Citation
Methods for promoting retinal ganglion cell neuroprotection and axon regeneration

Material Information

Title:
Methods for promoting retinal ganglion cell neuroprotection and axon regeneration
Creator:
Laughter, Melissa
Place of Publication:
Denver, CO
Publisher:
University of Colorado Denver
Publication Date:
Language:
English

Thesis/Dissertation Information

Degree:
Doctorate ( Doctor of philosophy)
Degree Grantor:
University of Colorado Denver
Degree Divisions:
College of Engineering, Design, and Computing, CU Denver
Degree Disciplines:
Bioengineering
Committee Chair:
Hunter, Kendall
Committee Members:
Park, Daewon
Ammar, David
Mestroni, Luisa
Soranno, Danielle

Notes

Abstract:
Retinal degenerations, such as glaucoma, affect millions of people worldwide and ultimately lead to death of retinal ganglion cells (RGCs) ultimately resulting in blindness. Unfortunately RGCs do not regenerate on their own making cell transplantation necessary to restore function and vision to these patients. Cell transplantation therapies to treat these diseases are extremely complex; to begin, the cells must first survive the implantation process in which they are injected into a diseased and damaged environment. Once implanted the replacement cells must then navigate their way into the ganglion cell layer and extend functional axons. During development RGC axon extension is guided by a series of neurotrophic factors and guidance cues; however, these factors are typically not expressed during adulthood. Considering the unfavorable environment and lack of an adult modality, successful cell transplantation would require a multifunctional environment to be implanted along with the cells to promote survival and implantation. Herein, we developed a reverse thermal gel system capable of encapsulating cells while being functionalized with a wide variety of moieties to cater to specific applications. We first functionalized this polymer with a cell-binding motif (RTG) intended to promote survival and axon extension of encapsulated cells. Furthermore we tailored this polymer system to possess mechanical and morphological properties similar to that of native retinal tissue. 3D culture and analysis of RGCs grown within this polymer system showed increased survival and axon extension when compared to controls. Next we conjugated the reverse thermal gel system system with charged sulfonate groups intended to promote prolonged release of NTFs from the polymer system. CNTF was loaded within the sulfonated polymer and injected into the vitreous of an optic nerve crush rat model. The sulfonated polymer system loaded with CNTF showed increased neuroprotection following optic nerve crush injury for up to 4 weeks when compared to the controls, indicating a prolonged release of CNTF from the polymer system. In the end, treatment for glaucoma will likely require a multifaceted approach due to the complexity of the disease state. However, in this work we were able to show successful in vivo and in vitro results through different manipulations of the same base reverse thermal gel system. Future work would involve combining these two functionalities in the same polymer for a multifaceted approach to the treatment of glaucoma.

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Auraria Library
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University of Colorado Denver
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Copyright Melissa Laughter. Permission granted to University of Colorado Denver to digitize and display this item for non-profit research and educational purposes. Any reuse of this item in excess of fair use or other copyright exemptions requires permission of the copyright holder.

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Full Text
METHODS FOR PROMOTING RETINAL GANGLION CELL NEUROPROTECTION AND
AXON REGENERATION by
MELISSA LAUGHTER B.S., University of Colorado Boulder, 2012 M.S., University of Colorado Denver, 2014
A thesis submitted to the Faculty of the Graduate School of the University of Colorado in partial fulfillment of the requirements for the degree of Doctor of Philosophy Bioengineering Program
2019


This thesis for the Doctor of Philosophy degree by Melissa Laughter has been approved for the Bioengineering Program by
Kendall Hunter, Chair Daewon Park, Advisor David Ammar Luisa Mestroni
Date: May 18th, 2019
Danielle Soranno


Laughter, Melissa (PhD, Bioengineering Program)
Engineering an Injectable Microenvironment to Promote Substantial RGC Survival and Axon Extension
Thesis directed by Associate Professor Daewon Park
ABSTRACT
Retinal degenerations, such as glaucoma, affect millions of people worldwide and ultimately lead to death of retinal ganglion cells (RGCs) ultimately resulting in blindness. Unfortunately RGCs do not regenerate on their own making cell transplantation necessary to restore function and vision to these patients. Cell transplantation therapies to treat these diseases are extremely complex; to begin, the cells must first survive the implantation process in which they are injected into a diseased and damaged environment. Once implanted the replacement cells must then navigate their way into the ganglion cell layer and extend functional axons. During development RGC axon extension is guided by a series of neurotrophic factors and guidance cues; however, these factors are typically not expressed during adulthood. Considering the unfavorable environment and lack of an adult modality, successful cell transplantation would require a multifunctional environment to be implanted along with the cells to promote survival and implantation. Herein, we developed a reverse thermal gel system capable of encapsulating cells while being functionalized with a wide variety of moieties to cater to specific applications. We first functionalized this polymer with a cell-binding motif (RTG) intended to promote survival and axon extension of encapsulated cells. Furthermore we tailored this polymer system to possess mechanical and morphological properties similar to that of native retinal tissue. 3D culture and analysis of RGCs grown within this polymer system showed increased survival and axon extension when compared to controls. Next we conjugated the reverse thermal gel system
m


system with charged sulfonate groups intended to promote prolonged release of NTFs from the polymer system. CNTF was loaded within the sulfonated polymer and injected into the vitreous of an optic nerve crush rat model. The sulfonated polymer system loaded with CNTF showed increased neuroprotection following optic nerve crush injury for up to 4 weeks when compared to the controls, indicating a prolonged release of CNTF from the polymer system. In the end, treatment for glaucoma will likely require a multifaceted approach due to the complexity of the disease state. However, in this work we were able to show successful in vivo and in vitro results through different manipulations of the same base reverse thermal gel system. Future work would involve combining these two functionalities in the same polymer for a multifaceted approach to the treatment of glaucoma.
The form and content of this abstract are approved. I recommend its publication.
Approved: Daewon Park
IV


ACKNOWLEDGEMENTS
I would like to thank all the people that contributed to the development of this work. I am extremely grateful to all my mentors, professors, family and friends that helped me achieve this goal. This experience has helped me not only define my interests and passions but has been amazing in terms of personal development.
First of all I would like to thank Dr. Daewon Park who has supported me throughout this whole project and from whom Fve learned a tremendous amount. I will be forever grateful for the opportunities you have given me and hope to continue to work with him in the future. At the same time, I would like to thank my committee members, Dr. David Ammar, Dr. Danielle Soranno, Dr. Luisa Mestroni, and Dr. Kendall Hunter for their encouragement and advice.
Similarly, I would like to thank Maria Bortot, Lindsay Hockensmith, Anna Laura Nelson, Madia Stein, James Bardill, and Adam Rocker for sharing this experience with me and always making lab a fun place to be.
I would also like to thank my friends in the department of ophthalmology for their support throughout my work. Their guidance and advice were an invaluable and essential part of this work.
Finally, I dedicate this thesis to my family, to my dad for being so inspiring and encouraging me everyday to have goals and motivate me to strive for greatness. To my mom who has been with me every step of the way and who has supported me with everything I do. To my sister, who I share every moment with and who has believed in me immensely. She has always given me the confidence to continue pursuing my dreams.
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TABLE OF CONTENTS
CHAPTER
I. INTRODUCTION......................................................2
II. SYNTHESIS AND CHARACTERIZATION OF A BIOMIMETIC REVERSE THERMAL GEL FUNCTIONALIZED WITH INTEGRIN BINDING RGD MOTIFS FOR 3-DIMENTIONAL RETINAL TISSUE
ENGINEERING.......................................................40
III. A SELF ASSEMBLING INJECTABLE BIOMIMETIC MICROENVIRONMENT
ENCOURAGES RETINAL GANGLION CELL AXON EXTENSION IN VITRO.............60
IV. AN INJECTABLE NEUROTROPHIC FACTOR DELIVERY SYSTEM SUPPORTS THE SURVIVAL AND REGNERATION FOLLOWING OPTIC
NERVE CRUSH INJURY................................................77
V. DISCUSSION LIMITATIONS AND FUTURE DIRECTIONS....................115
REFERENCES...........................................................129
APPENDIX
A. GPS analysis of PSHU........................................145
B. HPLC standard curve and quantification of RGD conjugation.......147
C. Loss modulus of RTG-RGD and RTG.............................148
D. Maximum intensity 3d fluorescent images of RGCs cultured in RTG.149
E. Synthetic route of RTG and SRTG.............................150
F. Characterization of RTG and SRTG............................151
G. Intravitreal injection control..............................154
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LIST OF FIGURES
FIGURE
1.1 Schematic showing the effects of rgd-integrin binding on cellular behavior.21
1.2 Polymer solubility behavior at the lower critical solution temperature.....24
2.1 Schematic synthesis of PSHU-PNIPAAm-RGD....................................41
2.2 NMR spectrum of PSHU.......................................................51
2.3 NMR spectrum of PSHU and deprotected PSHU..................................52
2.4 FT-IR of DPSHU, PSHU, and PSHU-RGD.........................................53
2.5 Temperature-dependent phase transition of PSHU-PNIPAAM and PSHU-
PNIPAAM-RGD ...................................................................55
2.6 Thermal gelling property of PSHU-NIPAAM-RGD ...............................56
2.7 SEM images of PSHU-PNIPAAM-RGD and PSHU-PNIPAAM............................57
2.8 Rheological properties of PSHU-PNIPAAM-RGD and PSHU-PNIPAAM................58
3.1 Schematic of 3D RGC culture process........................................67
3.2 Maximum intensity projections of RGCs cultured within 3D polymer scaffold....71
3.3 Live/dead staining of RGCs 3 days of culture on PDL-laminin glass coverslips.72
3.4 Quantification of average axon length for RGCs.............................73
4.1 Experimental protocol and animal groups....................................91
4.2 Expression of BRN3a in the retina following ONC............................94
4.3 Expression of BRN3a in retina following ONC and SRTG intravitreal..........99
4.4 Intravitreal treatment of SRTG-CNTF induced RGC axon growth...............104
4.5 Images of GAP-43 immunostained on downstream from crush site..............105
4.6 Response of Muller glial cells to ONC after intravitreal injection........107
4.7 Larger amounts of SRTG-CNTF system increased the survival of RGCs.........109
5.1 Subretinal injection of RTG mixed with GFP................................124
vii


5.2 Flatmount retinas stained with BRN3, B-Tubulin, and dapi................127
5.3 Cross-sections of retina showing integration of transplanted RGCS.......128
A. 1 GPC refractive index and light scattering curves for PSHU...............145
A. 2 GPC refractive index and log mw indicating polymer size separation......145
B. 1 HPLC curve of molar ratio of RGD/free amine group.......................147
C. 1 Rheological properties of PSHU-PNIPAAM-RGD and PSHU-PNIPAAM............148
D. 1 Maximum intensity projections of representative 3d fluorescent images of
RGCS cultured inside PSHU-PNIPAAM......................................149
E. l Schematic of synthesis process for SRTG.................................150
F. l Elemental analysis of SRTG and RTG......................................151
F.2 Temperature-dependent phase transition of SRTG and RTG..................152
F. 3. In vitro release profile of CNTF released from SRTG and RTG.............153
G. 1 Expression of BRN3A in retinal cross sections following ONC and SRTG...154
viii


ABBREVIATIONS
ANOVA analysis of variance
BDNF brain-derived neurotrophic factor
BRN3A brain-specific homeobox/pou domain protein
BSA bovine serum albumin
CNS central nervous system
CNTF ciliary neurotrophic factor
DMF dimethylformamide
DMSO dimethyl sulfoxide
ECM extracellular matrix
EDC //-(3-dimethyl ami nopropyl )-//-ethylcarbodii mi de hydrochloride
FT-IR fourier transform infrared
GFAP glial fibrillary acidic protein
GPC gel permeation chromatography
HDI hexamethylene diisocyanate
IOP intraocular pressure
LCST lower critical solution temperature
MMP matrix metalloproteinase
MW molecular weight
MWCO molecular weight cut-off
NHS //-hydroxy succini mi de
NMR nuclear magnetic resonance
NSC neural stem cells
NTF neurotrophic factor
ONC optic nerve crush


OPP
ocular perfusion pressure PBS phosphate-buffered saline
PNIPAAM poly (//-isopropyl acrylamide)
PSHU poly(serinol hexamethylene urea)
RGC retinal ganglion cells
RTG reverse thermal gel
SEM scanning electron microscopy
SRTG sul-pshu-pnipaam
x


