Development of an In Situ Photopolymerizable Fixative

for a Novel Keratoprosthesis

Joshua N. Markham

A thesis submitted in conformity with the requirements

for the degree of Master of Applied Science

Graduate Deparmient of Chernical Engineering and Applied Chemistry

University of Toronto

@ Copyright by Joshua N. Markham, 1999

National Library 1*1 of Canada Bibliothèque nationale du Canada

Acquisitions and Acquisitions et Bibliographie Sewices services bibliographiques 395 Wellington Street 395, rue Wellington OttawaON K 1 A O M Ottawa ON K1A ON4 Canada Canada

Your iib Voire relsrmce

Our fiie Notre rel6renc.e

The author has granted a non- exclusive licence allowing the National Library of Canada to reproduce, loan, distribute or sel1 copies of this thesis in microform, paper or elecîronic formats.

The author retains ownership of the copyright in this thesis. Neither the thesis nor substantial extracts fiom it may be printed or otherwise reproduced without the author's permission.

L'auteur a accordé une licence non exclusive permettant à la Bibliothèque nationale du Canada de reproduire, prêter, distribuer ou vendre des copies de cette thèse sous la forne de microfiche/film, de reproduction sur papier ou sur format électronique.

L'auteur conserve la propriété du droit d'auteur qui protège cette thèse. Ni la thèse ni des extraits substantiels de celle-ci ne doivent être imprimés ou autrement reproduits sans son autorisation.

Abstract

Development of an In Situ Photopolyrnenzable Fixative for a Novel Keratoprosthesis

1999

Josh MarWiam

Department of Chemicai Engineering and Applied Chemistry

University of Toronto

For patients with opaque corneas that can not be treated with corneal transplants,

keratoprostheses represent a moderately successful alternative. Current designs tend to

fail afler several months due to complications such as: poor mechanical strength,

inadequate integration of the device with surrounding tissue, and poor nutrition o f the

anterior comea due to compromised stromai permeability.

This thesis presents initial development and evaiuation of a covalently crosslinked

keratoprosthesis fixative that is injected into the stroma and polymerized in situ. A series

of PEG-CO-PVP crosslinked hydrogels, synthesized fiom solutions of varying

composition, were characterized by their swelling behaviour, permeability, and

mechanical strength. Crosslinked networks of this polymer, formed within bovine

comeas, were investigated to determine their extent of interaction with stroma1 collagen

and the mechanical properties of the resultant interpenetrating network.

Histological examination of in situ polymerized networks indicated moderate

interdigitation between polyrner and stromal collagen fibrils. Permeation triais

dernonstrated that the PEG-CO-PVP hydrogel disks had adequate permeability for stromal

implantation. Tensile testing of in situ networks showed failure at relatively low stresses.

Samples did not fail at the stroma-polymer interface, which is encouraging.

The mechanical strength of this keratoprosthesis fixative mode1 must be improved prior

to implementation. This may require further characterization of hydrogels of different

composition, or use of a bifûnctional macromer of lower molecular weight.

Acknowledgments

There are many people whom 1 must thank for their contributions toward this document,

the culmination of over two years of my life. Dr. Yu-Ling Cheng, rny supervisor,

patiently guided me through my research, providing me with a great deal of valuable

insight. Jenn Hansen, Henry Peng, Hai Lin, Charlene Ng, and Pina Turner each made my

time at school much more interesting and enjoyable, and were invariably willing to assist

me in the lab and during the final stages of my thesis. Without Chris Pereira, from the

Centre of Biomaterials, and his tensiometer expertise, my mechanical strength analyses

would not have been possible. May Griffith, fiom the Eye Institute at the University of

Ottawa, generously gave of her time and resources in order to prepare the histology slides

presented in this thesis. Dave Small at Ryding-Regency Meat Packers ensured a steady

supply of eyes. Hai Lin helped photograph these beautifid subjects, and Jem and Craig

Hansen scanned these photographs. Mike May of Rimon Therapeutics was kind enough

to allow me to use his computer facilities to print many of the figures contained within

this document. The new friends 1 made through CEGSA and at Hart House have been,

and will always be, an important part of my life. My "older" fiiends know who they are,

and displayed unending interest and enthusiasm toward my research, which buoyed me at

tirnes when my own enthusiasm began to wane. Most importantly, 1 would like to thank

my family, and especially my parents. Without your love and support, none of this would

have been possible.

1 would like to thank the Nahird Sciences and Engineering Research Council of Canada

for the financial support with which 1 was provided over the past two years.

iii

Table of Contents

1) Introduction

2) Background 2.1 ) Corneal Physiology

2.1.1 ) The Cornea 2.1.2) Functions 2.1.3) Epithelium 2.1.4) Stroma 2.1.5) Optical Clarity 2.1.6) Endothelium

2.2) Keratoprostheses 2.2.1 ) Design and Properties 2.2.2) History 2.2.3) Keratoprosthesis Implantation

2.3) In Situ Polymerization 2.3.1) Selection of Systern 2.3.2) Reaction Mechanism

3) Experimental 3.1 ) Gel Characterization - Swelling S tudy 3 -2) Permeation Shidies 3.3) Mechanical Strength Testing - Hydrogels 3 -4) Mechanical Strength Testing - Comeal Webs 3.5) Histology of Comeal Webs 3.6) Comeal Gel Characterization

3.6.1) Swelling Study 3.6.2) Hydroxyproline Assay

4) Results 4.1) Gel Characterization 4.2) Permeation Studies

4.2.1 ) Gel Characterization

4.3) Mechanical Strength of Hydrogels 4.3.1 ) Gel Characterization

4.4) Mechanical Strength of Comeal Webs 4.5) Histology of Comeal Webs 4.6) Comeal Gel C haracterization

4.6.1) Swelling Study 4.6.2) Hydroxyproline Assay

5 ) Discussion 5.1 ) Pemeability of PEG-CO-PVP Hydrogels 5.2) Tensile Strength of Fixative 5.3) Interaction Between Polymer and Collagen

6 ) Conclusions and Recommendations

7) References

Appendix Caffeine Calibration Curve Hydrox yproline Cali bration Curve Miscellaneous Solutions Winer's Algorithm (ANOVA) ANOVA Tables

List of Tables

Table 3.1 - Full-Factorial Experimental Design for Gel Synthesis and Swelling Study

Table 3.2 - Molar Ratios of Solvent to Macromer

Table 3.3 - Preparation of Corneal Webs for Histological Examination

Table 4.1 - Equilibnurn Swollen Water Content of Gels 1 - 16 (Series A, B, and C)

Table 4.2 - Yields of Gels 1 - 1 6 (Series A, B, and C)

Table 4.3 - Swollen Weight Loss for Gels 1-16 (Series A, B, and C)

Table 4.4 - Perrneabilities of PEG-CO-PVP Hydrogels

Table 4.5 - Swollen Weight Loss for Gels 9-16 (Series E)

Table 4.6 - Failure Stress and Strain of Hydrogel Strips (84-85% Water Content)

Table 4.7 - Swollen Weight Loss for Gels 9-16 (Series L)

Table 4.8 - Swelling Characteristics of Corneal Gels

Table 4.9 - Collagen Content of Comeal Gels

Table A-1 - ANOVA (HzO Content); al1 50% PEGbMA gels

Table A-2 - ANOVA (HzO Content); al1 25% PEGbMA gels

Table A-3 - ANOVA (H20 Content); a11 120s gels

Table A-4 - ANOVA (HzO Content); al1 60s gels

List of Figures

Figure 2.1 - Cross-section of the eye

Figure 2.2 - Cross-section of the cornea

Figure 2.3 - Cross-section of stroma

Figure 2.4 - Typical core and skirt keratoprosthesis

Figure 2.5 - Cordona "through-and-through" keratoprosthesis

Figure 2.6 - Chirila keratoprosthesis with core and skirt IPN

Figure 2.7 - Reaction mechanism

Figure 3.1 - Difision ce11 apparatus

Figure 3.2 - Template for tensile samples

Figure 3.3 - Instron setup for tensile testing

Figure 3.4 - Rectangular, partial-depth comeal injury

Figure 3.5 - Comeal injection

Figure 3.6 - Multiple, adjacent comeal injections

Figure 3.7 - Transfer of PEGbMA solution into reservoir

Figure 3.8 - Comeal strip for mechanical testing

Figure 4.1 - Permeant concentration in receptor half ce11

Figure 4.2 - Stress-strain cuve for hydrogel strip

Figure 4.3 - Failure point of corneal web samples

Figure 4.4 - Histology of comeal webs

Figure 4.5 - Comeal injections

Figure 6.1 - Template for reproducible corneal web and hydrogel tab synthesis

Figure A-1 - Caffeine Calibration Curve

Figure A-2 - Hydroxyproline Calibration Curve

vii

1) Introduction

The comea is the clear tissue at the front of the globe of the eye, and serves as the eye's

main source of refiactive power, as well as its protective barrier. The comea's stroma1

layer comprises 90% of the comeal thickness, and contains primarily water (78 wt%) and

collagen (15 wt%)'. The collagen fibrils within this layer are arranged into approximately

200 sheets, or lamellae. Each lamella traverses the entire comea, parallel to the comeal

surface. The collagen fibrils within a single larnella run parallel with each other, have a

diameter of 20 to 30 nm, and are separated by about the same distance'. The collagen

fibrils of adjacent lamellae lie approximately perpendicular to each other. This regular

spacing of collagen is the pnmary explanation of the comea's transparency2. Under

normal conditions, the comea will maintain its transparency indefinitely.

Injury or any of a wide variety of diseases can disrupt this regular orientation of collagen,

ieading to comeal opacity and blindness. Penetrating keratoplasty is usually an effective

treatment for such disorders; transplantation of the comea is more successful than that of

virtually any other organ. There are certain situations, however, when comeal transplants

are not successful (e.g., severe chernical burns) or when donor organs are not available.

In such an event, prosthokeratoplasty becomes the primary option. It is generally agreed

that the charac teristics of an ideal kerato prosthesi s include3:

i) biocompatibility with the host comea;

ii) the ability to prevent infection andor epithelial downgrowth by promoting growth of

the epithelium over the anterior implant surface;

iii) high permeability to corneal nutrients to prevent necrosis of anterior comeal tissue;

iv) the ability to heal and integrate with the sunounding comeal tissue, thereby

withstanding intraocular pressure and preventing infection andor epithelial downgrowth;

and

V) the ability to avoid the formation of a retroprosthetic membrane.

The history of keratoprostheses dates back over 200 years, but these implants are still

only rnoderately successfül at best. Most keratoprostheses remain intact without

disastrous clinical reaction for a pied of only a few months. This relatively

unsuccessful procedure is performed only because of the desperation of patients who have

lost their eyesight. Comrnon complications that lead to implant failure include: poor

integration of the keratoprosthesis with the surrounding stroma1 tissue; epithelial

downgrowth; and diffùsional limitations induced by the implant, leading to anterior tissue

necrosis and implant extnisiodA.

Several curent core-and-skirt keratoprosthesis modelsS incorporate the application of

"porous" skirts to overcome some of these difficulties. in these models, the porous skirt

is situated around the clear optic core; the two parts are held together, either by covalent

bonds or by physical means. The skirts' porous nature allows infiltration and

proliferation of stromal cells, which synthesize and lay down collagen, leading to

improved implant integration. However, these polymer networks, which are

morphologically heterogeneous, often have poor mechanical strength6*'.

in this thesis, a keratoprosthesis mode1 is proposed that uses a novel in situ polymerized

network as its method of fixation. in the ideal scenario envisioned, a trephine will be

used to create a full-thickness injury in the centre of the cornea through which a clear

optic core will be inserted. A macromer solution containing a photoinitiator will be

injected at several locations within the comeal stroma, around the penphe~y of the optic.

