern of blood within the

vessel, and may sometimes cause turbulent flow which can

promote thrombosis (Hanson and Barker, 1987).

2.1.2 Properties of the Natural Vessel Physical and mechanical structure of the vessel

The natural blood vessel is a composite structure

comprised of three layers. The outermost layer is the tunica

adventitia or externa, the middle layer is the tunica media,

and the inner layer is the tunica intima (Figure 2.2).


Figure 2.2.

Cut away view of an artery and vein showing the
three distinct layers. (Taken from Fox, 1987).

These three layers are composed of an interwoven network of

collagen, elastin, smooth muscle, and other proteins and

cells, with an innermost layer of endothelial cells

(endothelium). Collagen, elastin, and smooth muscle provide

the vessel with necessary mechanical properties, and the

inner endothelial cell lining provides the surface chemistry

and function of a non-thrombogenic surface. The highly

ordered and abundant fibers of collagen in the adventitia

bear the major part of vessel stresses. This provides the

high modulus and tensile strength of the vessel. Collagen

has a modulus of elasticity of 0.1 to 2.9 X 10 9 Pa (145,000

to 420,000 psi) (Abbott and Cambria, 1982). The high

elasticity of blood vessels are due primarily to the elastin,

which has a modulus of 3 to 6 X 10 5 Pa (45 to 90 psi), a

tensile strength of 0.36 to 4.4 X 10 6 Pa (50 to 650 psi) ,

and an elastic strain of over 300% (Abbott and Cambria, 1982

and Abbott, 1987).

The overall properties of the vessel vary with body

position and distance from the heart. The higher pulse wave

regions closer to the heart damp out the energy with a higher

elasticity and lower modulus. Lower pulse regions do not

require as much damping, and have a lower elastin

composition, less elasticity, and higher modulus. Human

arteries have a modulus of about 1 X 10 5 Pa (15 psi)

longitudinally and 1 X 10 6 Pa (150 psi), circumferentially,
with variations according to location (Nichols and O'Rourke,

1986). Veins typically have a higher modulus than arteries

because the mean pressure within a vein is 2 mm Hg and 100 mm

Hg within an artery (Fox, 1987). The vessel components are

assembled in a complex composite structure that yields

anisotropic mechanical properties.

Most evaluations of vascular mechanical properties

involve measurements of compliance rather than tensile
strength or modulus. Compliance is a measure of the dynamic

circumferential elastic properties as determined by the

relationship shown in equation 2.1,

C ( yD.tolic Ddiutli )
C = (ytl PD i(2.1)
( D.tarc ) x (P~ntolico Pdiatolic )

where Da and Dd are the vessel diameters, and Pa and Pd are

the pressures (Mergerman et al., 1986). This is not to be

confused with the engineering definition of compliance, which

is the inverse of the modulus. Vascular compliance is a

dynamic property, and changes with pressure. A typical

compliance-pressure curve is shown in figure 2.3 for a canine

femoral artery (Mergerman et al., 1986). Matching the

mechanical properties of natural vessels and artificial

grafts is important. A mismatch of properties was reported

to be thrombogenic by Baird et al., (1977), Abbott et al.,

(1987), Hasson et al., (1985), and Kelly et al., (1992). The

mismatch of properties arises from the initial use of stiff

materials and changes in modulus resulting from intimal

hyperplasia, both of which may lead to turbulent flow and

coagulation (Hanson and Harker, 1987). A graft material

replacing the natural vessel must have favorable mechanical

properties, and not induce thrombus and intimal hyperplasia

to remain compliant during use.

25o 1
0 11
e \

60 80 t0 12 1to 16e
Figure 2.3 Typical curve for vessel compliance
measurements., where compliance is % radial
change per mm Hg X 10-2 (Mergerman et al., 1986). Surface properties and functionality of the vessel

The surface characteristics (chemical, physical, and

functional) of the endothelial cell lining make it the ideal

blood contact surface. The dynamic and active roles of the

endothelium and sub-endothelium have only recently been

realized (Gimbrone, 1987). The favorable surface properties

of the endothelium are related to both the production of

proteins and the availability of specific binding sites for

anticoagulant agents and factors. The complex homeostasis

within the vascular environment is an ongoing balance between

formation and lysis of fibrin, the insoluble polymer formed

to repair or close off damaged areas within the arteries.

Exposure of the sub-endothelium, damage to the endothelium,

or simple contact of blood with a non-endothelial surface is

sufficient to initiate platelet activation and the

coagulation cascade (Gimbrone, 1987). Activation causes the

release of mitogens from platelets, endothelial cells, and

monocytes stimulating a host of responses including smooth

muscle cell proliferation giving rise to thrombus formation

and intimal hyperplasia in an attempt to repair or close off

the damaged region (Fox, 1987). Following the formation of

fibrin, the endothelium produces plasminogen activators which

cause the conversion of plasminogen to plasmin, causing

fibrinolysis. Production of plasminogen activator inhibitors

allows control over the extent of lysis (Mustard et al.,

1987). For example, specific surface binding sites are

provided for glycosaminoglycans which associate with

antithrombin III, which in turn inhibits thrombin, a protease

which converts fibrinogen into fibrin, as well as factor Xa,

and possibly factors IXa, XIa, and XIIa (Mustard et al.,

1987). Specific binding sites are also provided on the

endothelium for thrombomodulin, which readily binds thrombin

(Mustard et al., 1987). The endothelium of a normal, healthy

individual can typically repair damaged segments of 1 to 2 cm

in length (Greisler, 1991). Neo-endothelialization (re-

growth of the natural endothelium) of a vascular prosthesis

which allows the retention of normal specific activities and

functions of the endothelium would be advantageous to the

success of vascular grafts.

2.1.3 Synthetic Vascular Prosthesis Materials and Properties

The primary synthetic vascular graft materials are PET

and expanded PTFE. PDMS has also been studied. The surface

and bulk properties of these materials contribute to the

success or failure of these materials as vascular prostheses.

The improvement of the surface properties of these substrate

materials through surface modification to make a prosthesis

that is nonthrombogenic and which is capable of healing was

the focus of this research. Properties and modification of PET

PET has been used as a vascular graft material

clinically for over 40 years, and this is the primary reason

it was studied here for surface modification (Brothers et

al., 1990 and Hufnagel, 1955). PET is used for vascular

grafts in the form of woven and knitted fibers. Because of

the higher porosity of some weaves, some grafts require pre-

clotting prior to use. Grafts with porosity greater than 300

to 400 ml/cm2 min at 120 mm Hg require pre-clotting to prevent

leakage (Greisler, 1991).

Typical mechanical properties of PET depend on the

degree of crystallinity and orientation. PET vascular grafts

are composed of woven fibers which are highly ordered and

contain approximately 35 to 60% crystallinity (Rodriguez,

1982). PET fibers have a glass transition temperature, Tg, of

80'C and a melt temperature, Tm, of 245 to 265C. Typical

mechanical properties of unoriented PET are 4.8 to 6.9 X 107

Pa (7,000 to 10,000 psi) tensile strength, 4.1 to 4.8 X 109

Pa (400,000 to 600,000 psi) tensile modulus, and a strain to

failure of 30 to 300% (Rodriguez, 1982 and Roslink, 1990),

although tensile strength values for oriented fibers can

exceed 100,000 psi. (Unlike the value reported earlier for

elastin, 300% elongation of PET is non-elastic.) Compared to

the natural vessel, PET has exceptional mechanical strength,

and in the form of a woven fabric material, has flexibility

and kink resistance. Radial and longitudinal modulus,

however, remain much greater than the natural vessel. The

need for a non-thrombogenic surface is therefore important to

reduce possible intimal thickening caused by mechanical


PET is a polyester, and polyesters are subject to

hydrolysis of the ester linkage. The rigid nature of the

backbone chain and high crystallinity make PET fibers more

resistant to hydrolysis than many other polyesters, and PET

is reported to be relatively stable in the biological

environment compared to other polyesters (Rodriguez, 1982).

Successful surface modification of PET by gamma

radiation polymerization has been reported (Stannett, 1981,

Rebenfeld and Weigmann, 1978, Kale and Lokhande, 1975, and

Nair et al., 1988). The use of PMMA and similar monomers

such as acrylic acid, di-methyl acrylamide, hydroxyethyl

methacrylate (HEMA), and other vinyl monomers have reportedly

been grafted onto PET. Polyester fiber stability during

gamma radiation has been reviewed by Nair et. al (1988).

Doses of 2.5 Mrad increased the crystallinity from 40 to 44%

and increased the breaking load of the fibers by 2.3%. The

changes, although minor, are attributed to cross-linking and

some degradation. The low doses used in this research (ca. s

0.15 Mrad), however, are not expected to significantly alter

the mechanical properties of PET.

PET has been reported to be an activator of the

complement system, and the addition of a barrier layer such

as a PMMA surface graft or PDMS coating may inhibit this

reaction (Shoenfeld et al., 1988, and Miyauchi and Shionoya,

1988). Figure 3.1 in section shows the chemical

structure of PET as well as the other substrate polymers used

in this research. Properties and modification of ePTFE

Hufnagel (1955) reports the use of expanded PTFE as a

vascular prosthesis as early as the late 1940s. PTFE is also

used in catheters, bone joint prostheses, and soft tissue

implants, and has been studied as a paste material to replace

endoscopic balloon inserts and other devices (Ratner, 1993,

Bonomini et al., 1969, and Atala et al., 1992). PTFE is a

hydrophobic, very low surface energy fluoropolymer (19

dynes/cm) which appears less thrombogenic than PET. PTFE is

generally used for vascular grafts in the expanded form as

ePTFE (GORE-TEX*), and implanted grafts of expanded PTFE and

Dacron" invoke a similar biological response (c.f., figure

1.1). Although porous, the high surface energy and

hydrophobic surface of ePTFE grafts prevent leakage during

implantation, and ePTFE grafts do not require pre-clotting.