CHAPTER I
INTRODUCTION
Overview
Glaucoma, amongst other retinal degenerations and optic neuropathies, is as it stands an incurable disease. Glaucoma refers to a group of diseases characterized by neuropathy of the optic nerve and loss of retinal ganglion cells (RGCs), the projection neurons located in the retina with axons extending through the optic nerve \ The death of these cells causes loss of a patient’s visual field and in the worst cases can leave a patient completely blind. The shock of losing vision can have devastating effects on a person’s quality of life and with Glaucoma being the second leading cause of blindness worldwide2 (second to cataracts), it is a problem in desperate need of a solution. This crux of this work focuses on taking key steps towards a cure for this disease.
There are many risk factors that increase the likelihood of developing glaucoma; however, increased intraocular pressure (IOP) is recognized as the principal risk factor3. This increase in IOP can be caused by excessive aqueous production, inadequate aqueous drainage, certain medications, trauma, and other eye conditions. In addition, factors such as race, age, and family history can play a role in increased IOP. Treatments are available to lower IOP (medications, laser surgery, microsurgery) and currently these are the only proven treatments for glaucoma. However, the enigma behind glaucoma is that although raised IOP remains the most reliable risk factor, many glaucoma patients develop this disease under normal IOP. Conversely, the majority of eyes with elevated IOPs do not develop glaucoma 4 Furthermore, a significant number of glaucoma patients see disease progression despite early diagnosis and successful lowering of IOP. This means that patients can suffer significant vision loss even when the appropriate treatments
2


are applied. In addition, there are some cases in which patients respond negatively to drugs or surgeries aimed at lowering IOP. In these cases, patients will see a more rapid disease progression and RGC loss 5. The bottom line is that many patients still suffer significant vision loss with current glaucoma treatments and the RGCs that are lost are unable to regenerate meaning that the vision lost is currently irreversible.
Pathophysiology of glaucoma
Glaucomatous optic neuropathy is characterized by damage to and subsequent loss of RGCs.
This damage typically occurs slowly, taking many years to develop. Although there is much debate as to the mechanism of primary and secondary insult, the end result is consistent, permanent loss of the neurons vital for sight. In this section, we will review the current topics in the pathophysiology of glaucoma, focusing on the steps leading to neuronal loss.
The pathophysiology of glaucoma is most likely multifactorial with various risk factors leading to damage of both the RGC soma and RGC axons. The complex nature of this disease has left many unknowns in the exact development of RGC death. Theories for primary injury in glaucoma begin with various risk factors including, elevated IOP and fluid dysregulation. Both of these factors can contribute to the initial insult seen in glaucoma, damage to the RGCs through variations in fluid flow and connective tissue. The secondary insult can be thought of as the damage resulting from injured neurons 6. Once damaged, injured RGCs release a variety of cytokines and reactive species leading to a vicious cycle of cell death.7
Following damage to RGCs in glaucoma, cell death occurs through the process of apoptosis 8 9. Apoptosis, also known as programmed cell death, is a mechanism of cell death in the absence of inflammation. This is because in apoptosis, cell death occurs without lysing of the cell and thus without the release of inflammatory factors. Instead apoptotic cells die through characteristic
3


DNA fragmentation, chromatin condensation, formation of apoptotic bodies, loss of mitochondrial membrane potential, and the breakdown of the cell into multiple smaller membrane-bound vesicles 6. These vesicles are then cleared or engulfed from the area by neighboring phagocytic cells. Although there is agreement that RGC death in glaucoma occurs mostly through apoptosis, there is some evidence that shows necrosis plays a part in the late stages of this disease 10.
The process of apoptosis is heavily regulated and progresses in an orderly fashion through a series of signals. This signally cascade is modulated by caspases, a family of aspartate specific cysteine proteases and members of the interleukin-1 B-converting enzyme family. Due to the destructive capabilities of these enzymes, caspases initially exist as inactive zymogens. Once cleaved, they work to initiate, propagate, and amplify the apoptotic process. There are three major pathways of apoptotic progression in mammals: the death receptor pathway (extrinsic), the apoptosome pathway (intrinsic), and the cytotoxic lymphocyte-initiated granzyme B pathway u. The extrinsic pathway is initiated when the Fas ligand (FasL) binds to the Fas receptor (death receptor) present on the cell’s outer membrane. This binding triggers a signally cascade ultimately leading to the activation of caspase-8 and caspase-10. These downstream caspases work to digest the cellular contents and begin apoptosis n. There is evidence that in humans, glaucoma progresses mostly through the extrinsic pathway. An increase in the expression of FasL and Fas receptor are seen throughout glaucoma progression and are also associated with increased IOP (a strong risk factor for glaucoma)1314 Elevated IOP can also cause an increase in TNF-alpha lending to further proof that the extrinsic pathway of apoptosis is highly involved in glaucoma. Due to this, researchers have begun to develop methods to intercept this pathway. Specifically antibodies against TNF-alpha have been shown to block some of the subsequent
4


RGC damage in mice with ocular hypertension 15. In addition, one study delivered caspase inhibitors to rats that had undergone axotomy of the optic nerve. This treatment recused 34% of RGCs that would have otherwise died 14 days post injury 16.
Elevated intraocular pressure as a risk factor in glaucoma
RGC death in glaucoma commonly occurs in the presence of increased IOP, making this a very important risk factor for disease progression 17 Several studies have shown the association between elevated IOP and RGC loss. More specifically, investigators have shown a positive correlation between a change in IOP and RGC death. Furthermore, it has been shown that the duration of increase in IOP can determine the significance of RGC death 18. However, among glaucoma patients, only one-third of patients have increased intraocular pressure at the early stages. Furthermore 30-40% of patients experiencing visual field loss due to glaucoma, have an intraocular pressure within normal range 19. Thus, while increased IOP is still considered an important risk factor, it is by no means the only player. Three factors that determine IOP are:
1. The degree of venous pressure
2. The rate of aqueous production by cilliary bodies, and
3. The rate of aqueous drainage across the trabecular meshwork-schlemm’s canal system. Generally elevated IOP results from the third mechanism, impaired drainage of aqueous humor due to changes in the trabecular meshwork. Most of the aqueous humor in the eye drains through the trabecular meshwork into the schlemm’s canal and empties into the collecting veins on the scleral surface. The trabecular meshwork is made of layers with one layer composed of connective tissue and the other of endothelial cells. As we age, the endothelial layer begins to diminish and the connective tissue layer begins to thicken. Thickening of the connective tissue increases resistance and impedes aqueous drainage thereby prompting an increase in IOP. By
5


definition, elevated resistance to outflow at the level of the trabecular meshwork is primary open-angle glaucoma where the iridocorneal angle remains anatomically open. On the contrary, angle-closure glaucoma occurs as a result of elevated resistance building between the front of the lens and the pupillary margin. This increase in pressure pushes the iris up blocking the iridocorneal angle and aqueous drainage.
Mechanisms of retinal ganglion cell apoptosis in response to increased IOP
Unfortunately, it is still not well understood how this increase in IOP leads to RGC apoptosis. One proposed mechanism is that increased IOP alters the cellular ECM through extensive collagen remodeling and the presence of matrix metalloproteinase (MMP)-l . MMPs are the enzymes responsible for the degradation of ECM. In a recent study, the exposure of eyes to increased IOP caused an increase in the release of MMPs into the ECM. Digestion and damage to the ECM by these MMPs can lead to disruption of RGCs and subsequent cell death. It is known that cell-ECM and cell-biomolecule interactions are crucial for the survival of most cells. Specifically integrin binding motifs present in many components of the ECM, are involved in supporting attachment, spreading, and survival of neural cell types 23. Thus, the degradation of the ECM caused by the release of digesting MMPs, could compromise these crucial cellular attachments leading to RGC apoptosis. The second explanation is that elevated IOP increases the mechanical force exerted on RGC axons passing through the lamina cribrosa of the optic nerve. Damaged RGCs begin to secrete MMPs, which in turn causes digestion and damage to the ECM further perpetuating this disease state 4 Finally, the damage of RGC axons can hinder retrograde transport of critical growth factors. When these growth factors are no longer able to make it through the RGC axon and up to the cell body, the cell may no longer be able to regulate
6


metabolism and cell survival leading to apoptosis (this will be discussed further in subsequent sections)24.
Vascular insufficiency
Undoubtedly, elevated IOP likely plays a huge role in the development of glaucoma. However, this theory does not account for the large number of patients that develop glaucoma under the normal range of IOP. In addition one study reported that over 90% of people that had high IOP failed to develop glaucoma over a 5-year period. Similarly, there are many patients that develop glaucomatous neuropathy with a normal IOP. It was also noted that patients that have glaucoma and increased IOP continued to show disease progression even when their IOP was reduced to a normal range 25. Thus, there are seemingly other risk factors involved that contribute to the development of this disease. Recent works point towards a correlation between the presence of vascular insufficiency and glaucoma. Vascular insufficiency includes reduced ocular perfusion pressure and defective vascular regulation.
Ocular perfusion pressure (OPP) describes the relationship between blood pressure (BP) and IOP as shown in the following equation:
OPP = BP - IOP
Studies have shown that a decrease in OPP is a risk factor for glaucoma and that systemic hypertension may in fact be protective against the development of glaucoma (likely because this would increase the OPP)26. Defective vascular regulation can also contribute to the formation of glaucoma. Constant blood flow is necessary to provide oxygen and nutrients to the eye, most importantly the retina and optic nerve. In a healthy eye, the maintenance of constant blood flow is controlled through autoregulatory mechanisms existing within the blood vessels. Unfortunately, as you age, these mechanisms worsen and the ability to adapt to fluctuations in
7


pressure decreases. In one study, researchers showed that older rats were unable to adapt to changes in ocular pressure as compared to younger rats 27. Thus, these defects in autoregulation that occur with age could lead to ischemia within the eye and damage to the RGCs.
Treatments for glaucoma
The first question to ask when devising a treatment for glaucoma is, how do RGCs die in glaucoma? It is now agreed that RGC death in glaucoma occurs via apoptosis or necrosis28 (as discussed above). Apoptosis is the active process of programmed cell death. This process may initiate when the cell is deemed no longer necessary or becomes damaged in some way 29. On the contrary, RGC necrosis occurs when surrounding toxins injure the cellular membrane and destroy the cell. This process is passive in nature and can lead to an inflammatory response, which can contribute to more cell death. Although both methods of cell death can occur during glaucoma, apoptosis is known to be the primary method of cell death with necrosis occurring during the late phase of the disease30.
New therapeutic approaches
Until recently, most therapeutic approaches for glaucoma were focused on lowering IOP. However, as mentioned above, a large portion of these patients continue to show disease progression even when their IOP is brought down to normal conditions. In addition, there is a significant portion of patients that develop glaucoma under normal IOP ranges. As new information on the pathophysiology of glaucoma began to emerge, new treatments were developed that focused on these mechanisms instead of solely lowering the IOP. One new group of therapeutic agents are aimed at preventing apoptosis of the RGCs. As discussed above, the process of apoptosis is mediated through enzymatic proteins called caspases. Caspase inhibitors, such as erythropoietin, have shown efficacy in preventing apoptosis in a glaucoma rat model31.
8