The macromers will polymerize upon exposure to ultraviolet light, forming a crosslinked

network. This network will interact intimately with the collagen lamellae of the stroma,

producing an interpenetrating network (IPN). %y grafting macromer to the stroma-

contacting portion of the optic's surface, it will be covalently attached to the crosslinked

network forrned.

It is expected that this method of fixation will:

i) reduce postoperative healing time by eliminating the need for the distinct skirt

implantation procedure that is typically required for most keratoprostheses in clinical use;

ii) produce an inunediate, intimate interpenetration between the stromal collagen and the

in situ polymerized network; and

iii) be mechanically stronger than current porous skirt models.

An in situ pdymerization system was selected based on the in situ hydrogel coating of

Hubbell el The macromer chosen was a 3000-MW bifunctional derivative of PEG,

an exceptionally biocompatible materid4. The goal of this project is to develop and

evaluate the applicability of such a system. The work presented in this thesis includes:

synthesis of hydrogels based on the bifunctional PEG derivative;

analysis of the effect of various reaction parameters on hydrogel swelling behaviour;

evaluation of the permeability of the hydrogels, relative to comeal requirements;

determination of the hydrogels' mechanical properties; and

mechanical and histological studies of the polymenzation system's behaviour in situ.

2) Background

2.1) Corneal Physiology

2.1.1 ) The Comea

The comea is a clear, colourless, concavo-convex disk of highly-organized collagenous

tissue located at the front of the eye (see Figure 2.1). The comea comprises one sixth of

the eye's outer shell", and is surrounded by the white, opaque tissue that comprises the

sclera. The interfacial region between the comea and the sclera, known as the limbus,

takes on some of the characteristics of each.

Figure 2.1 - Cross-section of the eye

The comea appears to be slightly elliptical fiom a fiontal view, slightly longer

horizontally than transversely, but it is circular when viewed from behind. This

discrepancy is due to the asymmetrical overlap between the sclera and the comea The

typical diameter of the mature human cornea is between 1 1 and 12 mm, and its thickness

ranges fiom 0.52 mm at the centre to 0.65 mm at the comeal periphery'*'s. The adult

human comea has a radius of curvature of 7.86 mm'.

2.1.2) Functions

The comea's functions are both protective and optical in nature; its optical functions

involve both transmission and refiaction of light. The comea provides about 70% of the

refractive power of the eye3, with the lens and other ocular components in the visual

pathway making up the remainder. This high refractive ability of the cornea is pnmarily

due to the large difference in refiactive index at the interface between air and the oily t e s

film that coats the anterior surface of the cornea.

The cornea is able to perfon these functions due to the unique structures of each of its

most significant sections: the epithelium, the stroma, and the endotheliurn (Figure 2.2).

Epithelium - 4-6 cells Bowrnan's Membrane - unorganized collagen Stroma - collagen sheets or lamellae

with 2-3 vol Oh cells (keratocytes) Descemet's Membrane - basement membrane Endothelium - single celt layer

Figure 2.2 - Cross-section of the cornea

2.1.3) Epithelium

The epitheliurn consists of four to six ce11 layers, and represents about 10% of the entire

corneal thickness. At the anterior epithelial surface is a single layer of flat, thin squamous

cells, coated by the aforementioned tear film. These cells are regularly sloughed off by

the eyelid during blinking. Posterior to this are two to four layers of irregularly-shaped

wing cells, and a single layer of columnar basal cells. The basal cells undergo mitosis

fiequently to replace wing cells that have migrated antenorly to replace lost squamous

cells.

One major fùnction of the epithelium is to provide a smooth refractive surface. This is

achieved by the tear film, and by the constant polishing of this surface by the eyelids.

The epithelium and tear film also minimize comed dehydration while allowing some

pemeation of nutrients such as glucose and dissolved oxygen.

The epithelium is separated fiom the stroma by Bowman's layer. This is a clear. uniform.

8-12 pm thick membrane that is secreted by epithelial cells, and consists of minute

collagenous fibnls in a homogeneous, glycoprotein-rich matrix.

2.1.4) Stroma

The stroma is the most substantial portion of the comea, cornprising about 90% of the

comeal thickness. It is composed of 78% water, 15% collagen, 5% other proteins, 1%

proteoglycans, and 1% dissolved salts'.

Both the strength and the unique clarity of the stroma are due to the regular orientation of

collagen fibres. The collagen fibres, primarily Type 1, are grouped into bundles called

fibrils, which have a diameter of between 24 and 28 n ~ n ' ~ . These fibrils are arranged

parallel with each other, and are spaced approximately 20 to 30 nm apartl, forming

regularly ordered sheets or lamellae. There are over 200 lamellae in the stroma2, each

between one and two micrometres thick. Each larnella traverses the entire comea,

parallel to the comeal surface. Adjacent lamellae lie parallel to each other, but the fibrils

within these adjacent sheets are onented at right angles with each other (Figure 2.3).

Additional strength is provided by the interchange of fibrils between adjacent lamellae.

lamellae

Figure 2.3 - Cross-section of stroma, showing

orientation of collagen fibrils within adjacent lamellae

The stroma also contains a number of cells, mostly keratocytes, distributed within and

amongst the lamellae. The keratocytes are flat cells that are c o ~ e c t e d to each other via

very long, thin, protoplasmic bridges, forming a communication nehvork. These cells are

aiways moderately metabolically active, but activity increases greatly during wound

healing. The activated keratocytes, or fibroblasts, synthesize and secrete collagen fibre

precursors and ground substance. Keratocytes also act as phagocytes, breaking down

collagen fragments during wound heaiing. Mitosis of keratocytes ends early in life,

resuming only upon death of cells due to injury.

The remainder of the stroma1 volume is filled with ground substance, an aqueous solution

consisting mainly of dissolved proteins, glycoproteins, and mucopolysaccharides (mostly

keratan sulfate and chondroitin sulfate)15. Most of the water in the stroma is likely held

in fixed positions, bound to the long polymer chahs that make up the collagen and

mucopolysaccharidesl. It is through the ground substance that difhsion of inflammatory

cells, nutrients, and waste products occur.

2.1.5) Optical Clarity

Several different theones have been developed to explain the clarity of' the cornea2.

Earlier researchers theorized that optical clarity was due to the homogeneity of comeal

components, leading to a uniform refiactive index within the cornea. Later, the lattice

theory proposed that the strict spacing of collagen fibres produced destructive interference

of scattered light. The current belief is that scattering does not occur for variations in

refkactive index over distances less than half the wavelength of light. In other words,

"light cannot resolve structures substantially smaller than the dimension of its

wave~en~th'". The distance between comeal fibres (up to 30 nrnl) fits this description.

Any disruption to the ordering of collagen fibres in the stroma, due to disease, injury, or

changes in hydration, will compromise comeal clarity.

2.1.6) Endothelim

The comeal endothelium represents the postenor surface of the comea, consisting of a

single layer of flattened epithelial cells. These cells are about 5 pm thick and are

generally hexagonal in shape. The endothelium is separated from the stroma by

Descemet's membrane, a 5 to IO-pm thick elastic membrane that is secreted by the

endothelial cells and is composed mostly of collagen and carbohydratesl'. Although

comeal endothelial cells are extremely active metabolically, they do not undergo mitosis

in adult humans; when cells of this layer are lost to injury, neighbouring cells spread and

migrate to cover the injured area'.

The main fiinction of the endotheliurn is as a banier that controls the level of corneal

hydration. The endothelial cells' membranes intemnrine to fonn intercellular bridgesi5,

maintaining the integrity of the single-ce11 layer and completely separating the anterior

chamber from the comea. Using an ion pump, the comea is naturally maintained in an

unswollen state, a condition necessary for transparency. Nutrients and metabolic wastes

also diffise across this thin layer.

2.2) Keratoprostheses

2.2.1) Design and Properties

The typical modern keratoprosthesis (Fig. 2.4) consists of two main parts: a clear optic

core and a thin, annular "skirt". The optic core is inserted in a hole cut through the centre

of the cornea. The skirt is embedded between stroma1 larnellae and is used to hold the

optic in place. Both portions of the device are generally made with synthetic polymers.

In most cases, the two segments are made fiom different materials.

Figure 2.4 - Typical core and skir t keratoprosthesis

There is general agreement regarding the desirable properties of an ideal keratoprosthesis.

Among these requirements are':

i) biocompatibility with the host comea;

ii) the ability to prevent infection a d o r epithelial downgrowth by promoting growth of

the epithelium over the implant's anterior surface;

iii) sufficiently high pemea5i:ity to comeal nutrients to prevent necrosis of anterior

comeal tissue;

iv) the ability to heal with and form a tight connection with the surrounding corneal

tissue to withstand intraocular pressure and to prevent infection ancilor epitheliai

downgrowth; and

v) the ability to avoid the formation of a retroprosthetic membrane.

2.2.2) History

Keratoprostheses, or artificial comeas, have been in existence for over 100 years. The

17 keratoprosthesis was first proposed by de Quengsy in 1789 , but the first reported

clinical trial did not occur until Nussbaum's glass crystal model was inserted into a

18 comea in 1855 . This first keratoprosthesis remained in place for 7 months. Further

efforts with this model showed that the major complications leading to failure of the

device were: filtration of the aqueous humour, formation of a retroprosthetic membrane.

19 aseptic necrosis, and extrusion of the device .

Further atternpts were made throughout the rernainder of the 19th century, using such

materials as glass, quartz, and celluloid. Begiming late in the 19th century, work focused

mostly on keratoplasty - corneal transplants. Early in the 20th century, however, it

becarne evident that not al1 patients could be treated in this manner. Thus, interest in the

keratoprosthesis was rekindled.

The "modem" period in keratoprosthesis research began with the use of polymers as

keratoprosthetic materials following World War II. During the war, the canopies of many

military aircraft collapsed and shattered, with the fiagments becoming embedded in the

pilots' eyes. These fiagrnents, made of poly(methy1 methacrylate) (PMMA), were very

20 well tolerated in the cornea, with virtually no immune response . Shortly thereafter, the

feasibility of using this "biomaterial" in a keratoprosthesis was being investigated.

In the 1950's, researchea designed a wide array of keratoprosthetic models, varying both

materials and geometries. It was during this period that the "core and skirt" rnodels

gained popularity. Among the first of these to be developed was the Cordona "through-

4 and-through" keratoprosthesis (Fig. 2.5)- in which the eyelid was sewn shut over the

prosthesis for added strength.

Figure 2.5 - Cordona b6through-and-through" keratoprosthesis

2 1 22 Work continued with PMMA, but silicone-based materials and ceramics were also

exarnined. Perforated devices with smaller diameters were introduced in order to

overcome the diffusion limitations and subsequent malnutrition of the anterior comea

caused by implantation of large, impermeable disks. Still, tissue necrosis and implant

extrusion were very common.

In the 1960's and 1970ts, research centred on clinical trials. Over the course of countless

studies, keratoprostheses evolved into thinner implants with even larger perforated areas.

These characteristics M e r improved nutrition of the anterior comea, and also reduced

19 the separation of the stromal lamellae between which the skirt was inserted .

Autologous tissues such as teeth, bone, cartilage, and fingemails were used, but fared

pooriy in most cases.

Another significant advancement during this penod was the development of surgical

techniques and speciaiized instruments. A two-step implantation technique was

designedZ3 to minimize the damage inflicted on the corna during surgery, thereby

facilitating post-operative healing, and improving implantation success rates.