The porous surface of the material is reported to be

advantageous, allowing greater flexibility with tissue and

cellular in-growth, and the adhesion and formation of a

neointimal layer on the graft surface (Sprugel et al., 1987).

However, the inability of certain materials to adhere to PTFE

may prevent the necessary interactions between circulating

factors and the substrate, hindering the re-

endothelialization and healing process. The major

manufacturers of ePTFE vascular grafts are Gore and Impra.

The process to form expanded PTFE first patented by W.

L. Gore & Associates to make GORE-TEX* entails extrusion of

the material at high temperature (ca. 390'C) and high strain

rate (ca. 40,000 %/sec), producing a microstructure of

interconnecting nodes of 0.5 to 400 pm and fibrils of 5 to

1,000 pm and a porosity of approximately 85 to 95% by volume

(Gore, 1970 and 1975). A high Tm of 310 to 330'C allows the

expansion process to take place without significant

degradation of the PTFE, providing the material with higher

flexibility. PTFE has a tensile strength of 1.4 to 3.5 X 107

Pa (2,000 to 5,000 psi), a modulus of 4.0 to 5.5 X 108 Pa

(58,000 to 80,000 psi), and an elongation of 200 to 400%

(Rodriguez, 1982 and Roslink, 1990).

Surface modification attempts on ePTFE have often

utilized radio frequency plasma polymerization or oxidation

of the surface as an initial step, but the use of gamma

radiation alone has been shown to be effective. Pre-swelling

the surface with a monomer-solvent mix reportedly yielded

successful gamma induced surface modification with N-

vinylpyrrolidone in pyridine (Sayed et al., 1981). Razzak et

al. (1987) reported successful radiation grafting of N,N-

dimethylacrylamide (DMAA) onto PTFE using ethyl acetate and

acetone as solvents. Pre-irradiation of the substrate to

induce a high concentration of free radicals followed by

addition of monomer has also been used. However, the higher

doses required for this pre-irradiation step often degrade

PTFE. This research utilized very low doses and simultaneous

irradiation of monomer and substrate in an attempt to surface

modify PTFE without significant degradation.

Expanded PTFE has mechanical properties similar to the

natural vessel with respect to modulus and elasticity. A

surface graft to improve the biocompatibility is not expected

to significantly alter this advantage of the material, but to

provide a more compatible surface through the altering of the

surface chemistry. Properties and modification of PDMS

Of the substrate materials studied, PDMS is the only

elastomer, and therefore has the potential for a close match

to the modulus of the natural vessel. Silicones are used as

soft tissue implants, IOLs, skin replacements, burn and wound

dressings, catheters, mammary implant components, and for

other devices (Ratner, 1993, Taylor, 1985, Meaburn et al.,

1978, Tsai et al., 1991, Christ et al., 1989, and Grabow,

1991). PDMS has not been used alone clinically as a vascular

graft material, but has been studied for the application as

early as 1955 (Hufnagel, 1955). PDMS has also been used by

the National Institutes of Health/Heart, Lung, and Blood

Institute (NIH/HLBI) as a low-thrombogenicity reference


PDMS is typically used in the form of a crosslinked

elastomer for applications demanding mechanical strength.

(Lower molecular weight and lower cross-link density oils and

gels find other applications.) The tensile and tear strength

of silicone is somewhat lower than other vascular graft

materials, and reinforcement is needed. Reinforcement is

achieved with either silica filler particles dispersed within

the silicone, or continuous fiber reinforcement. Typical

mechanical properties of silica filled and unfilled PDMS

elastomers fall in the range of 2.4 to 6.9 X 106 Pa (100 to

1200 psi) tensile strength, 1 X 105 to 1 X 107 Pa(15 to 1450

psi) modulus, and 20 to 700% elongation. The high elasticity

and low Tg (ca. -123 *C) of silicone elastomers afford

modulus values similar to the natural vessel. Table 2.1

shows compliance measurements made by Abbott and Cambria

(1982) for vascular graft materials as they compare to the

natural vessel before and two weeks after implantation.

However, one major concern for PDMS has been the long-term

changes which may accompany the known affinity of PDMS for

lipids, which results in a more brittle material.

Silicone elastomers used as IOLs and mammary prosthesis

shells could benefit from surface modification as well. The

most successful material used for IOL applications to date

has been PMMA (Joo and Kim, 1992, and Balyeat et al., 1989).

The new small incision techniques developed to insert

foldable IOLs with less invasive surgery has utilized

silicone as the implant material (Christ, 1989 and Grabow,

1991). If PMMA surface properties were to be applied to

silicone elastomers without increased rigidity, the success

of the PMMA implants might be achieved with modified PDMS

using less invasive techniques.

The recent criticism of silicone gel mammary implants

has created a frenzy of research for alternative materials.

Although the major problems are gel bleed of the silicone gel

and low molecular weight oligomers through the elastomer

shell and rupture of the shells due to gel swelling and

possible biodegradation, some research shows adverse tissue

reactions, such as fibrous capsule formation and hardening of

the implant, to be problematic as well (Habal et al., 1991).

A PMMA surface graft may reduce the adverse tissue response

based on the more favorable biocompatibility of PMMA.

Because of the difference in solubility parameters between

PMMA (9.4 (cal/cm3)1/2) and silicone (7.3 (cal/cm3)/2)

reduction of gel bleed in PMMA modified mammary prostheses is

expected (Rodriguez 1982).

Table 2.1 Initial and 2 week post implant dynamic
compliance values for canine vascular graft
materials. Compliance is given as percent radial
change per mm Hg (X 10-2) Data taken from Abbott
and Cambria, 1982.

Graft Implant External Initial 2 Week
Material Diameter Compliance Compliance
Normal Femoral Artery 4.69 0.007 5.86 0.26
Femoral Arterial Graft 4.53 t 0.50 4.41 0.80 5.67 1.1
4 mm ID DacronO 6.41 t 0.16 1.46 0.12 1.19 0.1
PTFE 5.08 t 0.05 5.08 0.05 1.3 0.4
Silastic Rubber 6.06 0.50 5.95 0.09 6.2 0.7
Polyurethane 5.27 0.15 6.10 1.10 2.3 0.4

Successful surface modification of PDMS elastomers using

gamma radiation with hydroxyethylmethacrylate, (HEMA),

ethylene glycol dimethacrylate, and copolymers has been

reported by Heaburn et al. (1978). Hoffman et al. (1983)

also report grafting of HEMA and ethyl methacrylate polymers

and copolymers onto silicones. Lin (1995) has successfully

modified PDMS elastomers, Shin-Etsu silicone in particular,

with PVP.

2.1.4. Advantages of PMMA. SEMA. and PDMS Surfaces Advantages of a PMMA surface

PMMA and other acrylic polymers have been used as

biomaterials with much success. Acrylics are currently used

for intraocular lens (IOL) implants (Amon and Menapace,

1990), hemodialysis membranes (Falkenhagen and Brown, 1991),

dental resins, and bone cements, and are studied for blood

compatibility for general biomaterial applications (Apple et

al., 1984, Feuerstein et al., 1991, Feuerstein et al., 1992,

Lentz et al., 1985, Ito et al., 1992, and Sherman et al.,

1963). A general summary of the recent literature on PMMA

and other acrylics would be to label these materials as

relatively bioacceptable for many invasive applications.

Apple et al. (1984) provide an excellent review of the

studies done by Ridley and others for work with PMMA

concerning IOL implants.

Acrylic chemistry has also been reviewed because of

concerns regarding hydrolysis of the ester-carbonyl bond in

the side groups of polyacrylates such as PMMA. Acrylics are

surprisingly resistant to degradation, however, and acrylic

elastomers such as ethyl acrylate have been used widely

because they are more resistant to oxidation, ultraviolet

light damage and hydrolysis than other traditional elastomers

(Rodriguez, 1982 and Sperling, 1986). Long alkyl pendant

groups provide less stability. Thus, methyl methacrylate is

more stable than ethyl, butyl, and isopropyl acrylate, and

the stability decreases with increasing pendant group size.

The best example of the vascular compatibility of PMMA

was presented many years ago at the 1954 annual meeting of

the Society for Vascular Surgery. Hufnagel (1955) reported

that tubes of methyl methacrylate polymer (Lucite) had been

used as thoracic artery replacements in canines for up to 6

years with complete maintenance of patency. Other materials

tested in this study included polyethylene, Teflon*, Kel-F,

nylon, woven stainless steel mesh, Vitallium, other metals,

and silicone rubber. These materials were tested with and

without silicone coatings, and no material was as successful

with respect to low thrombogenicity as methyl methacrylate

(Hufnagel, 1954). The silicone coated prostheses had poor

compatibility, which Hufnagel attributed to incomplete

removal of acids formed in the curing reactions.