However, apoptosis is the final step in the cellular progression of glaucoma and thus these treatments act far downstream to where the problems is initiated. Other agents, such as N-acetyl-L-cysteine, NOS inhibitors, and N-methyl-D-aspartate (NMDA) antagonists, have shown promising results at preventing RGC death in glaucoma animal models 32 16. NMDA antagonists showed a lot of potential at neuroprotecting RGCs in animal models and were subsequently taken to clinical trials. NMDA is a glutamate receptor that, once activated, leads to the activation of ion channels and the influx of Ca2+ and Na+ into the neuron. The influx of these extracellular ions works to trigger a signaling cascade that leads to cell death 33 34. Thus, NMDA antagonists prevent the activation of the NMDA receptors and the resulting cell death. However, once in clinical trials, it was quickly noted that these antagonists not only block the NMDA receptors in the eye but all NMDA receptors essential for neuronal function throughout the body 35. Neurotrophic factor treatments for glaucoma
The second question to ask when investigating RGC death as a result of glaucoma is, what causes the RGCs to die? The answer to this question is unfortunately much more complex, with likely several molecular pathways contributing to RGC cell loss. However, recent progress in this field has increased our understanding of the pathways that lead to RGC degeneration following optic nerve injury. The answer to this question is now believed to lie with neurotrophic factors (NTF) and their role in a healthy retina. Neurotrophins are a family of small diffusible molecules that are implicated in the control of adult neurogenesis, axon extension, proliferation, and cellular survival36. Two NF exhibiting key roles in the survival of RGCs are brain-derived neurotrophic factor (BDNF) and ciliary neurotrophic factor (CNTF). In healthy eyes, BDNF is produced in the lateral geniculate body where it binds to its receptor on the RGC axon and moves to the RGD cell body through retrograde transport37 When BDNF reaches the RGC cell
9


body, it binds to its corresponding receptors to affect survival of RGCs 24. On the contrary,
CNTF is released as a response to diseased conditions. Early studies first identified increased CNTF following injury of retinal neurons, after axotomy, ischemia, and experimental glaucoma 38 39. However, CNTF works to provide a similar result to BDNF by inducing neuroprotection of the RGCs.
As mentioned above RGC axons are responsible for transmitting signals from the photoreceptors to the cortex of the brain; however, RGC axons are also responsible for the uptake and retrograde transport of distant NTF to the cell body \ Although cells within the retina also produce NTF, studies have shown that both sources of NTF may be important in preserving the survival and function of RGCs 40. The hypothesis that RGC death occurs due to ‘axonal transport failure’ derives from this knowledge. As glaucoma proceeds, RGC axons get damaged decreasing axonal transport of vital NTFs. Insufficient or unbalanced levels of NTF can cause RGCs to go into an apoptotic cascade. Evidence for this hypothesis stems from studies showing decreased in retrograde axonal transport following glaucoma-like optic nerve damage 41 42. Additionally, increased IOP and subsequent optic nerve damage has been correlated with retinal accumulation of dynein, a motor protein required for retrograde axonal transport 43. Although the ‘axonal transport failure’ hypothesis and ‘neurotrophic factor deprivation’ hypothesis have not been completely proven, significant evidence points to these reasons for the loss of RGCs in glaucoma. Neuroprotection of retinal ganglion cells
Recent discoveries in the mechanisms behind RGC death in glaucoma have led to new potential treatments. The first line of treatment for glaucoma have been aimed at neuroprotection of RGCs. Neuroprotection is a therapeutic strategy aimed at preventing RGC death and loss of function 18.
10


Instead of traditional preventative measures to treat glaucoma, such as decreasing IOP, neuroprotective strategies focus on guarding RGC survival and function.
Neurotrophin supplementation strategies have been one area of research extensively studied for RGC neuroprotection. BDNF and CNTF have been the primary NTFs used for these neuroprotection strategies. Initial data supporting these therapies have involved optic nerve injury models (e.g. optic nerve crush or transection model) 18. Although optic nerve crush (ONC) in a rodent shows a different progression from glaucoma in a human, it does provide researchers with the ability to study RGC death during optic neuropathy. In addition, this animal model has been able to provide an effective model for optic nerve trauma as well as regeneration failure. Furthermore, the optic nerve is highly accessible increasing the consistency of the model and making it an effective tool to investigate new treatments. Studies have been able to show repeatable results in RGC survival following nerve crush with a predictable number of RGCs remaining at various time points. For instance, during the first week following optic nerve crush there is a small difference in RGC cell count (depending on distance between crush site and optic cup) 44 After the first week post nerve crush, changes in RGCs occur much more rapidly. A significant number of RGCs die after the first week reducing the number of RGCs in the retinal layer to 10%. This abrupt cell death allows researchers to see differences in RGC survival between treatment and control groups. For instance, studies using an optic nerve crush or transection animal model followed by intraocular injection of BDNF or viral-mediated BDNF gene transfer using adenovirus have been able to show an increase in RGC survival at just one week post lesion (100% survival of RGCs in treatment group compared to 50% in control groups). Two weeks post lesion, similar treatments have been able to rescue 48% of RGCs compared to the 10% survival rate in the control eyes 45 46 41. Although BDNF does show a
11


profound effect on RGC survival following injury, it unfortunately has little effect on regenerating axons. Previous studies have shown that intravitreal injection of BDNF following optic nerve crush did not increase RGC axon regeneration past the lesion site. In fact, it is now believed that BDNF may have an inhibitory effect on RGC axon regeneration 48 49. Pernet et al showed that not only did the application of endogenous BDNF fail to stimulate axon regeneration, but it also led to hypertrophic axonal swelling potentially causing further damage to segments of the optic nerve 48. Furthermore, neuroprotective strategies using BDNF (either via recombinant protein or gene transfer) have failed to extend the time-course of RGC survival. Groups have attempted repeated intravitreal injections of BDNF as well as mini pumps to provide sustained NTF delivery, but have not seen RGC survival past certain time points 50. This result stems from decreased expression of the BDNF receptor (TrkB) following RGC injury. The decreased amounts of the TrkB receptor from cellular internalization limits the ability of RGCs to respond to neurotrophin stimulation 51 45.
Ciliary neurotrophic factor
Similar to BDNF, CNTF has a neuroprotective effect on RGCs following optic nerve injury. Administration of exogenous CNTF following optic nerve axonomy displayed increased survival of RGCs although not as increased as BDNF 52. However, therapies with CNTF using gene transfer appear to be more effective at preserving RGC survival post injury. These studies showed a modest 25% to 50% survival of RGCs (compared to 10% control eyes) two-weeks post lesion 53. Although CNTF doesn’t have as great an effect on RGC survival as BNDF, CNTF has been shown to stimulated RGC axon regeneration post injury. Studies using AAV-mediated CNTF gene transfer following optic nerve crush stimulated significant RGC axon regeneration 54 This information indicates that BNDF and CNTF operate using different signaling pathways and
12


therefore stimulated different cellular responses such that BDNF is more effective at stimulated neuroprotection and CNTF is more effective at stimulating axon growth.
Neuroprotective strategies involving NTF can be divided into three categories, ones that supply RGCs with exogenous NTFs, ones that use genetic modifications to stimulate the cellular expression of NTFs, and finally ones that use genetic modifications to increase the expression of NTF receptors (such as TrkB). As mentioned above, NTF receptors can become downregulated following optic nerve lesion 51. This can lead to RGC death by making the cells less responsive to NTF survival signaling. Different strategies have been used to upregulate these receptors in hopes of increasing the potency of surrounding NFTs and thus promoting RGC cell survival55. Most strategies use vectors to transfer genetic information coding various receptors. Additionally, strategies have been investigated to increase endogenous expression of NTF from the RGCs themselves. Similar to strategies used to increase receptor expression, these strategies include viral vector transfection of genes coding NTF into RGCs 56. Although promising results have been seen in optic nerve crush animal models, both of these strategies include variable genetic alterations that may be too complicated for current clinical applications. Lastly, direct injection into the vitreous has been used to deliver NTFs to damaged RGCs. As mentioned above, these strategies have shown prolonged survival of RGCs as well as some axon regeneration after optic nerve crush or axonomy 57 58. Although this approach is minimally invasive and does not require any genetic manipulations, this method prompts diffusion of the protein out of the intended treatment area and quick proteolytic degradation. Therefore, applying this method in the clinical setting would require multiple and frequent doses of the GF but would still not likely result in the full effect59.
13


Cellular scaffolds for retinal ganglion cell regeneration
As it stands, RGC loss from any optic neuropathy is irreversible. This means that for many patients suffering from glaucoma, the loss of one’s visual field is permanent and cannot be improved. The only treatments that are available are preventative measures aimed at reducing IOP and protecting surviving cells from death. For many individuals, these preventative treatments are not sufficient and do nothing to improve their quality of life. For these cases the only hope of restoring vision is optic nerve regeneration; however, this goal presents immense challenge. Firstly stem cells induced to differentiate into RGCs must be sourced to replace the depleted cells. It is important that these cells are delivered to the appropriate area and integrate successfully into the surrounding tissue. Unfortunately, transplantations consisting solely of cells will be hindered by a diseased microenvironment characterized by growth-inhibitory signals. In glaucoma, the diseased state consists of myelin associate molecules, scar formation, and lack of passage across a lesion, all which need to be mitigated to allow successful cell transplantation60. The environment of transplantation needs to be made permissive to cell survival and axon regeneration.
Peripheral nerve grafts
Peripheral nerve (PN) grafts have been used to promote survival of injured RGCs (when transplanted into the vitreous) or to promote RGC axon regeneration through the lesion (when grafted to the optic nerve). Studies have shown that implanted fragments of PN into the vitreous can stimulate RGCs to regenerate axons across the lesion site 61. However, these results are mostly attributed to the release of NTF from the Schwann cells and the presence of macrophages that can enhance axon regeneration. Furthermore, the use of ‘conditioned’ or pre-crushed grafts increase RGC survival significantly more than normal grafts due to the increase in release of
14


NTF 62. PNs can also be grafted to the end of a transected optic nerve 63. Studies have shown that RGCs are capable of regenerating axons through the grafted PN extending all the way to the superior colliculus and forming correct synaptic connections 64 65. The ability of RGCs to regenerate axons through the grafted PN is likely due to the increased amount of cell adhesion molecules and decreased amount of growth-inhibitory molecules present within the graft66.
Stem cell replacement therapies
Although regenerating the optic nerve seems like a very distant goal, stem cells and tissue engineering hold great potential for achieving this goal. As mentioned above, the loss of RGCs due to glaucoma or other optic neuropathies are currently permanent. Therefore, it is necessary to treat such diseases with cellular replacement therapies to replace the depleted RGCs and possibly reverse the vision lost. Stem cells have garnered significant attention due to their ability to self-renew and differentiate into multiple cell types, making them capable of replacing specific types of tissue. For this work, we are concerned with the differentiation of stem or precursor cells into RGC-like cells. Preliminary studies have investigated whether implanted cells can differentiate into an RGC-like phenotype 67 68. To do this, groups have used histological markers specific for RGCs to determine effective differentiation of the stem cells. Two of the most common immune markers are B-3 tubulin and brain-specific homeobox/POU domain protein 3 (Brn3). Furthermore, studies have been conducted to determine the best place for cell transplantation.
The first and most simple injection site investigated was the vitreous cavity. Studies have shown variable results depending on the type of stem cell transplanted. One study showed that neural stem cells (NSCs) injected into the vitreous were unable to integrate into the retinal layer and efficiently differentiate into RGCs 69. However, other groups have demonstrated that retinal stem cells (RSCs) are more capable of incorporating into the retinal layer following intravitreal
15


injection. Regardless, few stem and precursor cell types injected intravitreally differentiate efficiently into RGCs and incorporate into the retinal layer 70. Transplantation can also be done using subretinal implantation. Such studies have shown that RSCs integrate better when implanted into the subretinal layer when compared to intravitreal injection 71.
However promising these early results are, cell transplant therapies are still plagued by a low rate of cell survival and limited integration of cells into target tissue. These results are likely caused by the injured microenvironment the cells are exposed to upon injection. The microenvironment or ‘niche’ can have a severe effect on cellular behavior by exposing cells to various spatially and temporally controlled biochemical cues, cellular attachment ligands, and extracellular matrix (ECM) molecules. In addition, even the microenvironment’s mechanical properties and topographical cues can illicit specific cellular behavior. Accordingly, cell implantation into an injured microenviroment that has been compromised by growth inhibiting myelin associated molecules, absence of growth-permissive molecules (e.g, laminin), scar formation, and lack of passage across a lesion is extremely detrimental to implanted cells 12. One way to mitigate these limitations is through the use of biomaterials to act as two-dimensional 2D or three-dimensional 3D cellular scaffold. These scaffolds not only provide physical support for the implanted cells but also can be engineered to possess specific mechanical properties and biochemical cues. Therefore, these biomaterials provide the exciting possibility of constructing an alternative microenvironment that can instead encourages cell survival, growth, and in our case axon extension.
Biomaterials for controlled delivery of alternative microenvironment
Not only can biomaterials be used to aid in the delivery of replacement cells, but they can also be injected into a diseased environment to provide an alternate bioactive niche. In this case the
16