A recent development in this field is the use of a wider array of synthetic polymers, such

3 1 (PTFEY-~O, polyurethanes, and a polybutylene/polyethylene blend . Most of these new

5 designs utilize a porous skirt . This porosity serves to enhance the proliferation of

stromal keratocytes into the skirt region. Cellular infiltration and subsequent collagen

production within the pores help to integrate the porous skirt with the surrounding

32-35 stroma .

2.2.3) Keratoprosthesis Implantation

The modem, core-and-skirt keratoprosthesis is typically implanted in two distinct steps.

In the first step, a small, partial-depth incision is made in the cornea. A specialized

spatula is inserted into the wound and is oscillaied between lamellae to create a stroma1

pocket. The annular skirt is inserted into this pocket, which is then stitched shut. Several

weeks later, after the initial injuries have healed, a trephine is used to create a full-

thickness injury through both the centre of the comea and the annulus in the skirt. The

optic core is inserted into this hole and is attached to the skirt. A conjunctival flap is

usually sewn over the implant for some temporary additionai strength.

Means of fixation between the stroma and skirt, and between the skirt and core, Vary

greatly. The heaiing period following insertion of a perforated skirt is required in order to

allow the collagen lamellae to grow up to and within the perforations. Porous skirts,

which are more easily invaded by c e l l ~ ~ ~ ~ ~ , are now being studied extensively. The

invading cells lay down collagen within the pores, enhancing integration between the

tissue and the implant.

Fixation between the skirt and the core may be classified into chernical and physicai

means. Physical methods include complementary threads on the outer surface of the optic

core and the inner rim of the skirt (see Fig. 2.5). Chemicai fixation techniques are

generally executed a priori, such as the formation of an interpenetrating network (IPN)

between the core and skirt as done by Chirila et ai3!

In Chirila's (Figure 2.6), a clear, homogeneous core and an opaque, porous,

"spongy" rirn are produced with poly(2-hydroxyethyl methacrylate) (PHEMA) in a two-

step process. In the fint step, a PHEMA solution of high water content is polymerized,

forming a macroporous ring. In the second step, a second PHEMA solution of lower

water content fills a cylinder in the rniddle of the ring. Before the second polymerization

begins, this solution penetrates into the porous skirt. Thus, the clear optic that is

produced is integrated with the surrounding spongy rim, forming a homo-IPN (since both

parts are made fiom the same material).

Figure 2.6 - Chirila keratoprosthesis with core and skirt IPN

2.3) In situ Polymerization

2.3.1) Selection of System

In situ polyrnerization is used for several different biomedical applications: in dental

3 7 3 8 39 fillings , in production of bioartificial polymeric materials , and in drug delivery . Of

panicular interest is the in situ polymerization method used by Hubbell et al&13. Using an

engineered bioerodible macrorner based on poly(ethy1ene glycol) (PEG), they were able

to form coatings on tissues, with the goal of preventing postoperative complications such

as intima1 thickening and aàhesion formation. While this macromer's bioerodible

properties are not particularly suitable for the present application, it was felt that a similar

system using a bifunctional PEG molecuie could be used.

PEG itself is widely used in biomaterial applications due io its proven biocornpatibility - it displays low imrnunogenicity and low toxicity, and is usually not hannhl to living

14 cells . One exception to these observations is low molecular weight (less than 400)

PEGs, which may exhibit some toxicity.

2.3.2) Reaction Mechanism

A 3000-MW PEG denvative, PEG-bis-methacrylate, or PEGbMA (Figure 2.7b), was

chosen as the bifunctional macromer that will form the backbone of the final, crosslinked

structure. 2,2-dirnethoxy-2-phenylacetophenone, or DMPA (Figure 2.7a) was selected as

the photoinitiator for the system. Since DMPA is relatively insoluble in aqueous

environments, N-vinyl pyrrolidone, or N-VP (Figure 2.7c), was selected as an organic

solvent for the photoinitiator.

The reaction begins upon exposure of DMPA to ultraviolet light of 350 nrn wavelength,

forming two radicals for each DMPA molecule (Fig. 2.7a). These radicals initiate the

reaction by reacting with either PEGbMA (Fig. 2.7b) or N-VP (Fig. 2.7c), forming the

start of a growing chain.

CH, DMPA

V PEGbMA (n=64 for MW 3000)

N-VP 0

Figure 2.7 - Reaction mechanism: a) radical formation;

b) initiation with PEGbMA monomer; c) initiation with N-VP monomer

Propagation of the reaction proceeds as growing chains react with monomers of either

PEGbMA (Fig. 2.7d) or N-VP (Fig. 2.7e). As the reaction continues, eventually a

crosslinked network will be formed (Fig. 2.7f). The junctions (e) in this network are

formed when each of the two reactive methacrylate groups in a PEGbMA molecule react

with different growing chains. Since N-VP has oniy one reactive vinyl group, it is

incapable of forming crosslinks. Instead, N-VP monomers are found in the linear portion

of the network, between crosslinks. The linear segments of the network may ais0 contain

PEGbMA molecules having one unreacted methacrylate group.

Incorporating N-VP into the crosslinked network may have an impact on the resultant

hydrogel's properties. PEGbMA is ordinarily very flexible, but introducing N-VP

molecules into the network may compromise this flexibility, producing a more rigid gel

with greater mechanical strength and lower swollen water content. Biocompatibility of

this material should not be affected, however, as PVP is known to have good

biocompatibility, and is thus widely used in biomaterial applications.40

Typically, termination of a polyrnerization reaction occurs either when al1 reactants have

been consumed or when two reactive groups, eacb with one fiee electron, participate in

either a combination or disproportionation reaction. In a crosslinking reaction such as

this, however, the propagation and termination reactions "autodecelerate" as the mobility

4 1 of macromea decreases in the increasingly viscous solution . In other words, the

unreacted macromer andior fiee chah ends can no longer reach potentially reactive sites

due to the reduced mobility of the growing crosslinked network. This phenornenon leads

to incomplete conversion, and the presence of unreacted monomen, that will eventually

be washed out of the resultant hydrogel. These monomen may pose a biocompatibility

concem, particularly for N-VP, which is a known irritant.

R * + IA - Y O r l

CH=CHz C O

R-CH-CH:

Figure 2.7 (cont'd) - Reaction mechanism:

d) propagation with PEGbMA monomer; e) propagation with N-VP monomer;

f ) crosslinked rietwork

3.1) Gel Characterization - Swelling Study

Poly(ethy1ene glycol)-bis-methacrylate (PEGbMA) was supplied by Sheanvater Polymers

(Huntsville, AB, USA) at a purity of 97% and a molecular weight of approximately 3000.

2,2-dimethoxy-2-phenylacetophenone (DMPA) and N-vinylpyrrolidone (N-VP) (Aldrich

Chernical Company Inc., Milwaukee, WI, USA) were used as photoinitiator and solvent

for the photoinitiator, respectively.

Four factors were selected for this study; each factor was studied at two levels:

i) the concentration of macromer (PEGbMA) in solution (50% w/v and 25% wh);

ii) the duration of exposure to ultraviolet light (1 20s and 60s);

iii) the concentration of photoinitiator (DMPA) in solution (1 800 ppm and 900 ppm); and

iv) the concentration of photoinitiator solvent (N-VP) in solution (10% v/v and 5% v/v).

Ushg a full-factorial design, 16 different formulations resulted, as s h o w in Table 3.1

below.

Initially, three hydrogeis of each formulation were produced, for a total of 48 gels. These

were produced with the use of tissue culture well plates (Costa Corporation, Cambridge,

Massachusetts), which were used as molds. Each 2.6-cm diameter well was filled with

one millilitre of well-mixed solution. When dl wells in the tissue culture plate were

filled, the plate was placed in a photochernical reactor (Rayonet Photochernical Mini-

Reactor, mode1 RMR-600, with 350-rn bulbs, The Southem New England Ultraviolet

Company, Branford, CT, USA) and photoinitiated for the required period of time.

Table 3.1 - Full-Factorial Emerimental Design for Gel Svnthesis and Swelline Study

1 Formulation 1 [PEGbMA] 1 UVCuring 1 [DMPA] 1 IN-VPI

Following this initial curing time, the well plate was removed from the reactor and lefl to

Number 1 2 3 4 5 6 7 8 9 10 11

1

react for 15 to 20 hours. Well plates were covered during this time to minimize

dehydration, thereby allowing maximum mobility of reactants during polyrnerization.

(% wlv) 50 50 50 50 50 50 50 50 25 25

Time (s) 120 120 120 120 60 60 60 60 120 120

25 1 120

(ppm) 1800 1800 900 900 1800 1800 900 900 1800 1800

(% vlv) 1 O 5 10 5 10 5 10 5 10 5

900 10

Following the 15- to 20-hour polymerization, the hydrogels were removed from their

molds and placed in separate containers of distilled water. Water was changed frequently

during the initial washing phase.

The disk-shaped gels were subjected to gravimetric andysis periodically during this

washing and swelling phase. Upon removal from the water, the disks were quickly

blotted dry before being weighed. Swelling was allowed to continue until the weight of

the gels changed by less than 1% over a 48-hour period (or by less than 0.5% over a 24-

hour period). The weights recorded at this point were designated as the "swollen weight"

of each gel.

Once fully swollen, the hydrogel disks were removed fiom the water and left out to dry.

Again, the gels were analyzed gravimetrically until their weight changed by less than 1%

over a 48-hour period, at which point each gel's "dry weight" was recorded.

In order to ver@ the first equilibrium swollen weights, each hydrogel was again swollen

in distilled water in a similar fashion to that described above. When the disks' weights

changed by less than 1% over a 48-hour period, the "reswollen weight" of each gel was

recorded.

The equilibrium water content (W) of each gel, expressed as a weight percentage, was

calculated by using equation 3.1 :

where: ws, = swollen weight (g); and

w, = dry weight (g).

The yield (Y) of each gel was determined using equation 3.2:

where: w, = weight of PEGbMA in original rnacromer solution (g);

3 Y, = volume of N-vinylpyrrolidone in macromer solution (cm ); and

3 p, = density of N-vinylpyrrolidone ( 1 .O40 g/cm ).

in order to gauge the swelling repeatability, the first and second swollen weights were

compared, and the diflerence (D) between these was calculated by using equation 3.3 :

(3-3);

where: w, = reswollen weight (g).

Table 3.2 has been included to demonstrate that, since the molar concentration of N-VP is

at least three times greater than that of PEGbMA, a significant amount of N-VP will be

incorporated into the hydrogels (provided that the reactivity ratio of PEGbMA vs. N-VP

is close to unity). The gel numbers in this table refer to the solution formulations outlined

in Table 3.1.

Table 3.2 - Molar Ratios of Soivent to Macromer

As the molar ratio of N-VP to PEGbMA increases, a greater proportion of

monofunctional monomers will be present. Thus, the hydrogels produced will have a

lower crosslinking density. As a hydrogel's crosslinking density decreases, the elastic

forces within this hydrogel decrease, and swelling increases. Another consequence of

less fiequent crosslinks is a lower mechanical strength4'. Conversely, as this molar ratio

decreases, the hydrogels will have a greater crossliliking density, a lower water content

when swollen to thermodynamic equilibrium, and greater tensile strength. Thus, by

controlling the solution composition, one c m predictably influence hydrogel parameters

such as swollen water content and tensile strength.