The use of methyl methacrylate as a vascular prostheses
has not been reported in the intervening years, probably

because the need for flexible materials has eliminated

consideration of this rigid polymer. Even though Hufnagel

collected data on over 400 implants of rigid vascular

replacements, he indicated the importance of using a material

which had mechanical properties similar to the natural

vessel. Work reported from the late 1950s to date has

therefore shifted to softer, more flexible, and more

hydrophilic surfaces. The potentially excellent

nonthrombogenic surface properties of PMMA thus, seems to

have been lost in the literature. Since the technology now

exists to graft polymerize very thin glassy polymer surfaces

onto flexible substrates without significant reduction in

elasticity, the surface modification of vascular and blood

contact devices now seems logical and promising.

Current work in our laboratory indicates PMMA has

favorable cell adhesion properties with respect to platelet

and endothelial cell adhesion (Goldberg et al., 1988-1995).

A reduction in the ratio of adhered platelet cells to

endothelial cells is sought for blood contact devices such as

a vascular grafts. Observations in our laboratories of in

vitro platelet and endothelial cell adhesion assays show

unmodified PMMA control samples to have a more favorable

ratio (greater endothelial cell adhesion than platelet

adhesion) than other materials studied, including surface

modified PMMA (Goldberg et al., 1988-1995). The affinity of

PMMA for endothelial cells becomes more apparent when

considering research on PMMA for IOL applications. A common

problem of unmodified PMMA IOL materials is excess epithelial

cell adhesion and growth on the implant (Goldberg et al.,

1988-1995, Yahiaoui, 1990, and Lin, 1995). Assuming that

endothelial and epithelial cells exhibit similar surface

adhesion properties on polymers, IOL studies suggest that

PMMA would favor endothelial cell adhesion. Finally, in

studying a new method to measure the adhesive strength of red

blood cells (RBCs) to biomaterials, Bowers et al. (1989)

found less adhesion of RBCs to control PMMA samples than to

hydrophilic glass, tissue culture grade polystyrene, and PET. SEMA surface advantages

Sulfonated surfaces have been investigated in the

vascular environment primarily because of the strong anionic

surface charge contributed by sulfonic acid functional

groups. Many blood components, including red blood cells and

platelets, have a slightly negative surface charge. It

therefore seems logical, that surfaces with negative charges

will repel these components of circulating blood. Our

laboratory has been investigating the use of sulfonated

monomers for surface modification such as potassium 3-

sulfopropyl acrylate (KSPA), sodium methacrylate (SMA),

styrene sulfonic acid, sodium salt (SSA), and other anionic

sulfur containing monomers since the late 1980s (Goldberg et

al., 1988-1995 and Yahiaoui, 1990). These monomers are

supplied as water soluble sodium or potassium salts. It is

often difficult to obtain an intimate mixture of monomer and

substrate molecules through presoaking in aqueous media,

because of the slightly hydrophobic surface of PET. Pre-

treatment with other monomers such as N-vinyl pyrrolidone

(NVP) is often required to obtain grafting of these monomers.

However, some sulfonated monomers, such as 2-sulfoethyl

methacrylate (SEMA), are available as organic, non-salt

compounds, and are therefore soluble in organic solvents such

as acetone and dimethyl sulfoxide (DMSO). Dissolution in

these solvents increases the swelling of substrates during a

presoak process, and grafting is more easily achieved.

Critical components of the vascular environment include

chondroitin sulfate, heparin sulfate, and dermatan sulfate,

which are all proteoglycans secreted by the endothelium.

Marcum et al. (1986) reported heparin sulfate is capable of

binding with antithrombin, and thus imparts antithrombogenic

properties to the vascular endothelium. Ofosu et al. (1989)

report that increasing the degree of sulfonation on heparin

and dermatan sulfate increases the catalytic effects on

thrombin inhibition. Other compounds with functional sulfate

groups are thought to have favorable reactivity within the

vascular environment. Kishida et al. (1991) studied

polyethylene films surface modified with cationic, anionic

(sulfonated and non-sulfonated), and non-anionic monomers,

and found the in vivo cell adhesion to be related to both the

charge and the presence of sulfonated groups, with sulfonated

polymers having higher HeLa S3 cell attachment and growth

than non-ionic and cationic surfaces, indicating a higher

affinity for binding, attachment, and growth of cells to

anionic surfaces. The incorporation of sulfonated functional

polymers on vascular grafts seems promising in reducing

thrombus formation. PDMS surface advantages

PDMS is a hydrophobic material with a low surface energy

(contact angle ca. 80). The low surface energy prevents

adhesion of many compounds from an aqueous environment, as

these molecules are often repelled by silicone surfaces.

Two key proteins determining the thrombogenicity of a

biomaterial are albumin and fibrinogen. Albumin adsorption

is preferred for a non-thrombogenic surface, and fibrinogen

adsorption usually indicates a thrombogenic surface, as

fibrinogen is converted to fibrin by thrombin. Cooper and

Fabrizius-Homan (1991) found silicone rubber to have a higher

affinity for albumin than fibrinogen in competitive

adsorption studies, and when the albumin was preferentially

absorbed, the thrombogenicity of the material was reduced.

This ratio of albumin to fibrinogen was most favorable for

silicone when compared to polyethylene, polyurethane, and

Teflon*. In a canine ex vivo AV shunt platelet adhesion

study by Ip and Sefton (1991), it was found that SilasticO

(PDMS, Dow Corning) and silica free PDMS (Thoratec) both had

significantly lower platelet cell adhesion than polyethylene.

Norgren et al. (1990) found silicone coated Dacron& to

have a reduced thrombogenicity, and Granke et al. (1993)

found reduced inflammatory reaction as well. In the study by

Granke, only the outer surface of the graft was covered to

create a prosthesis which did not require pre-clotting, but

there was significantly more tissue ingrowth and

endothelialization in the silicone treated samples than in

the control samples. Whalen et al. (1992) evaluated a novel

prosthesis made entirely of silicone. The prosthesis was

porous to allow tissue ingrowth, and after implantation in

canines for 8 weeks, had an overall patency of 86%.

Two studies within the research group of Cooper (Lin et

al., 1994 and Silver et al. 1995), report modification of

silicone elastomer surfaces in an attempt to improve the

biocompatibility. Lin et al. (1994) compared silicone to

polyurethane-silicone copolymers and found the compatibility

of the silicone to be superior with respect to clotting time

and platelet deposition. Silver et al. (1995) modified

silicone surfaces with alkylsiloxane monolayers of various

functionalities following exposure to an oxygen radio

frequency plasma. The materials were evaluated in a canine

ex vivo arteriovenous (AV) shunt, and the untreated silicone

had superior properties with respect to platelet and

fibrinogen deposition. In a study by Morel et al., (1989)

the endothelial cells cultured on thin silicone sheets were

found to have motility and contractility indicating a healthy

environment for cell growth and proliferation. Finally,

numerous studies using a canine AV shunt use Silastic, or

some other silicone elastomer, as the tubing in which samples

being analyzed are placed (Lin et al., 1994, Silver et al.,

1995, Goldberg et al., 1985-1995). The silicone tubing would

not be used in these studies if occlusion occurred.

The recent controversy surrounding the use of silicone

gel filled breast implants will obviously invoke a negative

response to the idea of using silicone as a vascular graft

surface. However, the silicone surfaces discussed and

studied in this research are crosslinked elastomers, and the

toxicity and poor biological response to silicone materials

related to breast implants are for the low molecular weight

oils and gels (Kimitoshi et al., 1990 and Kimitoshi et al.,


Current literature and research indicate a silicone

surface in the vascular environment could have favorable and

beneficial responses. This review led to the investigation

of PDMS coatings on Dacron' vascular prostheses presented


2.1.5 Gamma Radiation Initiated Polymerization

Gamma-rays are electromagnetic waves of short wavelength
(X= < 0.1 nm) which are emitted from a decaying radioactive

source. One of the most commonly used sources of gamma-ray

energy, or gamma radiation, is the radioactive isotope of

cobalt, cobalt-60 (60Co). Providing two sharp spectral lines

of radiation energy of 1.17 and 1.33 MeV (megaelectron

volts), 60Co is often chosen because of its ease of

preparation (nuclear activation of cobalt-59) and its long

half life of 5.3 years (Chapiro, 1962).

Gamma radiation energy generates free radicals on vinyl

monomers and polymeric substrates, making it an excellent

initiation source for free radical polymerization of monomers

onto polymeric surfaces (Chapiro, 1962). Gamma radiation

initiation polymerization (GRIP) has been studied since the

late 1940s, and is often chosen over other techniques because

of its low cost and cleanliness (does not introduce chemical

initiator molecules). Low cost and cleanliness are two

extremely critical factors in determining the success of many

materials for the biomedical industry.

Surface modification through GRIP involves the formation

of free radicals on monomer molecules leading to free radical

polymerization. The simultaneous exposure of the monomer and

substrate allows both initiation and cleavage of substrate

polymer. The growing homopolymer chains may propagate or

terminate with the available reactive sites created on the

substrate, creating a surface region of grafted, crosslinked,

and interpenetrating network molecules. Swelling the

substrate with monomer in solution presoakk technique)

provides a localized monomer-rich region within the substrate

surface, and the surface region becomes an intimate mixture

of substrate and monomer molecules (Yahiaoui, 1990). Upon

exposure to gamma radiation, this entire region becomes a

surface "graft," referring to a region of grafted and IPN

polymer molecules. The intimate mixture of the swollen

region also facilitates the diffusion of the monomer and

homopolymer chains to the activated substrate sites,

increasing the efficiency of the grafting. Polymer radiolysis and free radical reactions

Chapiro (1962) provides an excellent review of early

studies of radiation effects on polymers and monomers. Upon

exposure to high energy radiation, polymers undergo a series

of reactions including, but not limited to, radiolysis of

side chain atoms and functional groups, free radical

formation, crosslinking, chain scission, and degradation.