biomaterial would be used to alter the disease microenvironment to augment endogenous stem cell activity and subsequent regeneration. This can be achieved by incorporating biomolecules within the scaffold that can be released into the surrounding area or through the chemical conjugation of biomolecules. Both methods can lead to activation of endogenous stem cell populations and potentially to the self-renewal and differentiation into a replacement cell population. Studies with transplantation of melanoma cell-adhesion molecules (MCAM) have been able to show the activation of endogenous human bone-marrow stroma and subsequent differentiation into bone cells 73. Furthermore, in previous studies we have shown that a novel synthetic biomimetic polymer (PSHU-RGD) can induce significant differentiation of human neural stem cells (hNSCs) into a population of human motor neurons (hMNs) in 2D.
Designing biomaterials for a 3D scaffold/microenvironment Analysis of cellular behavior on 2D materials has been the first step to engineering an appropriate cellular scaffold. 2D materials allowed for controlled experimentation to determine how cells interact with individual components of a cellular niche 23. However, 2D approaches are limited to preliminary in vitro studies in that they are unable to represent natural tissue. Unlike 2D systems, 3D materials are more in tune with the cell-cell interactions, cellular organization, and microenvironment seen in native tissue. In 3D systems, cells will be fully encapsulated by a solid microenvironment and thus be more likely to respond to the scaffold specifically engineered to induce positive cellular behavior.
Fabricating a 3D scaffold is much more than just creating a 3D structure. Instead, there are many considerations depending on the desired cell behavior. Each characteristic and component of the 3D scaffold will have some sort of effect on cellular behavior so each property must be considered and engineered appropriately. Not only do 3D scaffolds require the same
17


considerations as 2D scaffolds (e.g. biomolecule density, material wettability, material charge) but also 3D scaffolds bring new complications such as mechanical properties (elasticity) and morphology (porosity).
In addition, practical considerations need to be taken when designing a 3D scaffold. First, the biomaterial must allow embedding of cells and ideally encapsulate the cells in situ. Typically, this is done through crosslinking of the material after implantation. However, this has to be well controlled to ensure harmful side products produced from the reaction do not damage the encapsulated cells or the surrounding tissue. Second, how the biomaterials will be delivered to the diseased area is another important consideration. Perhaps the ultimate goal is to create an injectable material that allows facile encapsulation of cells and gelation in situ. In addition, this characteristic would provide a more minimally invasive method less likely to further damage already damaged and possibly necrotic tissue present in a diseased state 74.
Biomolecule modification of biomaterials and cell behavior Surface modifications of biomaterials provide researchers with a way to fine-tune and manipulate materials to possess certain characteristics. One of the main goals of surface modifications is to make a material biomimetic. Biomimetic polymers combine the controllability of a synthetic polymer with the biocompatibility of natural material found in the body. Biomimetic polymers can be manipulated to prompt a specific cellular response or direct new tissue formation. Extensive work has been conducted to create biomimetic polymers suitable for cell transplantation therapies. One of the most prominent methods to create a biomimetic material is through surface modification of common ECM proteins such as fibronectin (FN), vitronectin (VN), and laminin (LN) 75. However, it was discovered that a small sequence of amino acids within the larger proteins is responsible for integrin interaction and
18


cellular attachment. It is now more common to simply incorporate the short peptide fragments to the material surface. This is because the larger proteins can present steric hindrance that may prevent cellular attachment. In addition, using smaller peptides can increase the amount of biomolecule that can be attached to the polymer surface thus increasing the cellular response. Finally, small peptide sequences are relatively more stable than the full protein and can thus better withstand the chemical conjugation process.
Arg-Gly-Asp (RGD) is the cell-binding motif found in fibronectin (FN) and laminin (LN). Because of its extensive presence and its highly effective cell binding capacity, RGD has become the most commonly employed peptide used to enhance cell attachment to a polymer’s surface 76. RGD is capable of binding to a plethora of cell adhesion receptors found throughout the body; however, the largest and most common group of RGD-binding cell adhesion receptors is the integrin family. This group of adhesion molecules is not only responsible for cell attachment, but they also play an important role in cell differentiation, embryogenesis, proliferation, and gene expression77 Binding of the RGD peptide to integrins causes a cascade of signals that can influence
cell behavior in different ways (Figure 1.1) 76. Research has shown that focal adhesion-associated proteins trigger the expression of anti-apoptotic protein Bcl-2 and the focal adhesion kinase pathway 78. These two pathways along with focal adhesion formation play an important role in cell survival, proliferation, cell spreading, and neurite extension23. It was also determined that the lack of this interaction between integrins and cell-binding motifs can induce apoptosis and cell death. This ‘integrin-mediated death’ is caused by the lack of a substrate to bind the ligand leading to the recruitment of caspases and the cell death cascade. Relating this finding to a
19


cell’s interaction with the ECM, we can see that the absence of integrin-binding ligands in a cell’s environment can cause the cell to undergo apoptosis 79.
Mechanical properties of biomaterials and cell behavior
In glaucoma, the diseased tissue is not only associated with cell death and release of toxic cytokines, but also with altered organization of the ECM. This disturbance in the tissue can lead to changes in the mechanical properties or stiffness of the matrix, which most likely contributes to the progression of the disease. It was recently discovered that the mechanical properties of the ECM, in which cells reside, can have an extreme effect on cell behavior and cell fate. In one of the first studies displaying this finding, human mesenchymal stem cells grown on polyacrylamide with varying stiffness (typical of brain, muscle, and bone) began to differentiate into the corresponding tissue-specific cells 80. This innovative study highlighted the strong
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Figure 1.1. Schematic showing the effects of RGD-integrin binding on cellular behavior. Integrin binding can induce cell survival, proliferation, and axon extension.
20


influence scaffold mechanical properties can have on cells and marked the beginning of many future studies relating matrix stiffness and cell behavior. Following suit, other groups were able to show that mechanical properties have significant control over proliferation and differentiation of other types of stem cells. One such study showed that adult neural stem cells (NSCs) grown on relatively stiff substrates triggered differentiation into glial cells, whereas adult NSCs grown on softer substrates (more closely resembling the stiffness of brain tissue) triggered differentiation into mostly neurons 81.
Compared to these 2D studies, examining mechanical properties in a 3D scaffold is significantly more challenging. This is due to the presence of other factors (physical, chemical, and mechanical) that are in play in a 3D matrix. Therefore, constructing a suitable 3D scaffold with the appropriate characteristics is complex. With this said, substrate elasticity and materials with mechanical properties closest to that of the native tissue are the most desirable and will likely have the most beneficial effect on surrounding cells 82. For our purposes, it was important to create a material with similar mechanical properties to native retinal tissue, as this would likely improve RGC culture in vitro as well as in vivo results. Rheology studies have been done to assess the mechanical properties of retinal tissue. Results from these experiments show that native retinal tissue is soft with a storage modulus not exceeding 100 Pa 83. Thus, a matrix that possesses softer mechanical properties and a modulus that does not exceed a couple hundred Pascals would be beneficial to RGC growth and survival.
In situ polymer gelling systems
Perhaps the ultimate goal of cellular scaffolds is to create a multicomponent, injectable material that can conform to the target tissue and provide a replacement niche. This characteristic is especially important when it comes to engineering neural tissue. As stated above, it is important
21


that scaffolds be tailored to match the mechanical properties of the native tissue. However, neural tissue has an extremely low mechanical modulus making implantation of a scaffold nearly impossible. To combat this obstacle, researchers have turned to in situ forming polymers. These polymer systems work by transitioning from a low-viscosity fluid to a solid gel upon application of various stimuli. This characteristic allows the polymer system to be deployed through a minimally invasive injection into the target site and conform to the site upon gelation. Not only does this allow for the use of polymer systems with an extremely low modulus, but it also allows the entrapment of molecules or cells before application. Initially, researchers used polymer systems that relied on photo-initiated free radial formation to provide this in-situ gelling behavior. However, the side products of this reaction proved cytotoxic to cells and made this technique no longer feasible for tissue engineering or drug delivery application 84
Avoiding polymer systems that produce harmful side products upon gelation, researchers turned to stimuli-sensitive polymers that undergo reversible volume phase-transition or sol-gel phase-transition in response to various stimuli. These stimuli might include external physical or chemical stimuli such as pH, charge, light, or temperature. Unlike in-situ gelling polymers formed by harmful chemical crosslinking reactions, stimuli-sensitive polymers can be reversibly transformed from solution to solid state by altering environmental conditions. Therefore, these stimuli-sensitive polymer systems may be polymerized and purified before application providing a simpler and safer method for injectable materials.
Thermally induced gelling systems
Among stimuli-sensitive polymer systems, temperature sensitive polymers (thermogels) are the most widely used due to their facile control and practical advantage for both in vitro and in vivo purposes 85. These systems work through an entropically driven phase transition upon reaching a
22


specific temperature, also referred to as the lower critical solution temperature (LCST). The LCST of a polymer system is determined by competing hydrophobic and hydrophilic interactions with the polymer backbone. More specifically, below the polymer’s LCST hydrophilic interactions between the backbone and water molecules prevail causing the polymer to go into solution. Above the
polymer’s LCST, the hydrogen interactions become weaker and hydrophobic interactions between the polymer backbone moieties prevail, causing the water to be expelled from the
Hydrophilic
Hydrophobic
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loss of bound water
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Figure 1.2. Polymer solubility behavior at the Lower Critical Solution Temperature (LCST). Left hand side shows hydrated polymer below LCST with entropic loss of water and right side shows chain collapse above LCST 208.
polymer and form a solid gel. A Schematic of this phenomenon can be seen in Figure 1.2. Because this process is controlled through hydrophilic and hydrophobic interactions and not through polymerization, this process is fully reversible. For this reason, these temperature-sensitive systems are termed reverse thermal gels (RTGs). The general principle behind RTGs, is to construct a block copolymer that has both hydrophobic and hydrophilic segments, named an ampiphilic polymer. This study focuses on one ampiphilic polymer specifically, poly (N-isopropylacrylamide) (PNIPAAm).
23


PNIPAam is a thermosensitive water-soluble homopolymer that has garnered a lot of attention in the biomedical field. As described above, this ampiphilic polymer exhibits a reversible hydration-dehydration solution to gel (sol-gel) change in response to temperature 86. In addition, this polymer has been shown to exhibit a sharp, reversible sol-gel transition sufficiently above room temperature (25°C) and sufficiently below body temperature (32°C) 87 making it extremely useful for many biomedical applications 88. When conjugated with other polymers, PNIPAAm can impart its thermosensitive behavior on otherwise non-stimuli sensitive polymers or molecules. For these reasons, PNIPAAm has been employed for many different applications including drug delivery, tissue engineering and, scaffolding applications 89 90 91.
From a drug delivery perspective, PNIPAAm offers significant advantage in that it shows good sensitivity, reversible transition, and low cytotoxicity when compared with other reversible thermal gels 92. In this sense, PNIPAAm provides a favorable environment for the entrapment of certain drugs and molecules. In addition, these molecules can be loaded into the polymer system through facile mixing below room temperature and delivered to the site through a minimally invasive injection, encapsulating the loaded molecules and conforming to the injury site upon gelation. Furthermore, the degradation of this polymer can be altered through changing the chemical structure of PNIPAAm, providing control of the release profile of entrapped molecules 92. Snowden et al first demonstrated the potential of PNIPAAm as a drug delivery system through investigating the release of fluorescein-labeled dextran from the PNIPAAm-based microgels 93. Researchers then moved to show the sustained delivery of other molecules such as insulin 94 and bovine serum albumin (BSA) 95 from PNIPAAm systems. However, studies soon found that PNIPAAm, used as a drug delivery system, is limited by a quick burst release. In one case encapsulated drugs were released from the polymer within the first 24 hrs of deployment96.
24