Gel No. 1 or5 2o r6

lo4 x Moles PEGbMA

1.64 1.64

lo4 x Moles N-VP 9.36 4.68

Moles N-VPI Moles PEGbMA

5.69 2.85

Mass N-VP/ Mass PEGbMA

0.2 1 0.10

3.2) Permeation Studies

Two sets of hydrogels were produced as described in section 3.1. Gels were placed in

containers of distilled water following an 18- to 19-hour reaction period, and were

washed and swollen for over two weeks prior to being used in permeation trials. As the

gels made with 50% PEGbMA solutions had already been screened out of the study (see

Section 4.1), only 16 gels were synthesized, two for each of formulations 9 ihrough 16

(see Table 3.1). Of each of these pairs of gels, one was used for permeation trials, while a

small section of the other was used to detemine the hydrogels' equilibrium swollen water

content.

Shown in Figure 3.1 is the experimental setup used for the permeation trials. A hydrogel

disk was removed from its water bath and patted dry, and its thickness was measured

prior to being clarnped between the two halves of a diffision cell. The inner diarneter of

the half cells was 1.6 cm. It was observed that this clamping provided an adequate seal to

avoid lealcage, and did not darnage the thick, robust hydrogel. The outer rim of the gel

QD disk, exposed between the half cells, was wrapped in Parafilm to prevent dehydration.

The diffision ce11 was positioned on a magnetic stirrer, and a stir bar was deposited in

each half cell. nie receptor half ce11 was filled with 15 mL of distilled water. The donor

half ce11 was filled with 15 mL of a 10 000-ppm solution of caffeine (Aldrich Chernical

Company Inc., Milwaukee, WI, USA). The top of each half ce11 was covered with

Q Parafillm to prevent dehydration during the trial.

Each tnal was one hour long, with samples taken at 5, 10, 15, 20, 30, 40, 50, and 60

minutes. At each of these times, a 3-mL sample was removed fiom the receptor half cell,

O and was replaced by 3 mL of distilled water. The Parafilm covering was replaced

following sampling .

stir

receptor - cell

disk

Figure 3.1 - Diffusion cell apparatus

(Cd = [permeant] in donor half cell; Cr = [permeantl in receptor half cell)

The permeation samples were analyzed on a WNis-spectrophotometer (Hewlett Packard

mode1 8452A Diode Array Spectrophotometer). A calibration curve (Figure A-1) was

generated for caffeine, over the range of 0.1 to 50 ppm, and used in the caffeine analysis.

To correct for the 3 mL sample size, the permeant concentrations in the receptor half ce11

were adjusted by using equations 3.4 through 3.6:

where: C, = permeant concentration in receptor ceIl at first sampling point (ppm);

C, = permeant concentration in receptor cell at second sampling point (ppm);

3 VI = sample volume (3 cm );

V = volume of fluid in each half ce11 (15 c d ) ; and

C, = permeant concentration in receptor ce11 at nth sampling point (ppm).

For each trial, a plot of Cr vs. t was generated, using these adjusted concentrations. The

first several data points for each trial represented the initial lag phase (prior to steady

state), and were discarded. A regression analysis of the remaining points gave the dope

(dC, /dr), where t is the elapsed time at each sampling point.

Assuming a linear concentration profile across the hydrogel disk, the flux across this

membrane may be expressed as:

where: J, = flux across hydrogel disk (g/cm2/s);

P = pemeability of disk to permeant (cm2/s); and

1 = thickness of hydrogel disk (cm).

Also, the rate of permeant diffusion across the membrane may be expressed as:

where: N, = rate of permeant difision across hydrogel disk (g/s); and

S = exposed cross-sectional area of hydrogel disk (cm2).

By equating the expressions for J. shown in equations 3.7 and 3.8, and setting Cr + 0, the

permeability P of each gel was found.

3.3) Mechanical Strength Testing - Hydrogels

Hydrogels were synthesized in 4.9-cm diameter pyrex" dishes, in a similar fashion to that

outlined in section 3.1. The previous gels that were produced in tissue culture wells were

very non-unifonn in thickness. The macromer solution adhered to the edges of the wells

to such an extent that the resultant hydrogel disks were much thicker around the

perimeter. For tensile testing purposes, sarnples of uniform thickness were required. The

larger pyrex@ dishes were used to provide for this, with a larger, central uniform region

away fiom the edges of the mold.

Again, gels were synthesized only with 25% w/v PEGbMA solutions (formulations

9 to 16). Only one set of eight gels was produced. AAer a polymerization period of 10 to

26 houn, the gels were removed from their molds and placed in distilled water for several

weeks. Sections were cut out fiom the perimeter of each of these gels and were analyzed

gravimetrically to determine their swollen water content.

Hydrogel strips were cut out of the central portion of each gel for mechanical strength

testing. A stainless steel template, the dimensions of which are s h o w in Figure 3.2, was

used to cut samples of appropriate size and shape. The dumbbell shape was selected in

order to force the samples to break within the narrower middle section. A broad, type-21

scalpel blade was used for cutting the relatively straight sections, while a narrower, type-

1 1 blade was used for cutting around the corners and curved sections. Once the strips had

been extracted, they were placed back in distilled water. -

Figure 3.2 - Template for teasile samples

The hydrogel strips were tested on an instron mode1 8501 tensorneter (Instron Canada,

Laval, PQ). All samples were stretched until failure at a constant rate of 1 m d m i n , with

a data acquisition frequency of 60 Hz. For the first two samples tested (gels 14L and

15L), a 25-lb load ce11 was used. Since it became clear that a smaller load would be

adequate, a 500-g load ce11 was used for the remaining 6 sarnples.

A diagram of the apparatus used is s h o w in Figure 3.3, showing the stationary top grip

and the vertically mobile bottom grip. It was found that as the grips were tightened, the

gripped portion of the gel became compressed. This, in effect, partially squeezed the

hydrogel out of the grips, resulting in a compressive force reading by the tensometer's

load cell.

hydrogei _, strip

lower grip

and down

Figure 3.3 - lnstron setup for tensile testing

A sample-loading method was developed to offset this compressive force. One end of the

hydrogel sample was loaded into the top grip, tightening the screws only enough to hold

the sarnple gently. The bottom grip was raised into position, and was tightened ont0 the

other end of the sarnple until a compressive force of 10 g was recorded by the load cell.

The bottom grip was lowered slowly, extending the sample until the compressive force

was reduced to zero. The top grip was similarly tightened to a compressive force of 10 g,

followed by another downward adjustment of the bonom grip until a reading of zero was

obtained.

3.4) Mechanical Strength Testing - Corneal Webs

Six fresh bovine eyes (Ryding-Regency Meat Packers, Toronto, ON) were placed in a

physiological salt solution (PSS; see page 80 for composition). Each eye was removed

from PSS just pnor to handling. A type-21 scalpel blade was used to make three

partial-depth incisions in one half of each comea, forming three sides of a rectangle.

With a pair of surgical scisson, this rectangular section of comea was removed (Figure

3.4) by cutting between stroma1 layers at the depth of the scalpel incisions. The surface

of fie cornea was patted dry, and a layer of silicone grease was applied around the

perimeter of the rectangular injury.

A PEGbMA solution was prepared, according to formulation #9 (see Table 3.1). This

was loaded into a disposable syringe, to which a 25-gauge needle was attached. The

needle was inserted into the comeal injury, parailel to the surface of the comea

(Figure 3.5). A small amount (less than 100 4) of the solution was manually injected

into the stroma, forming a raindrop-shaped opaque region. The needle was withdrawn,

then inserted in several more locations dong the edge of the corneal injury, forming a

nearly contiguous row of injections (Figure 3.6).

The eye was transferred into a shallow glas jar, which was filled with PSS such that the

entire eye, except the comea, was bathed in the saline solution. More PEGbMA solution

was transferred into the comeal injury itself, filling the reservoir that had been forrned

with the silicone grease (Figure 3.7).

Three of the six eyes were covered to block out al1 ambient light, and were lefi to sit for

approximately one hour before photoinitiation, allowing the injected material to diffuse

through the stroma. The other three eyes were transferred, immediately after the reservoir

was filled with macromer solution, into the photochernical reactor. Here, the contents of

the jar were exposed to 350-m light for two minutes.

After allowing four hours for the polymerization reaction to occur, the inj ected pol ymer

had forrned a comeal "web", interacting with the collagen in the surrounding stroma. The

solution within the reservoir had formed a "tab" of crosslinked macromer (Figure 3.6).

The eyes were entirely submerged in PSS and were placed in a refrigerator overnight.

The following day, the comeas of al1 six eyes were excised by cutting around the limbal

region with surgical scissors. A comeal strip was excised such that its midpoint was near

the interface between the comeal web and the PEG-CO-PVP tab (Fig. 3.8). To separate

the comea fiom the underside of the hydrogel tab, the comea was peeled away gently so

as not to damage the relatively fiagile, curved hydrogel. This lefi a strip that was

essentially half PEG-CO-PVP tab and half comeai web.

Figure 3.4 - Rectangular, partial-depth corneal injury

Figure 3.5 - Corneal injection

Figure 3.6 - Multiple, adjacent corneal injections

Figure 3.7 - Transfer of PEGbMA solution into reservoir

Figure 3.8 - Corneal strip for mechanical testing

Two of the six strips were loaded into the tensometer, in a similar fashion to that

describeci in section 3.3. Since the grips did not cornpress the comeal end of the strip in

the same manner as the hydrogel, the bottom grip was sirnply tightened enough to provide

a hold strong enough to avoid slippage of the sample. Sarnples were stretched until

failure at a speed of 1 mrn/min. Data acquisition occurred with a frequency of 120 Hz.

3.5) Histology of Corneal Webs

Five fiesh bovine eyes were placed in PSS, from which each eye was removed just pior

to handling. A partial-thickness rectangular injury was made in each cornea, as descnbed

in section 3.4.

With a 20-gauge needle, stroma1 injections of PEGbMA solution (formulation #9; see

Table 3.1) were made in four of the five eyes. Single injections were made in two

comeas, three adjacent injections in the other two. AAer injection, two of the eyes sat for

one hour prior to photoinitiation, while the other two were initiated immediately after

injection (see Table 3.3). The fiNi eye was used as a sharn.

Table 3.3 - Pre~sratioa of Corneal Webs for HistoIogical Examination

One hour following initiation, the cornea was excised fiom each eye. A section

containing the comea-polymer "web" was removed fiom each cornea and fixed in

Davidson's solution (see composition on page 80). M e r 40 hours in this solution, the

comeal sections were washed in 0.1 M phosphate buffered saline (PBS; see composition

on page 80) to remove any excess fixative. Sarnples were dehydrated through a graded

Eye A B C D E

No, of Injections one

three one

three none

Delay for Initiation none none

1 hour 1 hour none

series of increasingly strong alcohol solutions. Following clearing of the comeal sections

with Histosol (National Diagnostics, Diarned, Mississauga, ON), the samples were

infiltrated, then embedded, with p d ~ n wax.

Serial sections of seven to eight micron thickness were prepared. For samples with only

one injection site (eyes A and C; see Table 3.3), sagittal sections were taken in order to

observe the injection shape. Transverse sections were used for sarnples with multiple

injection sites (eyes B and D) to ascertain whether the macromer fiom these discrete

injections would form a continuous network. Transverse sections were also used for the

sham sample (eye E) to assess the darnage caused by insertion of the needle. Slides were

stained with haematoxylin and eosin (H and E) and were exarnined under a light

microscope.

3.6) Corneal Gel Chrtracterization

Fourteen fiesh bovine eyes were stored in PSS until just prior to handling. One or two

partial-depth incisions were made in each eye with a type021 scalpel blade. Macromer

solution of formuiation 9 or 10 was injected into each eye with a 25-gauge needle.