Although the exact nature of the mechanism of these reactions

with systems as complex as polymers is still not fully

understood, radiation polymerization following initial

radiation events is fairly well documented, and several

events may occur upon simultaneous exposure of monomer

solutions and polymer substrates.

When a polymer or monomer is exposed to gamma radiation,

cations, anions, and free radicals are created. The ions are

only stable at low temperatures, and usually dissociate to

yield radicals (Chapiro, 1962). Graft polymerization may

take place when a polymer and monomer in intimate contact are

simultaneously exposed to high energy radiation. Initially,

a radical may be formed on the substrate polymer or the

monomer. When a monomer molecule reacts with one of these

two radicals, either a graft polymer or homopolymer forms,

respectively. Graft polymerization is favored if the polymer

substrate has a higher ability to cross-link after chain

scission rather than degrade, and more grafting occurs if

radical yields are higher for the substrate polymer than the

monomer (Chapiro, 1962).

Radicals are also created on the growing homopolymer

chains, giving rise to higher molecular weight homopolymers,

branching, or crosslinking. Radical formation on side groups

of growing homopolymer chains or substrate chains yields

branching, and termination of radicals by combination with

branched chains from different molecules yields crosslinking.

These polymer reactions, if occurring within the substrate,

lead to the formation of an IPN.

Substrate main chains may be cleaved as a result of the

high energy of the radiation. The cleaved chains then form

crosslinks with other cleaved substrate chains; form graft

polymers or cross-links by the addition of monomer or

combination with a growing homopolymer chain; or re-combine

with the original site of cleavage (Yamamoto and Yamakawa,

1980). If cleaved chains terminate by disproportionation or

chain transfer, the molecular weight of the substrate polymer

is reduced, and degradation results.

During the course of the polymerization, the polymer

within the substrate surface begins to gel as higher

molecular weights are reached and as cross-linking and

branching occur. This decreases the mobility of larger

propagating molecules. The concentration of monomer

decreases and the concentration of radicals increases rapidly

as higher conversions are reached (Dob6, 1978), leading to

higher degrees of cross-linking and branching in the surface

polymer than in the surrounding solution polymer, as the

reaction auto accelerates.

The efficiency and stability of radical formation on

various substrates and monomers determine the ease or

difficulty of polymerization of a specific monomer to a

specific substrate polymer surface. The radiolysis and

radical formation on the various substrates and monomers

studied in this research will now be discussed. Radical formation on PET

Charlesby (1953) found PET to crosslink upon exposure to

gamma radiation, whereas Todd (1954) found it to undergo

degradation. Low dose exposures always gave an increase in

modulus of Dacron* fibers when studied by Teszler and

Rutherford (Chapiro, 1962). Radical formation on the phenyl

ring without atomic ejection is possible, but low yields are

expected because of the resonance of the ring (Chapiro,

1962). Cleavage of the ester bond or radical formation on

the main chain with ejection of an H atom are the most likely

reactions, as shown in figure 2.4 (Chapiro, 1962). Radicals

are formed on both the amorphous and crystalline regions of

the polymer, but polymerization is only expected in the

amorphous regions because of the reduced diffusion through

the crystalline phase. Radical formation on PTFE and ePTFE

Main chain cleavage is possible in PTFE, but radiation

more commonly yields free radicals from cleavage of the C-F

bond (Rye, 1988). Charlesby (1952) analyzed gas evolution of

PTFE upon exposure to gamma radiation and found the products

to be carbon tetrafluoride, indicating evolution of both

carbon and fluorine. In the presence of oxygen, doses as low

as 1 to 10 Mrad lead to significant degradation and

embrittlement of PTFE. When oxygen is excluded from the

system, however, much less damage occurs. Possible reaction

mechanisms are shown in figure 2.4. Radical formation on PDMS

Radical formation on silicone polymers was studied by

Charlesby and Omerad (1963), and the polymers were found to

crosslink upon exposure to gamma radiation. The Si-O bonds

are significantly more stable than the Si-C and C-H bonds,

leading to radiolysis of the pendant methyl groups. Possible

reactions are shown in figure 2.4. Radical polymerization of vinyl monomers

Methyl methacrylate monomer readily forms radicals when

exposed to gamma radiation, and is expected to polymerize

well on all of the chosen substrates. The resulting surface

polymer is expected to be an interpenetrating network of

grafted, cross-linked, and branched polymer.

---( --OH *CH--CH--O--

0O O
- -C-0 ...)-C-CH-CH2a-I-o-

Radiolysis of PET

--CFT-CFZ-CF--- Radiation -C C or --Cr

Radiolysis of PTFE

S- Gamma
--Si-Of-- Radiationb

-- &3 &3

Radiolysis of PDMS
Figure 2.4 Possible mechanisms of gamma radiation induced
radiolysis and free radical formation for PET,
PTFE, and PDMS. Polymerization of vinyl monomers onto substrates

Chapiro (1962) discusses the free radical polymerization
of vinyl monomers induced by gamma radiation. The mechanism
of reaction proceeds via traditional free radical processes,

with the rate of polymerization being proportional to the

monomer concentration and the square root of the dose rate.

The equation for the kinetic chain length (average number of

monomer molecules polymerized per initiating primary radical)

follows that of chemical free radical polymerization as well.

The kinetic chain length for gamma radiation polymerization

is proportional to the square of the monomer concentration

and inversely proportional to gamma radiation dose rate.

2.1.6 Polymer Solution Coatings and Techniques

When a solid material is removed from a liquid, (solution or

pure liquid) the surface attractive forces are typically

strong enough to adhere a monolayer or more of the liquid to

the solid surface. The relative surface energies of the

solid and solution determine the thickness and adhesion of

this surface coating. If a solute is dissolved in the

solution, and the solvent is rapidly evaporated after the

object is removed from solution, there will be a layer of

solute on the object. The relative surface energies of the

solute, solvent, and object as well as the evaporation rate

and solution concentration determine the amount remaining on

the surface. The surface energies of the solute and solid

determine the strength of the adhesive bond between the

surface layer and the object. This general principle of

surface phenomena was utilized in the dip coating of PDMS

onto Dacron* vascular prostheses.

Both PET and PDMS are relatively hydrophobic (contact

angle of PET ca. 60, and contact angle of PDMS ca. 80'), and

both materials swell in the organic solvent chloroform. For

this reason, solutions of the two part PDMS oligomers were

made in chloroform, and Dacron was soaked in this solution.

The swelling of the DacronO by the chloroform is believed to

allow the diffusion of PDMS molecules into the PET surface,

creating a swollen network of PET, PDMS, and chloroform.

Upon removal of the PET from the PDMS solution, the

chloroform evaporates, de-swelling the surface network,

leaving PDMS oligomers on and in the PET surface. The

relatively rapid evaporation of chloroform, and the similar

surface energies of PDMS and PET allow a layer of PDMS

oligomers to remain on the prosthesis after removal from the


In an attempt to improve the bonding between the PDMS

and Dacron*, a pre-modification step was implemented. MAOP-

t-PDMS was gamma polymerized onto Dacron* to provide a link

between the coating and the substrate. A covalent link

between PMMA-g-PET and the MAOP-t-PDMS is expected by using a

presoak of MMA followed by polymerization of MAOP-t-PDMS.

The incorporation of MMA into the system will provide a

propagating link between the substrate and the MAOP-t-PDMS

thus allowing polymerization of more MAOP-t-PDMS. The dip

coating of PDMS should then polymerize and crosslink with the

MAOP-t-PDMS. Thermal curing of PDMS

The curing of Shin-Etsu PDMS proceeds via vinyl addition

polymerization, and is initiated by a platinum catalyst. The

reaction scheme for this polymerization is shown in figure


----Si---Ca=CH + H-Si -CH
CH3 0 CH3
C13 -Si -H + CH2 ==CH -Si ---+
CH, o c13
-Si-CH=CH2 + H-Si--1CH3

Pt Catalyst

(--O-Si --- -'---Si ---M,
CH3 0 CH3
CH3-Si -H, -01k -Si -0-0
CH3 0 -13
-(-0-Si -O-r---lr--Si --13

Figure 2.5 Reaction mechanism for thermal curing of Shin-
Etsu PDMS.

2.2 Intracorneal Implants

2.2.1 Refractive Corneal Surgery and Intracorneal Implants

Myopia (nearsightedness caused by excessive corneal

curvature), hyperopia (farsightedness caused by reduced

corneal curvature), and astigmatism (irregular corneal

curvature) are disorders affecting the cornea which cause

impair vision. Accidental injuries that abrade, scratch,

deform, or scar the cornea also significantly reduce visual

acuity. These disorders and damages of the cornea are

currently corrected by the use of spectacles (glasses or

contact lenses), a corneal transplant for severe conditions,

or with some of the newer surgical techniques such as radial

keratotomy (RK) or laser keratotomy. Intracorneal lenses or

rings provide still another option, and are being studied by

numerous researchers and ocular device companies.