Groups such as Zhang et al. mitigated this obstacle through manipulating the PNIPAAm based systems. They were able to show controlled release of water-soluble protein BSA for an extended period of time without a significant burst release. This study gave rise to further studies using PNIPAAm as a drug delivery system for similar water-soluble growth factors and neurotrophic factors. Some groups showed the benefits of prolonged release of neurotrophic factors from a PNIPAAm drug delivery system in the regeneration of motor neurons and treatment of spinal cord injury (SCI). One particular study showed that six weeks of release of BDNF from a PNIPAAm-PEG injectable scaffold encouraged survival and attachment of motor
90
neurons
PNIPAAm has also been used for many tissue-engineering efforts. Typically polymer scaffolds used for tissue regeneration have been comprised of biodegradable polymer systems such as poly (lactic acid) PLA, PLA-based copolymers, alginate or collagen. Unfortunately these polymers degrade quickly and thus may not provide sustained mechanical support97 Disappearance of the scaffold before injured axons are able to regenerate can halt regeneration across the lesion as well as cause an increase in inflammatory response and glial scar formation. Furthermore, the polymer systems listed above all require surgical implantation. This process can lead to significant damage of an already damaged site. For these reasons, researchers have begun to investigate injectable scaffolds that can conform to the irregular shaped injury site and be deployed through facile injection. Unfortunately most of these polymer systems require crosslinking following injection and can release harmful unreacted chemicals into the surrounding tissue. PNIPAAm was the clear alternative to other injectable polymer systems in that it does not chemically crosslink upon injection. Furthermore PNIPAAm degrades slowly and can thus provide prolonged support for encapsulated cells once implanted 98 99.
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Biomimetic scaffolds for retinal tissue engineering
In the final two sections of this introduction, we will discuss the use of injectable materials in both retinal tissue engineering as well as neurotrophic factors/drug delivery for retinal regeneration applications. These sections will include the current state of research, notable pitfalls, and an introduction to the research performed in the subsequent chapters. We will focus this discussion on the use of synthetic polymers.
With the recent advances in stem cell technology, cell-based therapies have an increased potential to provide curative treatments for a wide variety of diseases. For retinal tissue engineering, transplantation of NSCs, RGCs, or RGC progenitors are of particular interest.
Solely the delivery of cell suspensions into the vitreal space have been shown to provide some amount of neuroprotection in neurodegenerative disease models 100101, likely due to the release of NTF from the implanted cells 102. However, the integration of cells injected in this fashion is erratic with most cells incorporating randomly within the target tissue. Thus, the efficacy of transplantation as a treatment is hindered by poor survival and improper integration of freely injected cells into the damaged tissue 103. A suitable vehicle to deliver replacement cells to the target tissue can mitigate these issues by providing appropriate biochemical cues and mechanical support to the encapsulated cells. However, traditional cellular scaffolds require surgical implantation, which can incur further damage to already, damaged tissue 104 For this reason, attention has turned to injectable materials that avoid surgical implantation while retaining all the benefits of a cellular scaffold. Various forms of ECM technology have emerged capable of delivering transplanted cells in an injectable scaffold. One commonly used platform for neural tissue engineering has been injectable hydrogels. These materials are highly tunable, capable of providing various mechanical and biochemical cues to encourage cell survival, deliver drugs, or
26


provide specific mechanical support. Although promising, many of the injectable hydrogel scaffolds require in vivo polymerization giving way to potentially harmful side products that may negatively affect implanted cells and surrounding tissue. Furthermore, in situ polymerization limits the controllability of each polymerization reaction because the body can be an extremely variable environment. Polymers that remain injectable but can be polymerized prior to injection would provide more control and limit any unwanted side effects.
In addition, researchers have focused on creating biomimetic scaffolds capable of providing appropriate biochemical cues to encapsulated cells. A number of previous works have created injectable materials that possess specific biomimetic properties for neural and retinal tissue engineering. In one such study, synthetic hydrogels were engineered to specifically mimic the biochemical composition and neurotrophic potential of native brain ECM. The constructed hydrogels were able to promote 3D neurite outgrowth in in vivo studies 105. In another study, researchers designed an injectable and bioresorbable hydrogel polymer composed of hyaluronan and methycellulose (HAMC). This hydrogel platform was shown to enhance cell survival and integration of transplanted retinal stem cell (RSC)-derived rods in the retina 106. However, limitations of the chemical conjugation process can constrain the amount and type of biomolecules conjugated to the polymer backbone.
Researchers have also begun altering the mechanical properties of injectable scaffolds to induce specific cellular behavior of implanted cells. Matrix stiffness varies with respect to the different organ systems in the body. For example retinal tissue shows a very low modulus of around 100 Pa83 while bone has a modulus closer to 50 kPa 107 Cells of different lineages naturally reside in tissue types with specific mechanical properties. Proving this point, researchers have shown that stem cells grown on stiff materials have a higher propensity to differentiate into bone cell
27


progenitors 107 while stem cells grown on softer material are more likely to differentiate into neural cell lineages 108. The different mechanical properties of each tissue type will alter how they interact with cells. To encourage incorporation and prevent damage, polymer materials used to replace native tissue should possess similar mechanical properties to the tissues they are aimed at replacing. Neural tissue has an extremely low modulus which makes surgical implantation of a material soft enough to replace it relatively impossible 83. For this reason, injectable scaffold are again beneficial when trying to replace damaged neural tissue specifically.
Finally, researchers have also begun to modify the properties of scaffolds to better mimic the morphological structure of native tissue. In native retinal tissue specifically, RGCs grow in a laminar formation with axons extending in a unilateral direction towards the optic nerve cup 109. However previously developed polymer scaffolds used in retinal tissue engineering simply form a porous structure forcing axons to extend in all directions. To better mimic native retinal tissue, researchers have begun to use electrospun scaffolds 110111 and retinal sheets to enhance transplantation into the retinal space 112. One research group was able to show axon growth in a pattern consistent with the structure of native retinal tissue through guidance of embedded RGCs using biochemical and morphological cues 11 \ In another study, retinal sheets taken from rats were used to construct a scaffolding system that better mimics the sheet formation of native retinal tissue 112. However, once again these materials require surgical implantation and likely do not possess a comparable modulus to soft neural tissue. The ideal material would be one that could be injected and take the shape of native retinal tissue while possessing the appropriate mechanical properties.
As you can see, much of the focus of retinal tissue engineering has been to construct a material that can mimic native tissue in the closest way. In this work, the goal was to engineer a injectable
28


material that can mimic the biochemical, mechanical, and even the morphological properties of native retinal tissue. The first component of this polymer system is a previously described highly funcationalizable backbone with a protein-like structure aimed at mimicking a cell’s native extracellular matrix 23. The second component of this polymer system is arginine-glycine-aspartic acid (RGD), an integrin/cell-binding motif found in many components of the ECM. It has been well established that cellular adhesions made through RGD-integrin binding can promote cell survival, cell spreading, proliferation, and neurite extension. RGD was chemically conjugated to the polymer backbone, in order to provide a stable linkage capable of withstanding the cellular contractile forces and promote strong cellular adhesion 113. Lastly, we modified this polymer system with poly (N-isopropylacrylamide) (PNIPAAm) to allow for an injectable cellular scaffold. PNIPAam, as discussed above is a thermosensitive water-soluble homopolymer that exhibits a sharp, reversible sol-gel transition point sufficiently above room temperature and sufficiently below body temperature (32°C)114 making it extremely useful for many biomedical applications 88. When conjugated with other polymers, PNIPAAm may impart its thermosensitive behavior on otherwise non-stimuli sensitive polymers or molecules. Following synthesis and characterization of this polymer system (Chapter 2), we investigated the 3D growth of RGCs cultured within this polymer scaffold (Chapter 3). The results of this work provide insight into the complex mechanical and chemical environment that support RGC growth in 3D and could have a direct application in cellular replacement therapies for the treatment of glaucoma and other neurodegenerative diseases.
Injectable systems for neurotrophic factor delivery
In this final section we will discuss the use of injectable materials for the delivery of protein drugs (such as NFs) in the setting of retinal tissue engineering. Importantly, the end goal of this work would be a polymer system that may be used for concurrent NF delivery and cell transplantation to maximize cell survival, integration, and function. Initial studies to administer
29


NF consisted of direct injection of the NF to the treatment site. This method presents with physiochemical instability, rapid diffusion, and a short half-life. Due to this, multiple high concentration doses would be required to realize the full effects of the NF, making the use of NFs impractical and potentially dangerous in the clinical setting 115. For this reason, an appropriate delivery system that can protect protein drugs from enzymatic release while controlling release is an encouraging approach to improve protein stability and control release within the body. In this section we will focus on synthetic, injectable materials.
A number of different polymer system have been investigated for controlled release of protein drugs. Amongst them, hydrogels have become a popular material for protein drug delivery with an injectable material. In one such study, investigators used photoinitiated polymerization to contruct hydrogels based on polyethylene glycol). They then used the degree of polymerization to control the neurotrophic diffusion from the hydrogel system. The release of NF (specifically CNTF) was assessed using a neurite outgrowth assay with cells taken from retinal explants. CTNF released from the hydrogel polymer triggered significantly more neurite outgrowth compared to controls 116. In a similar study, researchers studied PEGPLA photocrosslinkable hydrogels that release BDNF. They were able to show an increase in neurite survival and neurite extension when the NF eluting polymer was cultured with retinal explants. However, these results were absent by day 14 lending to the limitations of the hydrogel systems 117 There are many more variations of polymer based protein drug delivery systems such as microspheres, nanoparticles, etc; however, none of these can simultationously be used as cell transplantation systems. Furthermore, the hydrogel systems that may also be used as scaffolds remain limited by the length of NF delivery and the setbacks discussed above.
30


In this work, we utilized the same polymer system backbone discussed in the previous section, PSHU, capable of attaching a large quantity of functional groups. In this case, the polymer backbone will be further modified with PNIPAAm as well as negatively charged sulfonate groups similar to those seen in Heparin. Heparin, a naturally sulfated biopolymer with an intrinsic negative charge, exists in the ECM where it interacts with the overall positive charge of NFs 118. This electrostatic interaction between the polysaccharide and the NF protects the NF from proteolytic degradation, preserves its bioactivity, and increases its half-life 119. However the use of heparin itself as a NF delivery system has serious drawbacks including, batch-batch variability in structure and biocompatibility as well as unwanted biological interactions with nontarget tissue 5912°. Therefore, it is more advantageous to incorporate solely the functional elements of heparin (negative sulfonate groups) into the polymer system itself. The biomimetic polymer backbone and conjugated sulfonate groups are aimed at the improvement of NF stability and long-term stable release of the GF. Following characterization of this polymer system (Chapter 2 and Chapter 4), we investigated the release of NF from the polymer system using an optic nerve rat glaucoma model (Chapter 4).
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CHAPTER II
SYNTHESIS AND CHARACTERIZATION OF A BIOMIMETIC REVERSE THERM AT, GEL FUNCTIONALIZED WITH INTEGRIN BINDING RGD MOTIFS FOR 3-DIMENSIONAL RETINAL TISSUE ENGINEERING
Cell replacement therapies have become a promising area of research for the treatment of many diseases including neurodegenerative diseases such as glaucoma. Although this type of research has shown some promise, there are still many complications to overcome. One obstacle is the lack of an implantable cell scaffold that can provide transplanted cells with a growth-permissive environment following transplantation. We aimed to overcome this hurdle by synthesizing a bioactive injectable material that may be injected in conjunction with the replacement cells to provide mechanical and bioactive cellular support. We initially synthesized poly(serinolhexamethylene urea)-co-poly(N-isopropylacrylamide) (PSHU-PNIPAAm) and subsequently conjugated a pentapeptide, Gly-Arg-Gly-Asp- Ser (GRGDS) (PSHU-NIPAAm-GRGDS). In this work, the polymer synthesis was first optimized and each component characterized. Next we analyzed the gelling properties to show ideal temperature dependent thermal gelling properties. Finally, the physical structure of this polymer was analyzed to show the morphology once gelled. The results of this section show that this polymer system has the potential to be manipulated immensely and is highly suited for retinal tissue engineering.
Introduction
Substantial research has been conducted to investigate the signals responsible for promoting neural cell survival and axon extension in hopes of incorporating these signals to improve cellular scaffolds. Unfortunately, this task has proved to be extremely challenging due to the complex progression of signals required to induce specific cellular behavior 121. In the human
32