Photoinitiation occurred immediately after injections, for a period of two minutes. After

polymerizing for 2% to 3% hours, the corneas were excised with a pair of surgical

scissors. The section of each cornea containing the comeal gels was removed and plated

in 5 mL of a neutral-buffered, 0.222 mg/mL collagenase solution (see composition on

page 80). The corneal sections were placed in a 40°C oven, to activate the collagenase

(Sigma Chernical Co., St. Louis, MO, USA), for a minimum of 20 hours.

3.6.1) Swelling Study

Upon removal of the coneal sections fkom the oven, most of the comeal gels were easily

isolated fiom the surrounding tissue. These gels were placed in distilled water and were

analyzed gravimetrically over a period of 10 days, when the swollen weight was recorded.

This was followed by a drying phase of at least 10 days, at which time the dry polymer

weight was docwnented. A second swelling phase, over 8 to 10 days, provided the

reswollen weight of each gel.

The swollen water content was calculated for each gel, using equation 3.1. The weight

loss between the swollen and reswollen weights was aIso detemined with equation 3.3.

3.6.2) Hydroxyproline Assay

To determine if any collagen remained within the comeal hydrogels, a hydroxyproline

assay was conducted, accordhg to the method of ~ o e s s n e r ~ ~ , on four of the gels.

Hydroxyproline is an imino acid that is very specific to collagen, and typically comprises

about 15 wt% of type 1 collagen.

Each gel was placed in a test tube with 500 pi, of 6N HCI for several hours at 130°C.

After a few drops of methyl red indicator were added, the contents were neutralized with

concentrated NaOH. Final pH adjustrnents were made with dilute HCl and NaOH until

the indicator turned slightly yellow. The neutralized mixture was diluted with water to a

volume of 7.5 mL.

niree 1-mL aliquots of this solution were each diluted with water to a volume of 3 mL.

To each of these aliquots, 1.5 mL of chloramine T was added to initiate oxidation of the

hydroxyproline. After standing for 20 minutes at room temperature, 1.5 mL of perchloric

acid was added to each aliquot to destroy the chloramine T. Following an additional

5-minute period, 1.5 mL of p-dimethylaminobenzaldehyde was added to each aliquot, and

the mixture was placed in a 60°C water bath for 20 minutes. Finally, the contents of each

test tube were cooled with tap water.

A sarnple of each aliquot was analyzed with a UVNis-spectrophotometer. A calibration

curve (Figure A-2) had previously been generated with hydroxyproline standards, over

the range of 0.25 to 2.5 pg/mL. For each hydrogel, the three aliquot concentrations

yielded an average hydroxyproline concentration (Ha"&

The collagen concentration in the original comeal hydrogel was calculated by using

equation 3.9:

where: Ccoiiagc, = concentration of collagen in comeal hydrogel (wt%);

Hwg = average hydroxyproline concentration for three aliquots (mg1mL);

DF = dilution factor for aliquots (7.5);

V,, = volume of neutralized solution (7.5 mL);

HC = hydroxyproline content of type 1 collagen ( 1 5%); and

w, = weight of comeal hydrogel (mg).

4) Results

4.1) Gel Cbaracterization

As illustrated in section 3.1, hydrogel properties, such as crosslinking density and tensile

strength, may be controlled by the composition of the macromer solution from which the

hydrogel is synthesized. A relatively simple method of determining the effects of the

solution composition is by observing the hydrogel's swelling behaviour. A pariicularly

telling characteristic is the gel's water content when swollen to thermodynamic

equili brium.

Tables 4.1 to 4.3 surnmarize the results fiom the swelling study of the 48 gels produced in

section 3.1. Table 4.1 contains the equilibrium swollen water contents for each gel, as

determined by equation 3.1.

Analysis of variance was conducted for these values (see Tables A-1 to A-4 on pages 82

and 83) in order to determine which reaction parameters significantly affected the

hydrogels' equilibrium swollen water content. The algorithm used for these

ca~culations~ is shown on page 81. This analysis showed that PEGbMA concentration

(P > 0.999) significantly affects the swollen water content of the resultant hydrogels.

Table 4.1 - Equilibrium SwoUen Water Content of Gels 1-16 (Series A. B. and Cl

As PEGbMA concentration was increased fiom 25 to 50% w/v, water content decreased

fiom an average of 83.8 +/- 0.19% to 79.9 +/- 0.66% (n = 12 each). This result is due to a

greater crosslinking density within the network as the proportion of bifùnctional

macromer is increased. As the crosslinking density increases, the network becomes more

rigid and does not swell as fieely. The practical importance of this difference in swollen

water content will be discussed in sections 5.1 and 5.2.

Gel No. 1 2 3

Considering only the gels synthesized fiom 25% w/v PEGbMA solution (Table A-2),

N-VP concentration (P > 0.999), and W curing time (P > 0.99) also significantly affect

water content.

[PEGbMA] (% W/V)

50.0 50.0 50.0

Curing T h e (s)

1 20 1 20 120

[DMPA] (ppm) 1800 1800 900

m-VP] (% v/v) 10.0 5.0 10.0

Avg. Water Content (%)

79.6 79.1 80.3

Std. Dev. (%) 0.28 0.18 0.4 1

n 3 3 3

For these gels, as N-VP concentration was increased from 5 to 10% vlv, water content

increased slightly from 83.6 +/- 0.1 1% to 84.0 +/- 0.1 1 % (n = 12 each). With macromer

concentration remaining constant, an increase in the proportion of monohinctional N-VP

molecules will reduce the crosslinking density. A looser network will be forrned, which

will swell to a greater extent.

As W curing time was increased fiom 60 to 120s, the water content decreased slightly

from 83.9 +/- 0.17% to 83.7 +/- 0.20% (n = 12 each). As the duration of exposure to UV

light increases, so does the proportion of photoinitiator that absorbs this energy to

generate radicals. A greater number of radicals should affect only the rate of propagation,

and not the crosslinking density. The observed difference in water content, while of

statistical significance, is insignificant in terms of its effect on the hydrogels' properties

(e.g., mechanical strength, permeability).

Now, considering only the gels synthesized fiom 50% w/v PEGbMA solution (Table A-

l), DMPA concentration (P > 0.999) affects water content. As DMPA concentration was

increased from 900 to 1800 ppm, the swollen water content decreased slightly from 80.4

+/- 0.54% to 79.5 +/- 0.45% (n = 12 each). Again, initiator concentration, or radical

concentration, will increase the rate of reaction, but not the fiequency of crosslinks within

the resultant three-dimensional network. The statistically significant variation in swollen

water content between these two groups of hydrogels does not translate into practically

significant gel properties.

The hydration levels of these gels appear to make them particularly well suited for the

environment found in the stroma. Water content in the stroma is approximately 78%',

which is similar to the hydration of these crosslinked PEG-CO-PVP gels (79 to 84%).

The yields for each hydrogel disk, calculated with equation 3.2, are summarized in Table

4.2. Six of these values were discarded due to difficulties encountered during the

swelling and drying procedures. The primary reason for rejecting these values was

breakage of gels during the onginal swelling phase.

Table 4.2 - Yields of Gels 1-16 (Series A, B. and C)

- -

Gel [PEGbMA] Curing [DMPA] IN-VP] Avg. Std. No. (% W/V) Time (s) (ppm) (% v/v) Yield (%) Dev. (%) n

1 50.0 120 1800 10.0 71.2 8.21 3 2 50.0 120 1800 5.0 79.5 7.07 3 3 50 .O 120 900 10.0 66.6 6.43 3

Due to the unbalanced nature of these data, analysis of variance was not performed for

yield values. It is apparent, however, that gels 1-8, synthesized with 50% PEGbMA

solutions, have lower yields than gels 9-16, fiom 25% PEGbMA solutions. The average

yield for gels 1-8 is 73.3 +/- 10.33%, and for gels 9-16, 80.9 /+- 6.04% (n = 21 each). In

addition to the poor yields, the gels fiom 50% solution were noticeably heterogeneous.

Because of these factors, these formulations were not used in hiture experiments, in order

to focus on the more homogeneous gels synthesized fiom 25% PEGbMA solutions.

The dry weights of these hydrogels were found to exceed the weight of PEGbMA in the

macromer solutions, indicating that there was a significant amount of N-VP incorporated

into the gels. FTIR spectroscopy was performed on several hydrogel samples to confirm

the presence of N-VP. Qualitative analysis of the resulting spectra revealed several peaks

similar to the N-VP ~ ~ e c t r u r n ~ ~ , including the N-H peak at approximately 3400 cm-'.

Table 4.3 summarizes the differences between the swollen and reswollen weights for each

gel, determined with equation 3.3. The large differences s h o w were the result of the

drying process, which tended to darnage the gels. As descnbed in section 3.3, gels such

as these that were synthesized by using tissue culture wells as molds had a very non-

unifom thickness. Thus, when the gels were dehydrated, the thinner central portion dried

more quickly than the thicker region near the disks' perimeter. This drying pattern caused

undue stresses on the disks, which typically curved in toward their centres.

These physical stresses induced tearing in most gels, but did not cause sections to

completely break off. Since there was no gel fiacluring on a macroscopic scale, the

difference in weight encountered upon reswelling appears to be due to the severing of

many of the covalent bonds within the crosslinked network. This bond cleavage would

release many minute polymer fiagrnents within the bulk of the gel. Upon reswelling,

these fragments would be washed out of the hydrogel, which would reach a much lower

equilibrium swollen weight than when first swollen.

Ten of the weight loss values have been rejected, primarily due to gel breakage during the

reswelling phase.

As a result of this observation, smdler pieces of hydrogels were used in al1 future

swelling work in order to minimize gel damage.

Table 4.3 - Swollen Weight Loss for Gels 1-16 (Series A, B. and Cl

Gel No.

v 1 2 3 4 5 6 7 8 9 I O 11 12 13 14 15 16

[PEGbMA] (% W/V)

50.0 50.0 50.0 50.0 50.0 50.0 50.0 50.0 25.0 25.0 25.0 25 .O 25.0 25.0 25.0 25.0

Curing Time (s)

120 120 120 120 60 60 60 60 120 120 120 120 60 60 60 60

[DMPA] @pm) 1800 1 800 900 900 1800 1800 900 900 1800 1800 900 900 1800 1800 900 900

[N-VP] (% vlv)

10.0 5 .O 10.0 5 .O 10.0 5.0 10.0 5 .O 10.0 5.0 10.0 5.0 10.0 5 .O 10.0 5 .O

Avg . Lost (%)

26.3 28.3 29.6

- 28.0 29.1

,- 28.2 25.3 27.6 31.1 28.2 22.0 29.2 33.6 23.7 25.0

S td. Dev. (%)

0.82 O. 12 6.33

- 3.35 1.52 1.79

- 0.89 0.88 0.62 0.49 0.96 1 .O1 6.03 6.33

n 3 2 2 O 3 3 3 1 2 3 2 3 3 2 3 3

4.2) Permeation Studies

An essential characteristic of any material implantecl in the comea is its penneability.

Many corneal nutrients, such as glucose, amino acids, and vitamins, originate primarily in

the aqueous humour? From here, these constituents must diffuse through the comea to

reach keratocytes in the anterior stromal layers. Embedding a rnaterial of low

permeability within the comeal stroma will inhibit this diffusion, and cm lead to necrosis

of the anterior comeal tissue1*.

To determine whether crosslinked PEG-CO-PVP hydrogels are sufficiently permeable,

permeation studies were carried out on eight hydrogel disks, as outlined in section 3.2.

CafKeine, with a molecular weight of 194.2, was the permeant chosen in order to

approximate the behaviour of principal stromal nutrients such as glucose (MW = 180.2)

and amino acids (MWgl,,, = 75.1 to MWwpphn = 204.2).