Glasses and contact lenses will always be an option for

correcting many corneal deformations because they have the

advantage of always being reversible. If a problem occurs or

an error was made in fitting the diopter of the corrective

lens, the device may simply be removed. Corneal transplants,

are a major surgical procedure, and a donor organ is

required, restricting their use to severe cases of complete

corneal damage or damage which can not be repaired by other

means. Radial and laser keratotomy scars into the cornea and

relies on healing to alter the shape and modulus of the

cornea, and is an extremely unpredictable procedure (Waring,

1990). The use of RK and other similar procedures offers

patients a quick procedure with extreme pain and little

guarantee of success, as glasses are often still required.

However, it is often difficult to treat severe

astigmatism with contact lenses, and greater deformations

require large and often cumbersome spectacles, and damage or

deformations too great for other means of corrections often

requires treatment with a non-reversible therapy (Kerry,

1995). Keratoprostheses such as ICLs offer a more permanent

solution to refractive corrections of the cornea.

In the late 1940s, Barraquer began experiments on the

implantation of synthetic materials within the corneal stroma

of the eye to correct irregular curvatures (Barraquer, 1949).

The initial designs were lenses constructed of PMMA and

polysulfone that were surgically implanted within the corneal

stroma to provide both mechanical (corneal reshaping) and

refractive (lens power magnification) correction of the

cornea. More recent designs include hydrogel lenses,

intracorneal rings, and fenestrated PMMA lenses.

2.2.2 Intracorneal Lens Materials and Designs

Polysulfone is an optically clear, amorphous, stiff, but

flexible polymer with a glass transition temperature (Tg) of

190'C and a refractive index of 1.63. Polysulfone ICLs are

less rigid than those made with PMMA, and rely more on the

high refractive index for correction of the cornea. PMMA is

an optically clear, amorphous, glassy polymer with a Tg of

117 "C and a refractive index of 1.49 (Most commercial grades

of PMMA contain a small fraction of other acrylic polymer

such as polyethylacrylate which reduces the overall Tg to

around 105'C.). PMMA ICLs provide both refractive

corrections and mechanical correction, and PMMA intracorneal

rings rely solely on mechanical changes in the corneal shape

to provide correction.

The initial designs of a refractive lens made of

polysulfone and PMMA were proved to be successful at

correcting corneal correction, but implant studies have shown

them to be unsuitable as long term implants thus far (Kerry,

1995 and Lane et al., 1989). Both polysulfone and PMMA

lenses are impermeable to water and aqueous solutions, and

the initial design of these ICLs hinders the diffusion of

water, ions, proteins and other vital nutrients to the

corneal stroma anterior to the implant. The result of these

implants was corneal opacification and necrosis in the

deprived regions, and lipid deposits posterior to the

implants (Climenhaga, 1988).

Although the lens designs are still being studied,

recent studies have focused on a lens made with hydrogel

materials (McDonald et al., 1993). Hydrogel lenses (e.g.,

Permalens*, CooperVision and Lidofilcon A, Allerghan Medical

Optics) are polymers and copolymers containing hydrogels such

as polyhydroxyethyl methacrylate (pHEMA),

polyvinylpyrrolidone (PVP), and polyacrylic acid (pAA). The

hydrogel component in these lenses makes them water

permeable. Claims have been made that since the hydrogel

lenses are water permeable, it is therefore also permeable to

glucose and other metabolites necessary to the corneal stroma

(Werblin and Patel, 1992, Werblin and Peiffer 1992, McCarey,

1990, McCarey, 1981). However, diffusion or permeability of

water does not guarantee the permeability of other molecules,

especially those significantly different in size and

functionality. Hydrogel lenses also have fairly poor

mechanical properties, and tearing or damage to the lens

during handling, implantation, or use is possible (Menapace,


The intracorneal ring (ICR) design allows passage of

nutrients and offers curvature correction by flattening the

central cornea (Kerry, 1995). The ring is surgically

implanted by sliding and rotating it into an intrastromal

channel created by a radial incision of the stroma lamella

(Kerry, 1995). The design of a PMMA ring provides mechanical

correction of the cornea without hindering the diffusion of

nutrients to the stromal region anterior to the implant.

(The absence of a refractive center in the ring requires

application of more mechanical "pressure" to achieve the same

correction of the lens designs. Calculating the required

ring shape is often difficult, and may be complicated by

unpredictable healing of the cornea following the implant

surgery (Quantock et al., 1995).

The Surgidev intracorneal lens (ICL) is a fenestrated

PMMA lens designed to correct the corneal refractive power

both mechanically (by corneal reshaping) and refractivly

(magnification by the ICL). The holes in the ICL are present

to allow diffusion of stromal nutrients to the anterior

implant region and still provide refractive correction. The

primary component of interest which is vital to stromal

tissues is glucose. Glucose is delivered to the stroma via

the aqueous humor, and its normal concentration is constant

parallel to the stroma, and decreases across the stroma from

880 to 580 pg/ml (88 to 58 mg/dl), posterior to anterior

(McCarey and Schmidt, 1990).

2.2.3 Surface Properties of Ocular Biomaterial Implants

The corneal endothelium plays a critical role in the

balance of fluids and nutrients within the cornea itself as

well as the eye posterior to the cornea. The implantation

and residence of an ICL or ICR in the cornea requires a

minimally damaging procedure because the human adult corneal

endothelium does not regenerate when damaged (Bourne and

Kaufman, 1976). The work in our laboratories has shown that

PMMA placed in contact with the corneal endothelium or

epithelium will strip away vital cells upon removal

(Yahiaoui, 1990, Katz et al., 1977, and Sheets, 1983).

Studies by Andrade (1985) and Absolom et al. (1987) show the

tendency of hydrophobic surfaces (such as PMMA) to strongly

absorb proteins, whereas hydrophilic surfaces tend to readily

desorb proteins following initial adsorption. Likewise, it

was determined by our research that PMMA modified with NVP

(Hydrograft*) provides a surface which will not readily

adhere to these cell layers on contact. The Hydrograft*

surface imparts less damage to surrounding tissues during

implantation than would an unmodified PMMA lens by providing

long term lubrication of the tissues by the hydrophilic

grafted surface. Also reported by Yahiaoui (1990) and

continually observed in current research within this

laboratory, radiation induced graft polymerization with NVP

(Hydrograft* modification) significantly reduces the

adhesion, growth, and spreading of ocular epithelial cells on

PMMA ocular implant materials.

Surface modification of the PMMA ICLs NVP is expected to

provide a hydrophilic surface within the ICL fenestrations,

and increase the permeablity of the lens to vital stromal

nutrients. This research focuses on the determining suitable

surface modification conditions for low molecular weight PMMA

ICLs and resulting changes in permeability.

2.2.4 Surface Modification Techniques

The presoak surface modification techniques discussed

previously for vascular graft modifications will be used for

PMMA ICL modifications. The monomer solutions, however, are

aqueous because of the sensitivity of the ocular environment

and the low chemical resistance of PMMA to organic solvents. Radical formation on PMMA

Figure 2.6 shows several possibilities for the
radiolysis of PMMA The abstraction of a hydrogen from the
main chain may create Radical A, which upon
disproportionation leads to the formation of Radical B and
End Group A (Todd, 1954). Main chain homolysis is also
possible, yielding Radical C and D (Kirsher et al. 1965).
Degradation may occur upon ester group cleavage, yielding
Radical E. Scission of Radical E may then lead to formation
of a new Radical D and Chain End B (Ranby and Rabeck, 1975).

CH3 CH3 C13
H--CH--- -C + C=CH-C--
Radical A Radical B Unsaturated
Endgroup A

I I a aI I I
,-C^_ C-CH,._ ^ C^^. + CH-C--
2 Radiation" 2

PMMA I Radical C Radical D

C3 C3 CH3
--C-CH-C--CH,-- P --C==CH2 + Radical D
Radical E Unsaturaded
Endgroup B

Figure 2.6 Possible gamma radiation reactions and products
for PMMA.


3.1 Materials

3.1.1 Substrates

Surface modification was carried out on several vascular

prosthesis substrate materials. These included PET (Mylar D-

1000 and 700-Dl films from DuPont Electronics), woven PET

(Dacron fabric, Meadox), reinforced velour woven PET

(reinforced Dacron vascular graft, Bard), PDMS (KE-1935 A and

B, Shincor Silicones), ePTFE (sp.# 728-3 GORE-TE* expanded

PTFE, Gore), and PTFE (skived Teflon", Goodfellow). The

chemical structures of polymer substrates used in this

research are presented in figure 3.1.

All intracorneal lens (ICL) substrates for Hydrograft*

modification were PMMA, and were provided by Surgidev as

either PMMA sheets (V-811 low molecular weight PMMA),

fenestrated flat PMMA disks, or ICLs. These ICLs have three

hole sizes and percent surface coverage of holes, 10pm/5%

(209-99), 30pm/5% (220-121), and 80pm/5% (220-131), and all

are 6 mm/-8 diopter. The PMMA used in the manufacture of the

ICLs has a significantly lower molecular weight than PMMA

typically used for ocular applications. For this reason,

different conditions than those studied by Yahiaoui (1990)

were used for modification of ICLs. Table 3.1 shows

molecular weight data for several PMMA ocular materials.

PET -<>-CHr-CH2

PMMA -(-CH =c



F F n

Figure 3.1 Chemical structures of polymer substrates
subjected to surface modification.

Table 3.1 Molecular weight data for PMMA ocular implant
materials, from Goldberg, et al. 1988-1995).