122 123
body, cells reside in cellular niches that control cell behavior through various cell signals These signals are complex and multifaceted including cell-cell interactions, cell-biomolecule interactions, and ECM interactions. For this reason, great interest has been given to synthetic scaffolds that mimic the ECM environment124. Synthetic polymers have easy to control properties such as stiffness or morphology as well as allowing the attachment of functional biomolecules. Thus polymer systems may be tailored to promote cell adhesion, proliferation, and survival of the encapsulated cells giving potential for use with cellular transplantation. We propose the use of a synthetic polymer that can be fine-tuned and manipulated to mimic the signals found specifically in the retinal ECM.
PSHU (Figure 2.1) was employed due to its protein-like backbone structure and its potential to attach a large quantity of biomolecules (18 potential linkages per molecule). The protein-like backbone structure of the polymer may provide a cellular environment more similar to the naturally occurring proteins within the extracellular matrix. The functionalizable aspect of this polymer is extremely beneficial for these purposes because achieving a high concentration of biomolecules for cell-biomolecule interactions plays a crucial role in cell survival and axon extension 125.
The RGD sequence, an integrin-binding motif found in fibronectin and laminin (major components of the ECM), was found to be implicated in outside-inside cell signaling that can affect cell proliferation, migration, cell survival, and axon extension in the case of neurons 126. Therefore, incorporating enough of this RGD sequence into synthetic polymers to produce a synthetic scaffold has the potential to increase cell attachment and health 127 Not only do cell-biomolecule interactions play a role in directing stem cell fate, but also cell-ECM interactions can help to modulate neural stem cell behavior and differentiation 128. To cater to these
33


interactions, we engineered this polymer to possess multiple peptide-mimicking bonds to increase biocompatibility of the polymer. Furthermore, we aimed to create a material that is mechanically and morphologically similar to native retinal tissue. The use of a retina-like scaffold could induce the same morphology and axon extension seen in the retinal ganglion layer. Finally the polymer system was incorporated with PNIPAAm to allow delivery through simple injection. Perhaps the ultimate goal of tissue engineering is to deliver scaffolding through a minimally invasive injection 129. This way, damaged and diseased niches may be replaced with cell loaded injectable synthetic niches without causing additional damage through surgical or implantation procedures. PNIPAAm is a thermo-responsive polymer with a LCST of around 32°C. After conjugation of PNIPAAm to the polyurea backbone (PSHU-PNIPAAm), the entire polymer system will possess these RTG properties.
The first goal of this work was to synthesize and characterize an appropriate injectable scaffold. To confirm and characterize the polymer backbone we employed nuclear magnetic magnetic resonance (NMR) and fourier transform infrared spectroscopy (FT-IR). In addition, LCST and gelling studies were used to investigate the temperature and concentration required to form a stable gel. Finally, scanning electron microscopy (SEM) was used to investigate the morphology of the polymer once gelled and rheological studies were done to determine the storage modulus of the polymer once gelled.
34


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0 0 0
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Figure 2.1. Schematic synthesis of PSHU-PNIPAAm-RGD. The first stage is PSHU synthesis followed by the removal of the N-BOC groups from the PSHU backbone and finally RGD conjugation.
Materials and Methods
Materials
N-BOC-Serinol, urea, hexamethylene diisocyanate (HDI), anhydrous chloroform, and anhydrous N,N-dimethylformamide (DMF) were purchased from Sigma-Aldrich (St. Louis, MO, USA). N-(3-Dimethylamino- propyl)-N'-ethylcarbodiimide hydrochloride (EDC), N-hydroxysuccinimide (NHS), 2,2,2- trifluoroethanol (TFE), and trifluoroacetic acid (TFA) were purchased from Alfa


Aesar (Ward Hill, MA, USA). Anhydrous diethyl ether was purchased from Fisher Scientific (Pittsburgh, PA, USA). Anhydrous dichloromethane (DCM) was purchased from JT Baker (Phillipsburg, NJ, USA). The pentapeptide Gly-Arg-Gly-Asp-Ser (GRGDS) was purchased from Biomatik (Wilmington, DE). Dialysis tubing (Spectra/Por) was obtained from Spectrum Labo- ratories (Houston, TX). Equipment.
Equipment
Gel permeation chromatography (GPC) was recorded on a Viscotek GPC Max with 270 Dual Detector with right angle light scattering (RALS) and VE 3580 refractive index (RI) detector from Malvern Instruments (Houston, TX USA). Each sample’s number average molecular mass (Mn) and weight average molecular mass (Mw) was calculated using OmniSEC 5.02 software. 'H NMR spectra were recorded on a Varian-500 NMR (500 MHz, Varian) with CDCI3 as the solvent at 25°C. Chemical shifts are in ppm using the solvent peak as the internal reference. Fourier transform infrared (FT-IR) spectra were recorded on a Nicolet 6700 (Thermo Fisher Scientific, Waltham, MA) using polyethylene-windowed equipment. High-performance. LCST measurements were made with a Cary UV-Vis spectrophotometer using quartz cuvettes with 1 wt% solutions of polymer. Polymer morphology was imaged using a JEOL (Peabody, MA) JSAM-60101a analytical scanning electron microscope. Rheological measurements were performed on a stress-controlled rheometer (Rheostress Haake RS 150) using a cone- and-plate geometry (angle of 1°, diameter of 60 mm) and a solvent trap to prevent evaporation of the polymer solution.
Synthesis of poly (serinol hexamethylene urea) (PSHU)
PSHU was synthesized by combining HDI (1.928ml, 12mmol), N-BOC-Serinol (1.147g, 6mmol), and urea (0.360g, 6mmol) in 6ml of anhydrous DMF. This reaction was maintained at 90°C for 7
36


days. After the 7-day reaction period, the solution was cooled to room temperature and rotary-evaporated at 70°C to remove the DMF. Next, the product mixture was re-dissolved in a small amount of DMF (2-3ml) and precipitated into cooled anhydrous diethyl ether (100ml). The purification process was repeated twice and the solution was washed overnight in excess ether (100ml) to remove any unreacted reactants and remaining solvent. Finally, the product was rotary-evaporated at 45°C until dry; the final product was stored at room temperature until conjugation with RGD.
De-protection ofPSHU
In order to expose free amine groups on the PSHU backbone and subsequently conjugate PNIPAAm and RGD, the tert-Butyloxycarbonyl (BOC) groups must first be removed. To do this PSHU (l.Og) was dissolved in equal parts of methylene chloride (15ml) and trifluoroacetic acid (TFA) (15ml) in a round bottom flask. This reaction was held at room temperature with gently stirring for 30 min. Rotary evaporation was used to remove the solvent and the product mixture was dissolved in DMF (1ml). The solution was precipitated in cold diethyl ether and rotary evaporation was used to remove the residual ether. This precipitation and evaporation method was carried out once more and the deportected PSHU (dPSHU) was dissolved in 1ml of TFE and precipitated again in cold diethyl ether. After final rotary evaporation, the purified polymer was stored at room temperature.
Synthesis of poly (N-isopropylacrylamide) (PNIPAAm)
PNIPAAm-COOH was prepared as previously described 130. In brief, NIPAAm (44.19mmol) and ACA (0.22mmol) were dissolved in 25ml anhydrous methanol. Next, this solution was bubbled with nitrogen for 30 min and mixed for 3h at 68°C. This product solution was precipitated in hot water (60°C) to purify the PNIPAAm-COOH. The product was washed twice in hot water (60°C)
37


and further purified by dialysis (MWCO: 3500Da) for 2 days. This final product was freeze-dried and stored at room temperature (yield: 92%, Mw 11,000)
Conjugation of PNIPAAm to PSHU
PNIPAAM was conjugated to 25% of the free amine groups on the PSHU backbone using EDC/NHS chemistry. To begin, PNIPAAm-COOH (0.04mmol) was dissolved in 2 ml anhydrous DMF. EDC (0.048mmol) and NHS (0.048mmol) were dissolved in 1 ml anhydrous DMF and added drop-wise to the PNIPAAM-COOH dissolved in DMF. This activation reaction was allowed to proceed for 24hrs. Next, deprotected PSHU (0.1956mmol amine groups) dissolved in 1 ml DMF was added to the activated PNIPAAM and allowed to conjugate for 24 hrs. The resulting polymer was precipitated three times in cool diethyl ether and the solvent was removed each time with rotary evaporation. The dried polymer was dissolve in 5ml miliQ water at 4°C, dialyzed (MWCO: 12,000Da) for three days, and finally lyophilized).
Conjugation of RGD to PSHU-PNIPAAm copolymer
A similar conjugation protocol was used to conjugate RGD to PSHU as well as conjugating RGD to PSHU-PNIPAAm. RGD (0.125g, 0.25mmol), EDC (0.143g, 0.75mmol), and NHS (0.86g, 0.75mmol) were dissolved in 1 ml ultra-pure water for conjugation to PSHU-PNIPAAm. For conjugation to PSHU, 1ml of anhydrous DMF was used instead of ultra-pure water. This solution was reacted at room temperature for 2 h. PSHU-PNIPAAm (0.1 Og) or PSHU (0.1 Og) was dissolved in ultra-pure water or DMF respectively and added drop wise to the appropriate activated RGD solution and reacted for 24 h at room temperature, protected from light. The PSHU-PNIPAAM-RGD solution mixture was placed in dialysis tubing (MWCO: 3500) and dialyzed against 1 liter of ultra-pure water for 24 h, with one water change. After dialysis, the solution was freeze-dried, resulting in a white, flaky precipitate, which was stored at room
38


temperature. However, the PSHU-RGD solution mixture precipitated three times in excess ether, dried using rotary evaporator, and stored at room temperature.
Gel permeation chromatography
(GPC) (Malvern Instruments, Houston, TX USA) was used to determine the molecular weight distribution of the synthesized copolymers. Analysis was performed on each sample using a 100 pi injection into a single Viscotek D6000M column and 270 Dual Detector with right angle light scattering with DMF as the system solvent. The column and detector temperatures were kept constant at 45°C. The instrument was calibrated with polystyrene standards (MW: 105,000, dn/dc: 0.185 ml/g).
Solution to gel phase transition using LCST
To examine the thermal gelling temperature, the lower critical solution temperature (LCST) of PSHU-PNIPAAm and PSHU-PNIPAAm-RGD was measured by loading a 1% (wt/v) solution in deionized water into a UV/visible spectrophotometer fitted with a temperature-controlled cell and reading percent transmittance at 480 nm at temperatures between 20°C and 45°C.
Solution to gel phase transition using gelling tests
The sol-gel phase transition of the PSHU-PNIPAAm-RGD solutions was determined by a test tube inversion method. The polymer was dissolved in PBS with various concentrations and each polymer solution (2ml) was placed in a glass vial. The initial temperature of a water bath was set to 25°C and heated up to 50°C with l°C/min. The sol-gel phase transition was determined by inverting the vial horizontally at each temperature.
Scaffold morphology using SEM
Polymer solutions of 2.5, 5, and 10% were prepared in ultra-pure water and allowed to gel at 37 °C for 15 min. The gelled samples were then frozen quickly using liquid nitrogen, cut in half to
39


expose the center structure, and rapidly transferred to a freeze-dryer for 24 h (-48 °C, 38* 10'3). Liquid nitrogen was used to freeze the samples rapidly (~3 sec) to avoid de-gelling of the polymer and preserve its 3D structure. The samples were then sputter coated with gold for 30 s and the cross-section of the gel was analyzed using SEM.
Mechanical properties using rheology
First, polymers (PSHU-PNIPAAm-RGD and PSHU-PNIPAAm) were dissolved in RGC media at concentrations of 2.5, 5 and 10 wt%. Temperature sweep tests composed of heating ramps (at 5 °C/min) were conducted at constant frequency (1 Hz) and stress (0.05 Pa) between 25 °C and 45 °C.
Statistical analysis
Statistical significance between three or more data sets will be determined by ANOVA, while the t-test will be used to compare significance between 2 groups. A p value of < 0.05 will be considered statistically significant.
Results and Discussion
Polymer Characterization
The molecular weight distribution of PSHU was determined by GPC. GPC analysis showed molecular weights for PSHU, number average molecular mass (Mn): 1,610 Da, weight average molecular mass (Mw): 3,354 Da, and PI (Mw/Mn): 2.083 (See appendix A). NMR and FT-IR were both used to confirm the overall polymer structure and to detect any remaining BOC-protecting groups. The NMR spectrum of PSHU confirmed the expected copolymer structure, with peaks at 1.3 (-CH2-), 1.5 (-NH-CH2-CH2-), and 3.2 (-NH-CH2-) associated with HDI, at 1.4 (-C-(CH3)3), and 4.1 (-CH-NH-) associated with N-BOC-serinol (Figure 2.2). NMR
40


analysis was also able to show complete deprotection of the BOC-protecting groups from the PSHU backbone as well as the expected polymer structure. This can be seen by the absence of a peak at 4.1 (corresponding to N-BOC-serinol) in the NMR spectrum for deprotected PSHU (Figure 2.3).
A
o
O NH
'» A'
AN
Y
o
T
o
o o o AAA
H H H H
K K
Poly (serinol hcxamcthylcnc urea). Protected PSHU
I >MS( )
Chemical Shill (ppm)
Figure 2.2. 'H NMR (500 MHz, CDCI3) spectrum of PSHU to confirm overall structure of polymer chain.
41