Figure 4.1 shows the cafTeine concentration in the receptor ce11 during one of these trials,

and is typical of the results observed for each of the eight gels. Regression analysis was

performed only on the fuial four points of this curve. The samples corresponding to these

points were taken d e r the lag phase, when both the permeant concentration profile

within the gel and the rate of caffeine permeation across the gel reached steady state. Lag

phases for the eight samples investigated ranged fiom about 2 1 to 27 minutes.

Figure 4.1 - Permeant concentration in receptor half cell

Knowing the value of (dC,/dt) for each gel, perrneabilities were calculated using

equations 3.7 and 3.8. These values are summarized in Table 4.4 below. The values

show in the 1st column were normalized for hydrogel thickness (0. By studying the

swelling behaviour of smail gel sections, as described in section 3.2, the water content of

each hydrogel was determined.

Table 4.4 - Permeabilities of PEG-co-PVP Hydronels

Gel No.

Water Content (% H20)

Penneability x 1 o6 (cm2/s)

The permeabilities in Table 4.4 show no definite trend with respect to hydrogel water

content. This is not surpnsing, as the range of hydrations for these hydrogels is too

narrow for there to be any practicai difference in their permeabilities. Over a much wider

range of gel hydration, one would expect pemieability to increase as water content

increases, or as crosslinking density decreases.

4.2.1) Gel Characterization

Weight loss values for the gels created in section 3.2 are surnmarized in Table 4.5 below.

The difference between the swoilen and reswollen weights for these gels was much lower

than that for the gels produced in section 3.1. It was suggested in section 4.1 that damage

was infiicted on the gels of section 3.1 by an uneven drying process. Based on this

hypothesis, the substantially lower weight loss values presented in Table 4.5 were

expected, indicating that drying of smaller, more uniform gel sections does not have a

similar effect on the integrity of the hyàrogel structure.

Table 4.5 - Swollen Weiebt Loss for Gels 9-16 (Series El

Gel No. 9E 1 OE 11E

Water Content (%) 83.5 84.0 84.1

Weight Loss ('w 3.5 1.4 3.3

4.3) Mechanical Strength of Hydrogels

One potentiai complication that may lead to keratoprosthesis extrusion is mechanical

failure of the device. Failure may occur at the stroma-skirt interface, at the skirt-core

interface, or within the skirt material that bridges the gap between the tissue and the optic.

Any matenal being considered for use in a keratoprosthesis must be sufliciently strong to

withstand the mechanical forces encountered in the eye, such as normal intraocular

pressure and rubbing of the eye.

To determine the mechanical strength of PEG-co-PVP hydrogels, dumbbell-shaped

tensile strips were cut out of gels and were tested as described in section 3.3. Figure 4.2

shows a typical stress-strain cuve obtained fiom one of these samples.

O 2 4 6 8 10

Strain (Oh)

Figure 4.2 - Stress-strain curve for hydrogel strip

Results fiom the mechanical strength testing of hydrogel strips have been sumrnarized in

Table 4.6. Each hydrogel strip failed, as expected, within its shaight central portion (see

Fig. 3.2).

Table 4.6 - Failure Stress and Strain of Hvdroeel Str i~s 184-85% Water Content)

Failure Stress Failure Strain Gel No. (%)

The hi& vanability of these results is clearly noi due to hydration differences, as al!

hydrogel sarnples had hearly identical water content. Thus, an analysis of variance was

carried out to determine if there was any correlation between tensile results and any of the

four experimental factors given on page 21. This analysis demonstrated no such

correlation.

An altemate explanation for these variations in tensile properties relates to imperfections

in the hydrogels' structures. These faults may be chernical in nature, but in this instance

were more likely induced while rnanually cutting the tensile strips. Since the dimensions

of these strips were so small, even a minute flaw on the perimeter of a sample would

substantially weaken its structure at this point, resulting in prernature failure.

4.3.1) Gel Characterization

Table 4.7 contains the differences behveen the swollen and reswollen weights of hydrogel

sections produced in section 3.3. These relatively low values anim the suspected

damage inflicted on the gels fiom section 3.1 during their drying phase.

Table 4.7 - Swollen Weieht Loss for Gels 9-16 (Series L)

Water Content

4.4) Mechanical Strength of Corneal Webs

As stated in section 4.3, mechanical failure of a keratoprosthesis may occw at the

interface between the stroma and the skirt. To determine if the comeal webs produced in

section 3.4 were sufficiently strong to withstand typical ocular forces, strips were cut out

of these comeas and their tensile strengths were determined, as outlined in section 3.4.

Four of the six comeal web strips prepared were destroyed during handling. Although

precautions were taken to avoid such problems, many of the hydrogel tabs were too weak

to withstand even minimal manipulation. The poor mechanical properties of these gels

was possibly due to their heterogeneous structure - these samples were much less uniform

than the hydrogels produced in cylindrical molds. Their curved structure, con forming

with the surface of the comea, also contributed to the difficulties encountered, as

stretching of the curved surfaces in a linear manner imposed additional strain on the

samples. Regardless, the relative ease with which the hydrogel tabs ruptured attests to

their poor mechanical strength.

Each of the two remaining samples was tested on a tensometer. One of these had been

initiated immediately following injection of the macromer solution; the other had been

photoinitiated one hour following injection. Both samples failed not at the cornea-

polymer interface, but 2-3 mm away fiom this point, within the hydrogel tab (see Figure

4.3).

Fai lw stresses for the samples initiated immediately and one hour after injection were

14.9 and 14.2 kPa, respectively. Neither sample failed at the interface between the

comeal tissue and the hydrogel. Thus, it was not possible to determine the effect of

macromer diffusion through the stroma pior to initiation on the mechanical strength of

the comeai webs. Due to the imprecise elongation readings that were obtained during

elongation of these samples, failure strains were not calculated.

comeai

PEGbMA ta b

Figure 4.3 - Failure point of corneal web samples

4.5) Histology of Corneal Webs

Figure 4.4 displays the slides prepared for eyes A through E, respectively. Note that

comeal sample D swelled, and is not at a higher magnification. The larger, single

injections s h o w in slides A and C produced gels (labelled "PEG') that distended the

surrounding stromal matrix ("s"). These slides also show lighter staining around the

perimeter of the gels, suggesting that a minimal amount of stroma-polyrner interaction

occurs in these regions.

Figure 4.4 - Histology of corneal webs

A) one injection, immediate initiation (see Table 3.2); B) three injections, immediate initiation;

both samples sectioned transversely, stained with H and E; ep = epithelium, en = endothelium,

s = stroma, PEG or arrow = in situ hydrogel; Bar = 50 pm

Smaller, multiple injections were used for sarnples B and D. The corresponding slides

appear to have produced several smaller hydrogel domains, indicated by arrows. This

suggests that smaller injections produce greater interdigitation between the PEG-CO-PVP

gel and the surrounding stroma. Again, regions of stroma-polymer interaction are more

lightly stained, particularly in slide B.

Since slides A and C are so similar, as are slides B and D, it appears that an insignificant

amount of diffusion occurs during the hou-long delay between injection and

photoinitiation that was imposed on eyes C and D. Figure 4.5 contradicts this result,

showing a distinct clifference between the appearance of macromer injections that were

initiated immediately (at left) and those initiated one hour after injection (at right).

Figure 4.4 (cont'd) - Histology of corneal webs

C) one injection, one-hour delay; D) three injections, one-hour delay; E) no injections (sham);

al1 samples sectiooed sagittally, stained with H and E; ep = epithelium, en = endothelium, s = stroma,

PEC or arrow = in situ hydrogel, n = needle injury; Bar = 50 pm

The sharn sample (Figure 4.4E) indicates the damage caused by insertion of the needle

("n").

Figure 4.5 - Corneal Injections, with (right) and without (left)

one-hour delay prior to photoinitiation

4.6) Corneal Gel Characterization

Another method of establishuig how an in sPu polyrnerized network interacts with the

surrounding collagen is by examining the gels themselves. Comeal gels were prepared

and isolated as described in section 3.6.

4.6.1 ) Swelling Study

By observing the swelling behaviour of the corneal gels, one may be able to determine

whether collagen fibres were at any time incorporated into the bulk of these gels.

Thirteen gels were analyzed gravirnetrically; the results from eleven of these gels are

summax-ized in Table 4.8. The other two gels had excessive arnounts of collagen that

were not properly removed fiom their surfaces during treatment with collagenase. This

collagen evenhtally degraded and fell off of the gels.

Table 4.8 - Swelling Characteristics of Corneal Gels

Gel No, 2a 3a

Formulation No. 10 10

Water Content (%) 85.1 83.6

Weight Loss (%) 5.7 1.9

The swollen water contents shown in Table 4.8 are very similar to those (Table 4.1)

obtained for the gels made in section 3.1 fiom solutions of identical composition. The

hydration of gels made fiom formulation 10 ranged From 83.4 to 83.5%; gels made from

formulation 9 had hydrations of 83.8 to 84.0%.

The low weight losses s h o w in Table 4.8 the damage done by uneven drying of

the gels fiom section 3.1.

4.6.2) Hydroxyproline Assay

Table 4.9 contains the data generated by the hydroxyproline assay as descnbed in section

3.6.2. These data indicate that only a small arnount of collagen (about 2 wt%) was

contained in each of the corneal gels. This result is in agreement with the slides shown in

Figure 4.4, where large, single injections were shown to produce gels that did not interact

significantly with the surrounding stroma. The small amount of collagen present in the

comeal gels suggests that a minimal amount of interpenetration occurs between the gel

and the stroma's collagen fibrils, near the perimeter of the hydrogel.

Table 4.9 - Colla~en Content of Corneal Gels

Gel No. 10a lob 12a 15b

Avg. Abs. (562 nm) 0.354 0.27 1 O. 175 O. 187

Hsy~ (mg/mL) 0.0023 0.00 1 8 0.001 1 0.00 12

Gel Wt. (mg) 42.2 34.9 25.5 17.0

% Collagen 2.04 1.89 1.67 2.68

5) Discussion

5.1) Permeability of PEG-CO-PVP Hydrogels

It is essential that stromal implants have adequate permeability in order to ensure proper

nutrition of cells in the anterior stromal layers. Nutrients such as glucose and amino

acids, found primarily in the aqueous hurno~r'~, must diffuse across the entire corneal

thickness to reach these layers. Any implant with inadequate pemeability would inhibit

this difision and could lead to anterior comeal necrosis".

Caffeine was selected as the penneant for a set of permeation studies due to its molecular

weight of 194.1, which is similar to that of glucose (MW = 180.2) and most amino acids

(MW = 75.1 to 204.2). The results of these studies, reported in section 4.2, indicate that

the hydrogel disks produced in section 3.2 have an average permeability of 1.72 x 1om6 +/-

1.69 x 10" cm2/s (n = 8).

Sweeney et a p implanted a variety of porous membranes within the stroma, studying the

effect of membrane pore size on the post-implantation health of the comea. Membranes

with pore sizes ranging fiom 15 to 100 nm were investigated by monitoring treated eyes

and their levels of clinical response. The glucose permeability of each type of membrane

was measured both before implantation and foilowing explantation.

Only the membranes of 100-nm pore size were well-tolerated in the cornea, with 70% of

the treated eyes maintaining a clinical tolerance for these membranes for a penod of 100

days. The pre-implantation permeability of these membranes, adjusted for membrane

thickness, was found to be 5.08 x 1 ~ ' ~ +/- 1.56 x 1 ~ ' ~ cm2/s (n = 5). This value is

exceeded by the average permeability of the PEG-CO-PVP hydrogels, i -72 x 1 0 ' ~ +/- 0.17

x 1 o4 cm2/s, by over two orders of magnitude.