Name Lot # Mn MW MWD
Perspex CQCV 001544 3.3 X 106 5.1 X 106 1.5
Blue Perspex B# 001710 1.2 X 106 3.6 X 106 2.9
Opticlear P973 1.1 X 106 2.5 X 106 2.4
Low MW PMMA V-811 45,000 86,000 1.9
Nidek IOL PMMA 650,000 2.6 X 106 3.9
Storz IOL PMMA 60,000 110,000 1.8

3.1.2 Monomers and Reagents for Surface Modification

The monomers and polymers used for surface modification

included methyl methacrylate (MMA, Kodak), N-2-vinyl-

pyrrolidone (NVP, Polysciences), 2-sulfoethyl methacrylate

(SEMA), methacryloxypropyl terminated PDMS (MAOP-t-PDMS,

United Chemical Technology, Inc.), and PDMS (KE-1935 A and

B). MMA and NVP were supplied with MEHQ as an inhibitor, and

were purified by vacuum distillation (1-2 mm Hg at 40-60'C).

The Shin-Etsu silicone was supplied in a ready to use two

part oligomer mixture. SEMA was supplied with 5% MEHQ, and

no practical methods of its removal were found. Therefore,

SEMA was used as supplied. The MAOP-t-PDMS was supplied with

no inhibitor, and was used as received. The structure of the

monomers and oligomers used are displayed in figure 3.2.

3.2 Methods

3.2.1 Sample Preparation and Substrate Cleaning Preparation of Shin-Etsu PDMS Films

The KE-1935 PDMS was supplied by Shincor Silicones in

two separate parts. The mechanical properties of the final

polymer may be varied by altering the time and temperature of

the curing reaction. Samples used in this research were

prepared by the following standard procedure unless otherwise

noted. Twenty to twenty-two grams of each component were

poured onto a bordered glass plate (16 cm X 24 cm), mixed

thoroughly with a glass stirring rod, and the surface was
leveled with a glass microscope slide. The plate was
degassed in a vacuum oven (76 cm Hg) at 50'C for 45 minutes,
and then transferred to a preheated, 150 'C ,oven and cured
for one hour in air. The final thickness of the cured sheet
was approximately 1 mm.

//o CH2---C CH==C
C=0 =0

CH3 'CHT-CH2----OH

0 CH3 CH3 CH3
H- CHT-CH 'Ii-O -- i '--- i-CHy-CHr ,



Figure 3.2 Chemical structures of monomers and reagents
used for surface modification. Substrate cleaning prior to surface modification

All substrates were cleaned prior to use to remove

surface contamination and impurities. Samples referred to as

"controls" have undergone the cleaning process as well,

unless otherwise mentioned. Dacron", Mylar*, PTFE, ePTFE,

and silicone substrates were cleaned by sonication for 10

minutes each in acetone, isopropanol, and Ultrapure" water.

The samples were then rinsed in Ultrapure" water and placed

in a vacuum oven at 50'C to dry for 12 hours (76 cm Hg).

PMMA samples were cleaned by sonication in Ultrapure"

water for 10 minutes followed by a rinse in Ultrapure" water.

Samples provided as lenses were not cleaned prior to

modification because they were received in final manufactured

condition. Solution degassing

The samples for gamma radiation surface modification

were placed into borosilicate glass tubes with the grafting

monomer solution (sample completely submerged in the

solution). Before irradiation, the samples were "degassed"

to remove as much oxygen from the solution as possible using

one of two methods. The samples were degassed by vacuum (ca.

20-30 mm Hg) for 2 to 5 minutes, purged with argon, and

sealed with a polyethylene cap (Tainer Top, Fisher

Scientific). The second method used was to bubble argon

through the solution with agitation to replace the oxygen

with argon in solution. Bubbling was done through a glass

pasteur pipette for 2 to 5 minutes (2 minutes for 1 to 3 ml

solution volumes and 5 minutes for 5 to 10 ml solution

volumes). Bubbling was chosen over vacuum degassing for more

volatile solutions (e.g., methylmethacrylate monomer/acetone

solutions) to avoid evaporating of solution components and

changing solution concentrations. All samples were then

irradiated to the specified dose. Substrate cleaning after modification

Following surface modification with MMA, SEMA, or PDMS,

the residual polymer, monomer and solvent was removed from

the sample tube. The modified substrates (PET, PTFE, ePTFE,

and PDMS) were then placed in acetone to begin washing to

remove the remaining unbound polymer and residual monomer.

The acetone was removed and replaced three times per day, for

three days. The samples then were rinsed in Ultrapure" water

and vacuum oven dried.

PMMA samples modified with NVP were cleaned in the same

manner described, with Ultrapure" water being substituted for

acetone for the washing procedure.

3.2.2 Surface Modification Methods Presoaking

Some samples were subjected to a presoak step prior to

gamma polymerization. The presoak step involves placing the

sample to be modified in a monomer solution, usually at

higher concentration and temperature, to allow monomer and

solvent to diffuse into and swell the substrate surface. The

presoak conditions for a polymer-monomer system were usually

determined by swelling experiments. The samples were placed

in various solution concentrations in sealed borosilicate

tubes, and then placed in an isothermal water bath. The

weight increase of the sample with time was recorded, and the

solution concentration, temperature, and presoak time which

provided significant weight increases (5% or higher) were

chosen as the presoak conditions for that particular system.

Presoaking was also used for the solution dip coating of

PDMS onto Dacron*. A chloroform/PDMS solution was used swell

the PET fibers and allow diffusion of silicone oligomers into

the substrate surface.

Immediately following the presoak for surface

modification of PMMA with NVP, the samples were quenched with

ice water to reduce temperature and stop diffusion. Gamma radiation induced polymerization

The samples were immediately placed into the gamma

solution (if gamma solutions were different than presoak

solutions) following the presoak and degassed. The solution

concentrations for gamma polymerization were chosen based on

the final viscosity of the solution. That is, solutions

which could be easily removed following polymerization were

more desirable, and concentrations which caused gellation at

a given gamma dose were avoided. The samples were placed

into the gamma source immediately following degassing. Gamma

irradiation was conducted by simultaneously exposing the

substrate and monomer solution to a 600 Curie 60Co point

source. The samples were placed into a circular, motorized

carousel to provide uniform exposure to all samples.

Schematic diagrams of the gamma source chamber and carousel

are shown in figures 3.3 and 3.4, respectively.


\ 60Co

Figure 3.3 Schematic diagram of the irradiation chamber
used for polymerization.

Figure 3.4 Motorized carousel used to provide uniform
exposure within the irradiation chamber.

The total radiation dose varied from 0.02 to 0.15 Mrad,

and was controlled by time of exposure. The samples were

placed 4 inches from the point source, which provided a dose

rate of approximately 425 rads/min. The dose rate was

determined by measuring the absorbence of dosimeter film (GAF

Chromic) exposed on a calibrated irradiator at the Shands

Hospital Radiation Oncology Center, Gainesville, Florida

(Goldberg, et al., 1988-1995). Twenty-five films were

exposed from 0.025 to 0.200 Mrads and the absorbence at 540

nm was determined by UV/VIS spectroscopy (Perkin Elmer Model

Lambda b UV-VIS) at 540 nm wavelength. A calibration curve

was generated, and used for exposures on the gamma source

used in this research.

After exposure to gamma radiation, the samples were

removed from the grafting solution and systematically washed

according to the washing procedure discussed in section

3.2.3 Solution Dip Coating of PDMS onto Dacron

Dacron samples were coated with silicone as a surface

modification method. Equal amounts of a 10% of KE-1935 A and

KE-1935 B in chloroform (w/w% solution) were mixed in a

borosilicate glass tube with a screw cap (Kimax tubes, Fisher

Scientific). Dacron fabrics or Bard vascular grafts were

placed into the solution, and the caps were placed on the

tubes to avoid evaporation of the chloroform. The samples

remained in the oligomer solution for 4 hours at room

temperature to allow swelling of the PET and some diffusion

of the silicone oligomers into the substrate. Upon removal

from the solution, the samples were suspended (alligator

electrical clips) in a 60'C oven, in air The samples were

allowed to cure for 24 hours. Then vacuum was applied to the

oven (76 cm Hg), and the samples cured under vacuum at 60'C

for another 24 hours. To remove the uncured, low molecular

weight oligomers, the samples were washed in acetone, with

six solvent changes, for 48 hours, followed by washing in

hexane with three solvent changes for 24 hours. The samples

were then placed into a vacuum oven (60'C, 76 cm Hg) for 12

hours to remove the acetone and hexane.

The adhesion of the silicone layer to the substrate was

also studied (c.f. section 3.2.6). To improve the bonding of

the silicone layer to the PET substrate, a pre-dip coating

step was used. The Dacron was subjected to a presoak in MMA

solutions, followed by gamma irradiation in a chloroform

solution with 10% MAOP-t-PDMS and 10% MMA. This step is

believed to swell MMA into the Dacron. Upon exposure to

gamma radiation, the MMA and MAOP-t-PDMS polymerize into and

on the Dacron surface providing a covalent link from Dacron,

to PMMA, to MAOP-t-PDMS. The second step of solution dip

coating and curing PDMS onto this surface will then provide a

covalent bond between the Dacron and the PDMS.