Disappearance of peak at 1.4
Removal of BOC Groups
n
0 0
h»l> <%cnnol hc\*mcth\lcnc ureal. I)rpn*vlcd PSHl
B
o o
n
A
W) <«rnanl hruHK^Inr ml. Ph 2
Figure 2.3. *H NMR (500 MHz, CDCI3) spectmm of PSHU and deprotected PSHU to confirm removal of the BOC protecting groups during deprotection process.
FT-IR spectroscopy was used to confirm the conjugation of RGD to the polymer backbone by viewing the region of 1630-1680 cm-1 (Figure 2.4). This region is associated with the carbonyl groups found within the polymer backbone as well as the carbonyl groups found within peptide bonds of RGD. The wavenumbers correlated to the carbonyl groups of RGD are slightly lower than those of the carbonyl groups in the polymer backbone. In the PSHU-PNIPAAm-RGD spectrum, we can observe an obvious shift in this carbonyl peak toward the lower end of the spectrum, indicating the presence of carbonyl groups in the RGD peptide.
42


--dPSHU
1800 1700 1600 1500 1400 1300 1200 1100 1000 900 800 700
Wavenumber (cm-1)
Figure 2.4. FT-IR of dPSHU, PSHU, and PSHU-RGD. Confirmation of free amine groups on dPSHU after deprotection (B). Shift in carbonyl absorbance to confirm attachment of RGD to the polymer backbone (A).
High- performance liquid chromatography (HPLC) was also used to quantify the amount of GRGDS-COOH that was successfully conjugated to the free amine groups on the polymer backbone. A calibration curve was first constructed using known concentrations of GRGDS-COOH (0.78-200 pg/mL) (See appendix B). Each of these concentrations produced a corresponding HPLC peak area. Using the calculated area beneath the peaks and the corresponding calibration curve, we were able to determine the amount of RGD within each sample and thus the corresponding RGD conjugation efficiency. Results showed 93% conjugation of RGD to the free amine groups on the polymer backbone.
43


Analysis of gelling properties
LCST was used to analyze the gelling properties of both PSHU-PNIPAAm and PSHU-PNIPAAM-RGD. This data gave us valuable information on not only the temperature of sol-gel transition, but also the relative rate of gelation. As shown in Figure 2.5, the transmittance of both aqueous solutions of PSHU-PNIPAAm and PSHU-PNIPAAM-RGD decreased slowly upon heating from 20°C to 31°C, reached almost zero at 32°C, and turned to an opaque solid upon further heating over 33°C, indicating that the aqueous solution turns to a physical gel as the temperature increases. PSHU-PNIPAAm-RGD exhibited a LCST and phase transition profile very similar to PSHU-PNIPAAm, remaining in solution state at temperatures below 32 °C and rapidly undergoing a phase transition to a physical gel upon reaching body temperature. These unique characteristics will allow PSHU-PNIPAAm-RDD to be administered through a minimally-invasive injection at the desired location.
Lower Critical Solution Temperature
Temperature (°C)
Figure 2.5. Temperature-dependent phase transition of PSHU-PNIPAAm (orange) and PSHU-PNIPAAm-RGD (blue). Both polymers display similar gelling temperature while the incorporation of the RGD peptide slightly increased the gelling time.
44


Since the PSHU-PNIPAAm-RGD was designed to show temperature-dependent solution-to-gel phase transition, its thermal gelling properties were examined. Although not significant, the gelling temperature of PSHU-PNIPAAm-RGD decreased from 32°C to 31°C as the concentration of aqueous solution increased from 3 to 24% (wt) (Figure 2.6). More importantly all aqueous solutions remained a gel at 37°C indicating that PSHU-PNIPAAm-RGD is a promising temperature-dependent injectable material. Moreover, the gel status was maintained up to the highest temperature (50°C) with no phase separation.
Concentration (%, wt/v)
Figure 2.6. Thermal gelling property of PSHU-NIPAAm-GRGDS. All solutions turned to physical gel upon temperature increase and maintained gel status in a broad range (blue area) of temperature.
Morphology and mechanical properties
SEM was used to investigate the morphology of the 3D polymer scaffold after gelation at body temperature. As shown in Figure 2.7, higher concentrations (5 and 10 wt/v %) of both PSHU-PNIPAAm-RGD and PSHU-PNIPAAm assembled into a laminar sheet-like
45


conformation upon gelling. However, at a low concentration (2.5 wt/v %), both polymer systems formed a heterogeneous structure, implying unstable gelling conditions. The sheet-like conformation adopted by this polymer system is very similar to that of retinal tissue. Retinal tissue has a laminar organization in which different cell types are layered within specific retinal strata. To provide a more suitable 3D culture and induce greater axon extension, we engineered a polymer system with properties similar to that of retinal tissue structure. Retinal tissue is laminar in structure with RGC axons extending horizontally toward the optic cup. Therefore, recreating this sheet structure could provide topographical cues and, along with biomolecular cues of RGD, guide horizontal extension 131 4
Figure 2.7. Scaffold morphology. Representative SEM images of PSHU-PNIPAAm-RGD (bottom row) and PSHU-PNIPAAm (top row) at various concentrations. Concentrations above 2.5 wt/v% show a laminar sheet-like formation of the gelled polymer. Scale bar is 50 pm.
46


Rheology was used to assess the mechanical properties of the polymer scaffold as compared to native retinal tissue. Figure 2.8 shows the storage modulus (G') plotted against increasing temperature. All six polymer samples showed a G' that was greater than the loss modulus (G")(See appendix A), indicating the dominant elastic behavior of the polymer scaffold 13 \ The value for PSHU-PNIPAAm-RGD and PSHU-PNIPAAm both increased with increasing polymer concentration and temperature. In addition, the incorporation of RGD showed a slight increase in G' at all three polymer concentrations. In addition, we were able to create a polymer system with a G' similar to that of retinal tissue. At a concentration of 5 wt/v %, the polymer system possessed the laminar sheet-like morphology seen in the native tissue132 and had a G' of 200 Pa, which is comparable to the value of 100 Pa of retinal tissue,83 making this system the best choice for subsequent studies.
47


1400
_ 1200 ro
— 1000 b
to 3
o 600
800
m 400
cu
k_
o
200
. • *
• •
•x x
30 32
* * *
iHf
aaaaaaaaaaaaa
* * *
34 36 38 40
Temperature (C)
♦ PSHU-PNIPAAm 2.5% ■ PSHU-PNIPAAm-RGD 2.5% A PSHU-PNIPAAm 5%
PSHU-PNIPAAm-RGD 5% X PSHU-PNIPAAm 10% • PSHU-PNIPAAm-RGD 10%
Figure 2.8. Rheological properties of polymer scaffold. The storage modulus was plotted vs. temperature for various concentrations of each polymer system (PSHU-PNIPAAm-RGD and PSHU-PNIPAAm). Increasing either polymer concentration or temperature leads to an increase in the storage modulus of the material.
Conclusion
In this work, we have demonstrated that potential PSHU-NIPAAm-RGD has for retinal tissue engineering. The polymer backbone, PSHU, has the capacity to be modified with various moieties making it easy to manipulate for various tissue engineering applications. In this chapter we used RGD as a biomolecular conjugate; however, this polymer may be conjugated with a variety of molecules and chemical groups depending on the purpose. Furthermore, once conjugated to PNIPAAm, the PSHU polymer backbone possessed appropriate reverse thermal gel properties for injection and gelation at body temperature. In addition, the reverse thermal gelling property of this polymer system will allow facile encapsulation of cells in a three dimensional (3D) scaffold. Finally, this polymer system was created to possess both topographical and mechanical cues that mimic the native retinal microenvironment. In the
48


subsequent section, we will access the potential of this scaffold for RGC survival and axon extension.
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Chapter III
A SELF ASSEMBLING INJECTABLE BIOMIMETIC MICROENVIRONMENT ENCOURAGES RETINAL GANGLION CELL AXON EXTENSION IN VITRO
Sensory-somatic nervous system neurons, such as retinal ganglion cells (RGCs), are typically thought to be unable to regenerate. However, it is now known that these cells may be stimulated to regenerate by providing them with a growth permissive environment. We have engineered an injectable microenvironment designed to provide growth-stimulating cues for RGC culture.
Upon gelation, this injectable material not only self-assembles into laminar sheets, similar to retinal organization, but it also possesses a comparable storage modulus to retinal tissue. Primary rat RGCs were grown, stained, and imaged in this 3D scaffold. We were able to show that RGCs grown in this retina-like structure exhibited characteristic long, prominent axons. In addition, RGCs showed a consistent increase in average axon length and neurite-bearing ratio over the 7-day culture period indicating this scaffold is capable of supporting substantial RGC axon extension.
Introduction
Glaucoma, amongst other optic neuropathies, leads to the neurodegeneration of retinal ganglion cells (RGCs), the projection neurons located in the retina with axons extending through the optic nerve \ These cells play a crucial role in sight by transmitting visual information from the bipolar, amacrine, and interplexiform cells of the retina to the visual cortex of the brain 133. Due to the inability of these cells to regenerate in the normal human disease condition, the loss of these cells is permanent134 Additionally, clinical therapies for glaucoma are currently limited to treatments that prevent or limit further damage to the RGCs 135,5. Future interventions that seek to regain or improve visual function must not only include mechanisms for RGC neuroprotection
50


but also methods to facilitate the survival and axon regeneration of damaged RGCs and eventually methods to replace dead RGCs.
It was previously believed that RGCs, like many central nervous system neurons, do not possess the ability to regenerate following injury or death. However, it is now known that the limited regeneration of axonal regrowth of these cells is possible but inhibited due to the injured microenvironment (myelin associate molecules) 136’137’ 138, scar formation 139’ 14°, and lack of passage across a lesion 141,142,143. Therefore, the regenerative capacity of RGCs may be stimulated by creating an alternate extracellular microenvironment that will instead activate RGC growth, maintain RGC viability, and counteract the inhibitory signals of the injured nerve 12. To alter the fate of damaged RGCs, the cells must be encapsulated in a growth permissive microenvironment, protected from the diseased environment, presented with cell binding molecules, and exposed to appropriate mechanical properties to induce and cue growth.
Here, we have developed an injectable biomimetic three-dimensional (3D) scaffold with similar mechanical and morphological properties to native retinal tissue. The first component of this polymer system is a previously described, with highly functionalizable backbone aimed at mimicking a native ECM 23. The second component of this polymer system is the small peptide RGD, an integrin/cell-binding motif found in many components of the ECM. It has been well established that cellular adhesions made through RGD-integrin binding can promote cell survival, cell spreading, proliferation, and neurite extension23. In this study the peptide sequence GRGDS was used instead of RGD in order to preserve the integrity of the entire RGD binding motif. Studies have shown that the RGD tripeptide has little effect on cell attachment; however surrounding the RGD motif with flanking amino acids according to the natural sequence (GRGDS) can preserve activity of this integrin-binding motif144 Lastly, we modified this
51


polymer system with PNIPAAm to allow for an injectable cellular scaffold. PNIPAAm, a thermosensitive water-soluble homopolymer, has garnered a lot of attention in the biomedical field 145 114 146 147 This polymer has been shown to exhibit a sharp, reversible sol-gel transition point at 32 °C, sufficiently above room temperature and sufficiently below body temperature making it extremely useful for many biomedical applications 88. Following synthesis and characterization of this polymer system (as described in the Chapter II), we investigated the 3D growth of RGCs cultured within this polymer scaffold. The results of this work provide insight into the complex mechanical and chemical environment that support RGC growth in 3D and could have a direct application in cellular replacement therapies for the treatment of RGC-associated ocular neurodegeneration and other neurodegenerative diseases.
Material and Methods Materials
Dimethyl sulfoxide (DMSO) and bovine serum albumin (BSA) were purchased from Sigma-Aldrich (St. Louis, MO, USA). Succinic acid, tetrahydrofuran (THF), anhydrous methanol, 4-Dimethylaminopyridine (DMAP), N,N’-Dicyclohexylcarbodiimide (DCC), N-(3 Dimethylamino-propyl)-N'-ethylcarbodiimide hydrochloride (EDC), chloroform-d, and N-hydroxysuccinimide (NHS) were purchased from Alfa Aesar (Ward Hill, MA, USA).
Anhydrous N, N-Dimethylformamide (DMF) was purchased from EMD Millipore (Billerica,
MA, USA). N-Isopropylacrylamide (NIPAAm) was purchased from TCI Chemicals (Portland, OR, USA). Dichloromethane (DCM), Acetone, and 30% hydrogen peroxide were purchased from BDH Chemicals through VWR (Radnor, PA, USA). Anhydrous diethyl ether was purchased from Fisher Scientific (Pittsburgh, PA, USA). Gly-Arg-Gly-Asp-Ser (RGD) was purchased from Biomatik (Wilmington, DE, USA) and Cellmano Biotech Limited (Hefei, AnHui,
52


