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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
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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.
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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
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