3.2.4 Characterization Gravimetric analysis

Gravimetric analysis was used as a quick, non-

destructive and inexpensive method to provide information

concerning gravimetric yield for swelling presoakk) or

surface modification. All mass increases and decreases were

recorded on a Denver Instruments A-200DS electronic balance

with a precision of 0.02 mg. Percent weight increase (or

decrease) was determined by the percent change in final and

initial weights, and is defined in equation 3.1,

Percent Weight Change = {(Wf Wi)/Wi} 100 (3.1)

where Wi is the initial weight (usually referring to the

unmodified, clean, dry substrate) and Wf is the final weight

(usually referring to the modified, clean, dry substrate).

In some instances, wet weights were used as initial weights

(e.g., swelling experiments, c.f. section Contact angle goniometry

Contact angle goniometry is also non-destructive and

inexpensive, and provide information concerning the relative

wettability of the surface. This technique is dependent on

the outermost few monolayers of the polymer surface, and may

only be used on solid polymer substrates. Contact angle

values for all substrates were measured on a Rame-Hart

contact angle goniometer (Mountain Lakes, NJ) at room

temperature using the captive air bubble technique unless

otherwise mentioned. Some samples were evaluated using the

water drop in air technique.

The captive bubble technique involves suspending the

sample in water (clipped to the underside of an aluminum

block immersed in water), and injecting air bubbles (ca. 0.2

pl) with a microliter gas chromatography syringe, and

allowing them to come to rest underneath the sample. The

contact angles reported are averages of measurements on six

bubbles per sample. The contact angle measured is related to
the solid-vapor (ysv), solid-liquid (Y1), and liquid-vapor

(Y1v) interfacial free energies. The relationship between

these values is Young's equation, as shown in equation 3.2.

cos (8) = (Ysv Yl,)/y1v (3.2)

Schematics of the captive bubble technique and the angles

measured are shown in figure 3.5.

Typical values obtained with the goniometer used in this

research are 105-110', for hydrophobic substrates such as

PTFE, 50-60' for intermediate substrates such as PMMA, and

s20' for hydrophilic hydrogel substrates (Yahiaoui, 1990).


*Contact Angle

Air Bubble




Figure 3.5 Schematic representations of the captive bubble
technique and angles measured for contact angle
goniometry showing (a) sample analysis chamber,
and (b) bubble at interface. FT-IR/ATR and transmission microscopy FT-IR

Fourier-transform infra-red (FT-IR) spectroscopy

provides valuable information concerning functional group

chemistry and bonding in polymeric systems. In FT-IR, a

laser beam is passed through the sample, and the energy of

the beam is absorbed by the polymer through various phenomena

such as molecular vibrations, rotations, and stretching.

Percent transmission or absorbence (of the IR beam) is

collected for wavelengths of 2.5 to 20 pm (or wavenumbers of

4000 to 400 cm-1).

Attenuated total reflectance (FT-IR/ATR) utilizes a

crystal through which the beam is passed. There are

conditions under which the infrared radiation passing through

the crystal will be totally internally reflected. The sample

is mounted in contact with the crystal, and the evanescent

wave created within the sample is attenuated in the regions

where the sample absorbs energy. (Spectra Tech reference

Manual, Chapter 13) A schematic of the crystal, beam, sample

configuration is shown in figure 3.6. The depth of
penetration, dp, of the beam into the sample, and therefore

the depth of practical analysis is described by the

relationship shown in equation 3.3,

dp = X(3.3)
2 xni (sin8 n2/n)2 (3.3)

where X is the wavelength of the incident IR radiation, ni is

the refractive index of the ATR crystal, n2 is the refractive
index of the sample, 0 is the incident and exit angle of the

IR beam.

The major advantage to FT-IR/ATR is the ability to
provide chemical structure and bonding information on

polymeric surfaces. This allows for identification of
surface polymers present following surface modification.

FT-IR microscopy utilizes a microscope stage for

mounting the sample where the IR beam may be focused on a

specific sample surface area. The sample may be analyzed by

either transmission, where the beam is passed directly

through the sample, or by a special ATR microscope stage for

collecting ATR spectra. The main advantage to the

transmission microscopy FT-IR system is being able to

evaluate specific sample areas, however it requires thin

samples, typically less than 100 pm. Since transmission

spectroscopy is not surface sensitive as is the case with

ATR, information on the bulk composition of the sample is


FT-IR data were collected using a Nicolet 20SXB FT-IR

spectrometer using a ImW HeNe laser and a parallelogram KBr

crystal with a 60' entrance/exit face angle. Typically 128

scans at a resolution of 4 cm-1 were signal averaged to obtain

individual spectra. Data processing was done on the Nicolet

software provided with the equipment. Spectra presented here

were scaled and printed using OMNIC 1.2 software (Nicolet

Instrument Corporation).

Figure 3.6 Schematic representation of the IR beam,
crystal, and sample for FT-IR/ATR spectroscopy. X-ray photoelectron soectroscopy

X-ray photoelectron spectroscopy (XPS), also referred to

as electron spectroscopy for chemical analysis (ESCA), is a

fairly new technique for surface analysis. Developed during

the 1950s, XPS is an analytical technique which provides the

atomic chemical composition both quantitatively and

qualitatively, making it a powerful tool in the

characterization of surface modified polymers.

Based on the photoelectric effect, XPS bombards sample

surfaces with X-rays, which cause the ejection of core

electrons. The ejected electron emit discrete energy values

related to the binding energy of the electron and the

exciting radiation energy by equation 3.4 (Barr, 1994, and

Hercules and Hercules, 1976) Each atom has a specific

binding energy, and the atoms to which they are bonded cause

discrete shifts in the binding energies. This enables XPS to

measure atomic concentrations of surfaces as well as provide

information concerning the chemical environment of the

surface elements.

Eb = hu Ek (3.4)

In equation 3.4, Eb is the binding energy, hu is the photon

energy (where h is Plank's constant and v is the X-ray

frequency), Ek is the kinetic energy of the electron

(measured value), and 0 is the work function specific for the

instrument (Barr, 1990).

XPS data were collected using a Kratos model XSAM-800
spectrometer with a Mg Ka X-ray source. The X-ray gun was

operated at 12 kV and 19 mA, and the analysis chamber

pressure was maintained at 10-7 to 10-8 torr during analysis.

Quantification of the spectra was performed using DS800

Kratos software on a Digital computer system. Binding
energies were calibrated using Eb = 285.0 eV as the Cls peak

on all spectra, unless otherwise indicated. The depth of

analysis is in the range of 50A. Dynamic mechanical sampling

Dynamic mechanical sampling (DMS) measurements were made

with a Seiko DMS 200. The sample response to dynamic tensile

forces were recorded as a function of time, temperature, and

frequency (for frequencies of 0.1, 0.5, 1.0, 5.0, and 10.0

Hz) from -140 to 250'C. The sample thermocouple was

calibrated using the maximum value of the dynamic loss

modulus (E") at 1 Hz for the glass transition of PMMA

(Tg=117'C) (Feller, 1993). The elastic response was

calibrated with the tensile modulus of PMMA measured at an

elongation rate of 10%/min in uniaxial tension.

The samples analyzed had a gage length of 20 mm, and had

a range of cross-sectional areas of 4.9 to 6.4 mm2.

Information for the storage modulus (E'), loss modulus (E"),
and tan 6 (E'/E") are presented as a function of temperature

for silicone modified with PMMA, as a function of both

radiation dose and monomer concentration. Light/Optical microscopy

Sample inspection, graft thickness measurements, surface

topography, and other visible observations were made using a

Nikon optical microscope. Optical micrographs were taken

using a canon camera mounted on the same microscope.

The microscope was equipped with a graduated eyepiece

used for size measurements. A hemocytometer with precision

markings was used for calibration of the eyepiece for all

magnifications used, providing a resolution a 1 pm.

Organic and inorganic compounds were used for staining

surface modified regions of substrates to facilitate

visualization. A 10% silver nitrate solution was used for

staining the Hydrograft* materials. The modified samples

were placed in the silver nitrate solution for 12 to 24

hours, and then placed in phosphate buffered formalin (10%

formaldehyde) to precipitate silver oxide and silver

phosphate. A saturated crystal violet-acetone solution was

used to stain the PDMS coating on Dacron" and PMMA modified

PTFE. Scanning electron microscopy

High magnification evaluations of surface appearances

(before and after modification and evaluations) were made

using a Jeol 6400 scanning electron microscope (SEM).

samples were coated with a thin layer of gold palladium using

a Hummer V sputter-coater (Technics, Alexandria, VA), unless

otherwise mentioned. Typical accelerating voltages of 1 to 3

kv were used. These lower voltages were used to avoid sample

ablation typically caused by higher voltages (Goldberg et

al., 1988-1995). Micrographs were obtained at various

magnifications to provide representative records of surface


3.2.5 Diffusion/Flow Cell Testing of ICLs

The fenestrated ICL design attempts to provide a route

for the eye to supply nutrients to the stroma anterior to the

implant. A custom diffusion chamber provided by Surgidev was

used to evaluate the permeability of the fenestrated

intracorneal lenses. As displayed in figure 3.7, the

apparatus can be used to determine flow rates of liquids and

solutions through the lens, as well as evaluating the

diffusion osmoticc pressure controlled) of solution

constituents such as NaC1, glucose, and proteins. The

apparatus is constructed of PMMA with a PMMA divider in the

center. The ICL is placed in the right side of the chamber

on an o-ring, and the pressure screw cap is tightened to hold

the lens in place and seal the edges. Flow is in the right

to left direction mimicking flow through the lens from

posterior to anterior sides.