Thus, assuming that glucose and caffeine penneate similarly, the results of Sweeney et al

suggest that the PEG-CO-PVP hyàrogel disks produced in section 3.2 are sufficiently

penneable to be implanted into the stroma.

5.2) Tensile Streagth of Fixative

Any fixative system such as the one employed in this project should have a minimum

tensile strength that is adequate to at least withstand typical ocular forces. Normal

intraocular pressure is approximately 20 mm H~"*'' (2.6 kPa), and cm change by more

than 5 mm Hg (0.7 kPa) depending on whether a person is sitting, lying down, or

i n ~ e r t e d ~ ~ . Other normal ocular forces, such as squinting, blinking, or mbbing of the eyes

will increase this by several times, to a pressure on the order of 10 to 20 kPa.

Ideally, the fixative would demonstrate a tensile strength similar to that of the corneal

stroma. The stroma possesses most of the mechanical strength of the cornea4', although

the precise distribution of tensile strength between layers of the comea has not been

detemined. Full thickness comeai strips have been shown to have a failure stress of 19.1

+/- 3.50 M P ~ ~ * . However, it is believed that a fixative with tensile properties similar to

those of the sclera would be more than adequate. The ocular globe has a bursting

pressure of approximately 500 k ~ a ~ ~ , which is still two orders of magnitude greater than

normal intraocular pressure. Thus, the desired tensile strength of the comeal webs falls

between 10 and 500 kPa.

The tensile strength of the fixative system examined in this project was determined at two

locations: within the gel phase and at the fixative-strornal boundary. Failure stresses of

hydrogel strips have been reported in section 4.3. These values ranged fiom 52 to 97 kPa,

for swollen gels of 84.5 to 85.0% hydration. The hydrogels alone appear to be strong

enough to meet the minimum critenon established above.

Chirila et al have measured the mechanical properties of the porous skirt segment of their

PHEMA homo-IPN keratoprosthesis mode17. Failure stresses of a 70%-hydrated and an

80%-hydrated porous gel were reported to be 323.3 +/- 27.2 kPa and 52.24 +/- 1.15 kPa,

respectively. This latter value is comparable with the tensile strengths reported in section

4.3 (52 to 97 kPa) for PEG-co-PVP hydrogels. However, Chirila's group has encountered

problems with its porous skirt during implantation of the device. The porous PHEMA is

so brittle that sutures rnust be passed through the clear PHEMA optic core in order to

hold the keratoprosthesis in place7.

As the above data indicate, by creating a gel with considerably lower swollen water

content, Chirila et al were able to greatly increase its tensile strength. The PEG-CO-PVP

hydrogels codd be rnanipulated similarly.

Since the permeability of these hydrogels is more than adequate, it may be desirable to

modi@ their crosslinked structure by altering the composition of the macromer solution

fiom which they are formed. A substantiai increase in crosslinking density will decrease

the water content of the swollen gel, improving its mechanical properties while

maintaining sufficient penneability for water-borne nutrients. The increase in

crosslinking density could be achieved either by raising the concentration of bifunctional

macromer (PEGbMA) in the macromer solution, by lowering the concentration of

monofunctional monomer (N-VP), or by considerably decreasing the molecular weight of

the PEGbMA macrorner.

Notwithstanding the improved mechanical properties of the less-hydrated PHEMA skin,

Chirila's group has been forced to use the more highly hydrated PHEMA in order to

ensure adequate pore size for keratocyte into the s l~ i r r '~~ '~ , essential for their

keratoprosthesis fixation method. By using an in situ polymerization system, this

infiltration would not be a concern, as the crosslinked network would be formed in the

presence of, and around, these cells.

Comeal webs were produced as described in section 3.4, and their mechanical properties

were reported in section 4.4. Four of the six samples prepared were so weak that even

careful manipulation of these samples damaged them irreparably. The two samples that

were tested exhibited tensile strengths of 14.9 and 14.2 kPa. These values are marginally

within the suitable range, as designated above.

Both samples failed within the PEG-CU-PVP tab portion of the tensile strip, as indicated

in Figure 4.3. This suggests that the crosslinked hydrogels produced in contact with the

comeal surface were not as mechanically strong as those produced in molds. As stated

earlier, this was likely due to either the heterogeneous nature of these gels, or their

curvature, induced by the corneal surface upon which it was formed. When these curved

hydrogels were stretched, additional tension was encountered by the imer surface,

potentially causing a premature failure.

Since neither sample failed at the stroma-polymer interface, it is reasonable to assume

that the comeal webs exhibited a tende strength of greater than 15 kPa. This is an

encouraging result, as the mechanical properties of the PEG-CO-PVP tab could probably

be improved more predictably than could the strength of the comeal web.

5.3) Interaction Between Polymer and CoUagen

Although the comeal web formed is strong enough to withstand normal intraocular

pressure, it doesn't appear to interact as well as expected with the surrounding stromal

collagen. The histological findings presented in Figure 4.4 indicate that interaction

between the in situ polymerized network and the surrounding collagen fibres occurs only

near the periphery of macromer injections. Only when several adjacent injections are

executed (Figures 4.4B and 4.4D), does the collagen matrix become entrapped within the

crosslinked polymer. The observations presented in Figure 4.5 indicate that collagen may

become incorporated into the crosslinked network if suficient time is allowed for the

injected macromer solution to dif ise throughout the stroma prior to photoinitiation.

This lack of stromal-polyrner interaction is confinned by the comeal gel characterization

described in section 3.6.1. Results of this evaluation, presented in section 4.6.1, indicate

that the hydration of each comeal gel was nearly identical to that of the corresponding

hydrogel disk synthesized fiom the same solution formulation. But, since stromal

collagen swells to approximately the same degree (78%) as the PEG-CO-PVP gels,

collagen content would not necessarîly be detected by examining the gels' swelling

characteristics. Thus, the results of section 4.6.1 could suggest that these gels contained a

large number of embedded collagen fibrils that were not cleared by collagenase treatment.

However, results of the hydroxyproline assay, presented in section 4.6.2. indicate that

only a very srnall amount (2% w/w) of collagen was present in the comeal gels,

confirming the low level of stromal-polymer interaction discovered from histological

examination.

It may be difftcult to improve this interaction. One of the unique properties of PEG is its

tendency to exclude proteins and other macromolecules, including collagen, when in

aqueous ~ o l u t i o n ' ~ ~ ~ ~ . This property makes PEG ideal for applications such as protein

precipitation and coating of biomaterial surfaces (to prevent protein adhesion), but it

renders PEGbMA less effective in the present application. To ensure maximum

interaction with stroma1 collagen at injection boundaries, volumes of individual

PEGbMA injections should be minimized, and a delay should be incorporated into the

procedure, pnor to photoinitiation, to allow for diffusion of the injected solution.

6) Conclusions and Recommendations

Permeation studies showed that crosslinked PEG-CO-PVP hydrogel disks with 84% water

content have a permeability of 1.72 x 10" cm2/s, more than adequate for corneal

implantation.

PEG-CO-PVP hydrogels made fiom 25% PEGbMA solutions have an average tensile

failure stress of 70.3 kPa, while the cwrent fixative mode1 as a whole has a tensile

strength of approximately 15 kPa.

An attempt should be made to improve mechanical properties of the PEG-CO-PVP

hydrogels by decreasing their swollen water content. This effect is achieved by

increasing the crosslinking density of the three-dimensional network. Crosslinking

density may be increased either by reducing the concentration of the monofunctional

cornonomer N-VP, by increasing the concentration of bifùnctional macromer (PEGbMA),

or by using a macromer of lower molecular weight. It is recomrnended that this

polymerization system be studied M e r to understand more precisely the effects of

reaction parameters on swelling and mechanical properties.

Histology of corneal injection samples showed that in situ polymerized PEG-CO-PVP

does not integrate well with the surroundhg stroma. This is a result of the tendency of

PEG-based molecules in aqueous environrnents to exclude other macromolecules,

including proteins such as c ~ l l a ~ e n ' ~ * ~ ~ . It has also been demonstrated that srnaller,

adjacent injections interact with the stroma more intimately than larger injections.

Macromers of lower molecular weight, in addition to increasing the crosslinking density

of the network in situ, may also diffise more easily through the stroma, enhancing

interpenetration between collagen larnellae and the crosslinked network.

To facilitate funw comeal injections, it is recommended that a microinjection system be

used. This system could enable one to perform simultaneous injections at multiple sites.

Use of a syringe pump or similar mechanism will allow control over individual injection

volumes, which is also important. Smaller, discrete injections at well-defined locations

will improve cornea-polymer interaction.

A new template should be developed to create more reproducible and more easily tested

comea-polyrner interfaces. This template should produce a flat hydrogel tab of u n i f o n

thickness that does not adhere to the cornea. A possible configuration of this template is

shown in Figure 6.1, with a curved imer surface that rests on the eye, and a Bat upper

surface into which macromer solution can be injected prior to photoinitiation.

Figure 6.1 - Template for reproduciblc comeal web

and hydrogel tab synthesis

It appears that the development of an in situ polymerized fixative for a keratoprosthesis is

possible, but will be difficult to achieve. This thesis has shown that two crucial

properties of this fixative, its permeability and its mechanical strength, are not

complimentary. That is, as a hydrogel's crosslinking density decreases, its permeability is

enhanced, but its mechanical strength is weakened, and vice versa. The challenge will be

to find an optimal combination of these characteristics, as well as a precise, reproducible

method of introducing the fixative into the corneal environment.

7) Referenees

1) Maurice DM: The Comea and Sclera. pp. 1-129 In: The Eye (Vol. lb). Davson H (ed.), Academic Press, Orlando, FL (1984).

2) Freegard TJ : The phy sical basis of transparency of the normal comea. Eye 1 1 : 46 5- 471 (1997).

3) Chirila TV, Hicks CR, Dalton PD, Vijayasekaran S, Lou X, Hong Y, Clayton AB, Ziegelaar BW, Fitton JH, Platten S, Crawford GJ, Constable IJ: Artificial Comea. Progress in Poiymer Science 23: 447-473 ( 1 998).

4) Barron BA: Prosthokeratoplasty. pp. 879-896 In: The Comea (2nd ed.). Kaufman HE, Barron BA, McDonald MB (eds.), Butterworth-Heinemann, Boston (1997).

5) Chirila TV: Modem artificiai comeas: the use of porous polymers. Trends in Polymrr Science 2: 296-300 (1994).

6) Chen YC, Chirila TV, Russo AV: Hydrophilic sponges based on 2-hydroxyethyl methacrylate. II: Effect of monomer mixture composition on the equilibriurn water content and swelling behaviour. Materials Forum 17: 57-65 (1993).

7) Chinla TV, Yu DY, Chen YC, Crawford GJ: Enhancement of mechanical strength of po ly(2-hydroxyethy l methacrylate) sponges. Journal of Biomedical Materials Research 29: 1029-1 O32 (1995).

8) Pathak CP, Sawhney AS, Hubbell JA: Rapid photopolymerization of imrnunoprotective gels in contact with cells and tissue. Journal of the American Chernical Society 114: 83 1 1-83 12 (1992).

9) Sawhney AS, Pathak CP, Hubbell JA: Bioerodible hydrogels based on photopolymerized poly(ethy1ene glycol)-co-poly(a-hydroxy acid) diacrylate monomers. Macromolecules 26: 58 1-5 87 (1 993).

10) Sawhney AS, Pathak CP, Cox PR, Hubbell JA: Conformal buriers by interfacial po lymerization in contact with cells and tissue. Proceedings of the American Chernicol Sociew Division of Polymeric Muterials - Science & Engineering 69: 526 ( 1993).

I l ) Hill-West JL, Chowdhury SM, Sawhney AS, Pathak CP, Dunn RC, Hubbell JA: Prevention of postoperative adhesions in the rat by in situ photopolymerization of bioresorbable hydrogel barriers. Obstetrics & Gynecology 83: 59-64 ( 1 994).

12) Hill-West JL, Chowdhury SM, Slepian MJ, Hubbell JA: Inhibition of thrombosis and intima1 thickening by in situ polymerîzation of thin hydrogel barriers. Proceedings of the National Academy of Sciences of the USA 91: 5967-597 1 (1 994).

13) West JL, Hubbell JA: Cornparison of covalently and physically cross-linked polyethylene glycol-based hydrogels for the prevention of postoperative adhesions in a rat model. Biomuterials 16: 1 153-1 1 56 (1995).

14) Harris JM: Introduction to Biotechnical and Biomedical Applications of Poly(ethy1ene glycol). pp. 2-14 In: Poly(ethy1ene glycol) Chemistry - Biotechnical and Biomedical Applications. Harris JM (ed.), Plenum Press, New York (1 992).

15) Klyce SD, Beuerman RW: Structure and Function of the Comea. pp. 3-50 In: The Comea (2nd ed.). Kaufinan HE, Barron BA, McDonald MB (eds.), Butterworth- Heinemann, Boston (1 997).

16) Smelser GK, Ozanics V: New Concepts in Anatomy and Histology of the Comea. pp. 1-20 In: The Comea. World Congress. King JH, McTigue JW (eds.), Butterworth Inc., Washington (1 965).

1 7) de Quengsy P: Precis ou cours d 'operations sur la chirurgie des yeux. Didot, Paris ( 1 789).

1 8) von Nussbaum A, Nepomuk J: Cornea Art@als. Schurich, Munchen ( 1 853).

19) Fyodorov SN, Moroz 21, Zuev VK: Keratoprostheses. Churchill Livingstone, Edinburgh (1 987).

20) Stone W, Herbert E: Expenmental study of plastic material as replacement for the comea: A preliminary report. Arnerican Journal of Ophthalrnology 36: 1 68- 1 73 (1 953).

2 1) Ruedeman AD: Silicone keratoprosthesis. American Ophthalmological Society Transactionr 72: 329-359 ( 1 974).

22) Polack FM, Heimke G: Ceramic keratoprostheses. Arnericon Academy of Ophthahoiogy 87: 693-698 (1 980).

23) Stone W, Yasuda H, Refojo MF: A 15-year study of the plastic artificial cornea - basic principles. In: The Comea. World Congress. King .TH, McTigue JW (eds.), Butterworths, Washington (1 965).

24) Chirila TV, Constable U, Crawford GJ, Russo AV. US. Patent No. 5,300,116 (1 994).

25) Chirila TV, Constable IJ, Crawford GJ, Russo AV. ( 1995).

26) Jacob-Labarre .TT, Caldwell DR: Development of a new

74

U.S. Patent No. 5,458,819

type of artificial comea for treatrnent of endstage comeal diseases. pp. 27-39 In: Progress in Biomedical Polymen. Gebelein CG, Dunn RL (eds.), Plenum Press, New York (1 990).

27) Caldwell DR, Jacob-Labarre JT. U.S. Patent No. 4,865,601 (1989).

28) Caldwell DR, Jacob-Labarre JT. U.S. Patent No. 4,932,968 (1 990).

29) Legeais JM, Renard G, Parel JM, Savoldelli M, Pouliquen Y: Keratoprosthesis with biocolonizable microporous fluorocarbon haptic. Archives in Ophthalmology 113: 757- 763 (1 995).

30) Legeais JM, Renard G, Thevenin D, Pouliquen Y: Advances in artificial corneas. Investigative Ophthalmofogy & Visual Science 36 (Su ppl.): 3 1 4 ( 1 995).

31) Tsuk AG, Trinkaus-Randall V, Leibowitz HM: Advances in polyvinyl alcohol hydrogel keratoprostheses: Protection against ultraviolet light and fabrication by a molding process. Journal of Biomedical Materials Research 34: 299-304 (1 997).

32) Trinkaus-Randall V, Banwatt R, Capecchi J, Leibowitz HM, Franzblau C: In vivo fibroplasia of a porous polymer in the comea. Investigative Ophthalmology & Visual Science 32: 3245-325 1 (1991).

33) Trinkaus-Randdl V, Banwatt R, Wu XY, Leibowitz HM, Franzblau C: Effect of pretreating porous webs on stromal fibroplasia in vivo. Journal of Biomedical Materials Research 28: 1 95-202 (1 994).

34) Chirila TV, Constable IJ, Crawford GJ, Vijayasekaran S, Thompson DE, Chen YC, Fletcher WA: Poly(2-hydroxyethyl methacrylate) sponges as implant materials: in vivo and in vitro evaluation of cellular invasion. Biomaterials 14: 26-38 (1993).

35) Crawford GJ, Constable IJ, Chirila TV, Vijayasekaran S, Thompson DE: Tissue interaction with hydrogel sponges implanteci in the rabbit comea. Cornea 12: 348-357 (1 993).

36) Chirila TV, Vijayasekaran S, Home R, Chen YC, Dalton PD, Constable IJ, Crawford GJ: Interpenetrating polymer network (IPN) as a permanent joint between the elements of a new type of artificial cornea. Journal ofBiornedical Materials Research 28: 745-753 (1 994).

37) Potts TV, Petrou A: Argon laser initiated resin photopolymerization for the filling of root canals in human teeth. Lasers in Surgery and Medicine 11 : 257-262 (1 99 1).

38) Khor E, Li HC, Wee A: In situ polyrnenzation of pyrrole in animal tissue in the formation of hybnd biomaterials. Biomaterials 16: 657-66 1 (1 995).

39) Graham NB: Poly(ethy1ene glycol) gels and h g delivery. pp. 263-281 In: Poly(ethy1ene glycol) Chemistry - Biotechnical and Biomedical Applications. Harris JM (Ed.), Plenum Press, New York (1992).

40) Peppas NA: Hydrogels. pp. 60-64 In: Biomaterials Science: An Introduction to Materials in Medicine. Ratner BD, Hoffman AS, Schoen FJ, Lemons JE (Eds.), Academic Press, San Diego ( 1996).

41) Decker C: Photoinitiated crosslinking polymensation. Progress in Polymer Science 21: 593-650 (1 996).

42) Anseth KS, Bowman CN, Brannon-Peppas L: Mechanical properties of hydrogels and their expenmental determination. Biomaterials 17: 1647- 1657 (1 996).

43) Woessner JF: The determination of hydroxyproline in tissue and protein samples containing small proportions of this imino acid. Archives of biochemistry and biophysics 93: 440-447 ( 1 956).

44) Winer BJ: Statistical Pnnciples in Experimental Design (2nd ed.). McGraw-Hill Book Company, New York (1 97 1).

45) The Infrared Spectra Atlas of Monomers and Polymers. Sadtler Research Laboratories, Division of Bio-Rad Laboratories, Inc., Philadelphia (1 980).

46) Sweeney DF, Xie RZ, O'Leary DJ, Vannas A, Odell R, Schindhelm K, Cheng HY, Steele JG, Holden BA: Nutritional requirements of the corneal epitheliurn and anterior stroma: Clinical findings. Investigative Ophthalmology & Visuul Science 39: 284-29 1 ( 1998).

47) Davson H: The Intra-Ocular Pressure. pp. 147-196 Ln: The Eye (Vol. 1 ). Davson H (Ed.), Academic Press, New York (1962).

48) Pinsky PM, Datye DV: A microstructurally-based finite element mode1 of the incised human comea. Journal of Biomechanics 24: 907-922 ( 1 99 1 ).

49) Maurice DM: The Cornea and Sclera. pp. 489-600 In: The Eye (Vol.1, 2nd ed.). Davson H (Ed.), Academic Press, New York (1 969).

50) Bryant MR, Szerenyi K, Schmotzer H, McDonnell PJ: Corneal tensile strength in fully healed radial keratotomy wounds. Investigative Ophthalmology & Visual Science 35: 3022-303 1 (1994).

5 1) Maurice DM: The Comea and Sclera. pp. 289-368 In: The Eye (Vol. 1). Davson H (Ed.), Academic Press, New York (1962).

52) Lim K, Herron JN: Molecular Simulation of Protein-PEG Interaction. pp. 29-56 In: Poly(ethy1ene glycol) Chemistry: Biotechnicai and Biomedical Applications. Harris JM (Ed.), Plenum Press, New York (1992).

Appendix

Caffeine Calibration Curve

Hydroxyproline Calibration Curve

Miscellaneous Solutions

Winer's Algorithm (ANOVA)

ANOVA Tables

1 .O0 1.50 Concentration (pglmL)

Figure A-2 - Hydroxyproline Calibration Curve (dope = 0.1541)

Phosphate Buffcred Saline (0.1 M):

dilute with distilled H20 to 1 L adjust pH to 7.2-7.4 as necessary

Physiological Salt Solution:

32.03 g NaCl 0.799 g KCl 0.799 g CaC12 0.708 g MgCl2.6H20 0.375 g Na.H2PO4*H2O 4.00 g NaHC03 4.01 g glucose

dilute with distilled H20 to 4 L

Davidson's Solution:

100 mL glacial acetic acid 300 mL 95% ethanol 200 rnL 10% neutral formalin 300 mL distilled H20

NeutraCBuffered Collagenase Solution:

10 m M CaC12 50 rnM Tris-HC1 buffer

adjust pH to 7.4 add collagenase to a concentration of 0.222 mg/rnL

Winer's Aleoritbrn for Throc Factor ANOVA'~:

A, B, C = three factors being considered n = total no. of observations

X = any observation p = no. of levels of factor A

G = sum of al1 observations q = no. of levels of factor B

i, j, k = level of factors A, B, C, respectively r = no. of levels of factor C

m = replication no.

ss, = (3) - (1) SSb = (4) - (1) SSc = (5) - (1)

ss,b = @ ) - (3) - (4) + (Il ssu = (7) O (3) O (5) +(1)

SS, = (8) - (4) - (5) + (1) ss,, = (9) - (6) - (7) - (8) + (3) + (4) + ( 5 ) - (1)

ss = (2) - (9) CU

ssw, = (2) - (1)

Table A-' ANOVA (

l Source of Variation

i20 Content); al1 50% PEGbMA gels

Table A-2 ANOVA (H20 Content); al1 25% PEGbMA gels

l Source of Variation

A - UV Time

6 - conc. DMPA

C - conc. N-VP

A - UV Time

B - conc. DMPA

C - conc. N-VP

Table A-3 ANOVA (H20 Content); al1 120s gels

Source of Surn of Degrees of Mean Variation Squares Freedorn Squares

1

Total 9.2E-3 1 23

Table A 4 ANOVA (H20 Content); al1 60s gels

Source of Variation

A

B

C

AB

AC

BC

ABC

1 Total 1 9.9E-3 1 23 1 I

Sum of Squares

9.4E-3

45.1 E-6

, Residual

A - conc. PEGbMA

B - conc. DMPA

C - conc. N-VP

75.8E-6

46.3E-6

3.2E-6

6.3E-6

2.6E-9

A - conc. PEGbMA

B - conc. DMPA

C - conc. N-VP

Degrees of Freedom

1

1

278.4E-6

1

3

1

1

1

Mean Squares

9.4E-3

45.1 E-6

16

F IV, , vz] 540.55

2.5925

75.8E-6

46.3E-6

3.2E-6

6.3E-6

2.6E-9

17.4E-6

4.3537

2.6626

0.1859

0.3604

0.0001