Two types of tests were conducted using this apparatus.

The first test was to determine the water flow through the

lens at a constant pressure. One chamber was filled with

water and an inverted, water-filled flask (with an air bleed

port) was placed in one chamber to provide a constant

hydrostatic pressure. The other chamber was left empty. The

volume of water flowing through the lens was measured after

90 minutes had passed. This test was conducted five times

for each sample, and an average value for each lens condition

was determined.

The next test involved analysis of the permeability of

saline and glucose solutions though the lens as driven by

osmotic pressure and chemical potential. One chamber was

filled with the solution of interest, and the other was

filled with Ultrapure water. The solution concentration of

each chamber was evaluated as a function of time.

Osmolarity was measured using an osmometer (Osmette

Micro-osmometer, Precision Systems) to determine saline

concentration. A linear calibration curve (r2 = 0.999) for

solution concentrations of 0% to 0.9% saline was created

using a series of 10 solutions. Initial saline

concentrations used in the diffusion chamber were 0.9%.

A Glucometer* glucose reader (Ames) and glucose sticks

(Ames) were used to determine glucose concentrations. A

calibration curve (r2 = 0.997) for solution concentrations of

380 mg/dl to 38 mg/dl glucose was created using a series of

10 solutions. Initial glucose concentrations used were 380


Screw Cap

Figure 3.7 Schematic illustration of diffusion and flow
cell used for evaluating permeability of ICL

3.2.6 Leak and Stability Testina of Vascular Grafts

The silicone dipped Dacron vascular grafts may be

rendered non-porous by the silicone coating on the inside and

outside of the substrate. The leak-rate of water through the

pores was evaluated using a pressurized flow system. The

graft was placed in series with a pressure manometer and

water reservoir, and a back pressure of nitrogen. A

schematic of this test set up is illustrated in figure

3.8(a). The volume of water leaking through the graft

surface at 120 mm Hg in one minute was measured. The

reported values for leaking were normalized for the surface

area of the graft, and the leak rates are reported in


The stability of the silicone coating on the Dacron

substrate was also evaluated to determine if any material

delaminated or flaked off during dynamic flowing conditions.

A continuous flow of water at 300 ml/min at 110-120 mm Hg

through a vascular graft was maintained for up to 48 hours.

Samples of the water were taken and analyzed with inductively

coupled plasma (ICP) to determine the concentration of

silicon atoms in the solution. The sample volume taken for

analysis was 10 ml, to which 0.1 ml of Triton X surfactant

was added to suspend silicone molecules which may be in the

water sample taken. Since the vascular environment contains

both lipophillic and hydrophilic molecules, silicone coated

Dacron grafts were analyzed in the flow system with both

water and a 10% Triton X water solution. This test apparatus

is shown in figure 3.8 (b). Inductively coupled plasma

ICP utilizes an argon plasma torch to ignite a fine solution

mist. The emission spectrum quantized for each element, and

when the spectra of the ionized gas is compared to reference

solutions the atomic concentration of a specific atomic

species may be determined in parts per million (ppm). The

detection limit of silicon atoms at 251.6 nm by the ICP used

in this research is 0.1 ppm, which corresponds to a

solution concentration of t 0.1 pg/ml. A Plasma 40 ICP

(Perkin-Elmer) was used with a window size of 0.1 nm, a photo

multiplier tube voltage of 700 V, and an integration of 680

msec. Each sample was analyzed 3 times, and means and

standard deviations are reported.

-- Vascular Graft

Plug to Force
Leakage Through
Graft Pores


Water Beaker
Figure 3.8 Schematic illustrations of the leakage (a) and
stability (b) analysis set up.

3.2.7 Ex Vivo and In Vive Studies of Vascular Substrates Ex vivo AV canine shunt analysis

Using aseptic technique, an arteriovenous (AV) shunt was

constructed between the carotid artery and the jugular vein

of an adult mongrel canine. Samples to be tested were placed

into a section of Silastic* tubing, and were sealed in place

with silicone RTV. Autologus 111Indium labeled platelets were

injected into the dog, and blood flow over the samples

allowed for 60 minutes. The samples were then removed from

the shunt, and counted in a gamma-counter (Auto-logic, Abbott

Laboratories). The counts are normalized to surface area of

the samples, and are reported as counts/mm2. Canine in vivo evaluation of vascular grafts

Under general anesthesia and with heparin infusion

(100U/kg), a 6 mm diameter, 10 cm long vascular graft was

placed end-to-side on the infra-renal abdominal aorta,

proximally, and on the common iliac artery distally. Two

grafts are implanted in each animal, an unmodified control

graft, and the modified graft of interest, forming a

bilateral aorto-iliac bypass on each side. The implants were

allowed to remain in place for 30 days.

Twenty-four hours prior to graft explanation, 111Indium

labeled platelets were injected. The grafts were removed and

first counted for radioactivity, and then sectioned for


histological analysis. Counts of the grafts are reported in

counts/mm2 for the total (inner and outer) luminal surface



4.1 Gamma Radiation Induced Polymerization of Methyl
Methacrylate on PET. PTFE. and PDMS

A review of the literature and current studies within

this laboratory led to the belief that PMMA surfaces on

biomaterials for the vascular environment have the potential

advantages of reducing platelet cell adhesion and reactivity,

and increasing endothelialization and healing of vascular

prostheses (Hufnagel, 1954, and Bowers et al., 1989).

Gamma radiation induced polymerization (GRIP) of methyl

methacrylate has been reported in the literature. To create

a stable surface of PMMA on a polymeric substrate with gamma

polymerization, a mixture of monomer and substrate surface

molecules must be created. The first step in achieving this

mixture for all substrates swelling the substrate with a

monomer solution, allowing the monomer to diffuse into the

substrate surface. GRIP is a heterogeneous system, and Odian

reports the rate of graft polymerization (polymerization of

monomer directly to activated substrate molecules) to be

diffusion controlled (Odian, 1981). The solvent, monomer

concentration, temperature, and time were all evaluated to

some degree to determine conditions providing surface

modification of the substrates with MMA.

4.1.1 Swelling of PET. PTFE. and PDMS in MMA Solutions

The choice of solvents for the swelling of the

substrates with MMA solutions was based on solubility

parameter, compatibility with MMA and PMMA, relative

toxicity, and ease of removal. Table 4.1 lists the

solubility parameters for studied substrates and some typical

solvents. Typically, the more similar the solubility

parameters of the solute and solvent, the better the

solubility of the solute in the solvent (Sperling, 1986).

Solvents used must be solvents for PMMA to avoid

precipitating the homopolymer and surface grafted region

during polymerization. This facilitates solvent and un-

incorporated homopolymer removal following irradiation, and

allows swelling and diffusion within the surface region to

continue during polymerization. Complete removal of toxic

solvents is important for biomaterials applications.

TABLE 4.1 Solubility parameters and H-bonding groups for
selected solvents and polymers

Solvent Solubility H-Bonding Polymer Solubility
Parameter group Parameter
Acetone 9.9 a PMMA 9.0-9.5
Chloroform 9.3 p PET 10.7
Cyclohexane 8.2 p PTFE 6.2
DMSO 12.0 m PDMS 7.3
n-Hexane 7.3
THB 9.1 m
Toluene 8.9 _
MMA 8.8 m
Solubility parameter units are (cal/ca3)1/2. H-Bonding group refers to
the strength of hydrogen bonding by the material, where a strongly,
m moderately, and p poorly bonded. (Brandrup and Immergut 1975). Effect of solvent, concentration, and temperature

Swelling of PET. In an attempt to diffuse MMA into PET, a

solvent selection was initially difficult because of the high

solubility parameter (10.7) and crystallinity of Dacronp PET

fibers. Percent weight increase, or weight uptake, was used

to evaluate swelling (c.f. section Weights were

determined after blotting the samples with Whatman #3 filter

paper. Mylar* PET films were used initially in the swelling

studies for the ease of handling, table 4.2 shows the weight

uptake of selected solvents by PET at room temperature (25 to

28'C) after 24 and 48 hours. After observing these data, the

two solvent systems chosen for PET were chloroform and DMSO.

Figure 4.1 shows a plot of weight uptake of MMA-chloroform

solutions by PET (Mylar* D-1000) with time for various

solution concentrations. Percent crystallinity changes were

not evaluated in this research, but some solvent induced

crystallization of PET by chloroform is beleived to be


Table 4.2 Percent weight increases of Mylar* D-1000 films in
selected solvents at room temperature.

Percent Weight Uptake Percent Weight Uptake
Solvent at 24 hours at 48 hours
Acetone 1% 1
Cyclohexane < 1 -
DMSO 2% 2
Ethyl Acetate < 1% < 1%
Sexane 2% 2
THF < 1% 2
Toluene < 1% < 1%
Methyl methacrylate < 1% -
Chloroform 22-25% -


S15 -

2 10


0 300 60000 0 1200 1500 1800
Swelling Time (minutes)

Figure 4.1 Percent weight increase of Mylar D-1000 PET
films with time in MMA-chloroform solutions as a
function of solution concentration.

Figure 4.2 shows the same data presented in figure 4.1
plotted as Mt/M. versus t1/2/1 based on equation 4.1.

t = ( (4.1)

In equation 4.1, Mt is the mass uptake at time, t, Mm is the

maximum mass uptake (mass uptake at infinite time), 1 is the

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