China). Sprague Dawley rats were purchased from Charles River Laboratories (Wilmington, MA, USA). Optimal cutting temperature (OCT) compound was purchased from Sakura (Torrance, CA, USA). MTT Cell Proliferation Assay Kit was purchased from Invitrogen-Molecular Probes (Carlsbad, CA, USA). Goat serum, GFAP (mouse IgGl), GAP-43 (Rabbit IgG), Alexa Fluor 488 (goat anti-mouse IgG), Alexa Fluor 594 (goat anti-rabbit IgG), and SlowFade Diamond antifade mountant with DAPI were purchased from Life Technologies (Carlsbad, CA, USA), pill-tubulin (goat anti-rabbit, IgG) was purchased from abeam (Cambridge, MA, USA). Fluoromount G with DAPI was purchased from Electron Microscopy Sciences (Hatfield, PA, USA). Triton X-100 was purchased from MP Biomedicals. Retinal Ganglion Cell Isolation Kits (rat) were obtained from MACS Miltenyl Biotec (San Diego, CA, USA).
Equipment
Confocal images were collected using a Nikon Eclipse Ti C2 LUN-A microscope (Nikon,
Tokyo) equipped with two C2-DU3 high sensitivity PMT detectors, 4 diode lasers (405/488/561/640 nm), and a motorized microscope stage with 3-axis navigation (X, Y, and Z) controlled by the NIS-Elements software package. Laser and software setting were kept constant between specimens and to allow for comparison of different image acquisitions. Glass bottom culture dishes for RGC growth were purchased from MatTek Corporation (Ashland, MA, USA). Tissue was sectioned using a CryoStar NX70 Cryostat. Confocal images were taken 3i Marianas Spinning Disk microscope.
Synthesis of PSHU-PNIPAAm-RGD
See previous chapter for details on polymer synthesis and characterization
53


Animal Procedures
All animal experiments were performed in accordance with procedures approved by the Institutional Animal Care and Use Committee (IACUC) at the University of Colorado Denver Anschutz Medical Campus. All experiments were performed in accordance with IACUC guidelines and regulations. A total of two breeding pairs (Wistar rats male: 350-400 g, Wistar rats female: 250-300 g) were required to complete this project. These animals were used to obtain rat pups by housing one male and one female rat per cage. Once the rat pups reach postnatal day 5-7, the pups were euthanized by prolonged exposure to CO2 (~50 min) followed by a secondary form of euthanasia (decapitation).
Retinal ganglion cell isolation
RGCs were purified from rat pups (postnatal day 5-7) as these cells show higher survival rates following the separation process. To begin, eyes were carefully enucleated and transferred to a petri dish filled with D-PBS. Using a dissecting microscope, a small incision was made along the anterior part of the eye (behind the lens and cornea). Tweezers were inserted into the small incision and the eye was carefully pulled along this incision line to maintain the integrity of the retina. At this point, the retina was peeled away from the sclera and was moved with a transfer pipette to a 15 ml tube. The dissected retinas were then dissociated using the Neural Tissue Dissociation Kit for Postnatal Neurons from Miltenyi Biotec per the manufacturer’s instructions. Following this, the RGCs were purified using the Retinal Ganglion Cell Isolation Kit from Miltenyi Biotec per the manufacturer’s instructions. The isolated cells containing mostly RGCs were then resuspended in prewarmed RGC growth medium containing MACS NeuroMedium (130-093-570), NeuroBrew-21 (1:50 dilution, 130-093-566), sodium pyruvate (Sigma, 1 mm),
54


BDNF (Peprotech, 25 ng/ml), CNTF (Peprotech, 10 ng/ml), N-acetylcystein (50 pg/ml), insulin (Sigma, 5 pg/ml), Forskolin (Sigma, 10 pM), glutamine (2 mM), triiodothyronin (40 ng/ml), streptomycin sulfate (100 pg/ml), penicillin (100 U/ml). Cells were then plated on PDL-Laminin coated coverslips (Data not shown) or seeded in 3D polymer scaffolds. Three separate RGC isolations were completed on three different rat pup litters.
3D culture of Retinal ganglion cells
RGCs (8xl03) were suspended in a solution of 5 wt/v% PSHU-PNIPAAm-RGD or 5 wt/v% PSHU-PNIPAAm in complete media. 50 pi of the cell suspension in the polymer solutions were pipetted into each 35 mm glass bottom dish (MatTek, Ashland, MA, USA) (Figure 3.1) and placed in a 37 °C incubator for 10 min to allow polymer gelation and RGC encapsulation. After incubation, 1 ml of warm, RGC media was added to the culture dish, using a hotplate set at 37 °C to maintain gel stability when removed from the incubator. Cells were cultured for 3, 5, and 7 days, with media changes each day.
Confocal microscopy
Figure 3.1. Schematic showing 3D culture process. RGCs were first mixed in with the polymer solution and loaded into a glass bottom petri dish. After gelling at 37°C, RGCs became encapsulated in the polymer scaffold and culture for future analysis.
55


Immunostaining within the 3D scaffold
Cells were immunostained while remaining encapsulated in the 3D polymer scaffold. It is important to note that all steps of the staining process were conducted at 37 °C to prevent de-gelling and disruption of the polymer sample. First samples were washed twice with PBS lx and fixed with 4% PFA in PBS for 15 min at 37 °C. Next, samples were permeabilized with 1% Triton-X (in PBS) for 90 min, followed by a PBS wash overnight. Blocking buffer composed of 2% bovine serum albumin (BSA) in PBS was added to the cells for 90 min. After blocking, cells were incubated overnight with the first primary antibody Brn3a (1:200, prepared in blocking buffer). Cells were then washed with 1% Triton-X, 3x for 3 min each. The secondary antibody, anti-goat Alexa 594 (1:500) was added to each sample and incubated for 45 min. Cells were washed with PBS-Tween (0.002% in PBS) for 3 min and washed twice with PBS, 3 min each. Cells were then incubated with the second primary antibody pill-tubulin (1:100, prepared in blocking buffer) overnight at 37 °C. Cells were then washed with 1% Triton-X, 3x for 3 min each. The secondary antibody, anti-rabbit Alexa 488 (1:500) was added to each sample and incubated for 45 min. Cells were washed with PBS-Tween (0.002% in PBS) for 3 min and washed twice with PBS, 3 min each. Hoechst 33342 (1:2000, in PBS), a DAPI stain, was added to each sample and incubated for 5-10 min, followed by 3 washes in PBS, 3 min each. RGCs were imaged within the polymer scaffold using fluorescence confocal microscopy.
Analysis of average neurite length, average branchpoint, and neurite bearing cells
Using a 20x objective, z-stack projections of 4 pm thickness were sampled from three random visual fields in each sample. At least nine samples taken from three different RGC isolations were analyzed for each time point and for each polymer (PSHU-PNIPAAm-RGD and
56


PSHU-PNIPAAm). The simple neurite tracer plugin from FIJI (Longair, Baker, and Armstrong, 2011) was used to analyze the length of each process. Each neurite was traced starting from the cell body extending out into the image frame. The total length of the neuritis was divided by the cell count determined using DAPI to obtain the average neurite length.
Statistical analysis
Statistical significance between three or more data sets will be determined by ANOVA, while the t-test will be used to compare significance between 2 groups. A p value of < 0.05 will be considered statistically significant.
Results and Discussion
31) Culture ofRGCs in an Injectable Polymer Scaffold
The reverse thermal gelling property of this polymer system will allow simple mixing of the RGCs with the polymer at room temperature and encapsulation of the cells within the 3D polymer scaffold upon reaching physiological temperatures. Cells can then be stained and imaged within the scaffold to approximately the center of the gel (250 pm) (Figure 3.1). Unlike two-dimensional (2D) systems, this injectable 3D scaffold is more attune with the cell-cell interactions, cellular organization, and microenvironment seen in native tissue. In 3D, cells can be fully encapsulated in a solid microenvironment rather than solely exposed to one flat 2D surface 129.
Along with being injectable, this polymer system is functionalized with cell binding peptides to improve cellular localization and axon extension. For RGCs specifically, the LI integrin binding molecule that contains the peptide sequence RGD was shown to be present in axonal regeneration ofRGCs 148. Because integrins modulate how cells interact with their substrate, increasing the number of LI binding sites (RGD) within the polymer scaffold could increase the
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level of cell adhesion, cell growth, and extension of axons within the 3D matrix. RGD was chemically conjugated to the polymer backbone, to provide a stable linkage capable of withstanding the cellular contractile forces and promoting strong cellular adhesion 113.
The aim of this work was to determine how RGCs would behave once encapsulated in the 3D polymer scaffold. Following optimization and characterization of the injectable material, we were able to select the polymer system that was most similar to native retinal tissue and would likely support the growth and axon extension of RGCs. RGCs from rat pups postnatal day 5-7 were purified, and seeded simply through mixing with the polymer solution. Once the cell/polymer mixture was placed in the incubator, the cells became encapsulated in the 3D scaffold. Using confocal microscopy, we were able to visualize RGC axons growing in 3D throughout the scaffold. Figure 3.2 displays maximum intensity images from 200 pm thick z-stacks with 4 pm intervals. Cells grown within PSHU-PNIPAAm-RGD (5 wt/v%) showed robust axon extension whereas cells grown in PSHU-PNIPAAm (5 wt/v%) showed minimal axon extension (See Appendix D). Furthermore, RGCs grown on the PDL-Laminin coated cover slips displayed a ‘star-like’ morphology extending shorter axons in all directions (Figure 3.3). In contrast RGCs in their native environment in the retina typically extend fewer and longer axons in a unilateral direction. Interestingly, the RGCs grown within the polymer scaffold presented a morphology similar to RGCs in vivo 109 and less like RGCs grown in 2D culture dishes.
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Day 7 Day 5 Day 3
Figure 3.2. Maximum intensity projections of representative 3D fluorescent images. 3D images of RGCs (red) cultures inside of 5 wt/v% PSHU-PNIPAAm-RGD were taken using a confocal microscope after 3, 5, and 7 days in culture. RGCs (red) show long and mostly planar axon extension (green) for the 7 day culture period.
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A m . ' B • *« • * • r A • • * * • m * e % # * % 4 a *’
100 pm 200 Mm •
Live Calcein-AM/Dead Ethidium homodimer-1
Figure 3.3. Live/dead (green/red) staining of RGC 3 days of culture. RGCs were cultured on PDL-Laminin glass coverslips and stained with live/dead. RGCs grown on this coverslips display a star-shaped morphology typical of RGCs grown in vitro.
Image analysis of 3D RGC cultures
Average neurite length was determined by measuring the total length of all axons (in 3D) within a visual field and dividing this number by the total cell count within that field. The panels A-C in Figure 3.4 show the image analysis process, where Figure 3.4A is the maximum intensity image of a visual field and Figure 3.4B shows the measured axon length in the X-Y, X-Z, and Y-Z frames. Figure 3.4C graphs the average axon length vs. culture period. We can see that there is a significant difference between the mean axon length at day 3 and day 5 (p-value < 0.001) as well as a significant difference between day 5 and day 7 (p-value < 0.01). Ratio of neurite bearing cells was calculated by determining the number of neurite bearing cells per frame (only RGCs bearing neurites with a length greater than two cell bodies were counted) and dividing this number by the total cell count (determined by nuclear DAPI stain). Figure 3.4D shows the total
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cell count in a visual field and Figure 3.4E is an example of a maximum intensity image used for neurite bearing cell counting. From Figure 3.4F we can see that there is no significant different between the ratio of neurite bearing cell from day 3 to day 5 or from day 5 to day 7; however, there is a significant difference between the ratio of neurite bearing cells from day 3 to day 7 (p-value < 0.01).
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â–  PSHU-PNIPAAm-RGD
Figure 3.4. Average neurite length and ratio of neurite-bearing cells. Quantification of average axon length was performed by analyzing maximal intensity images (A) in all three planes (B) for both PSHU-PNIPAAm-RGD and PSHU-PNIPAAm cultures at 3, 5, and 7 day time points (C). The ratio of neurite-bearing cells (F) was quantified by calculating the number of cells bearing neurites in a visual frame (E) and dividing by the total number of cells within that frame (D). Numbers are plotted for both PSHU-PNIPAAm-RGD and PSHU-PNIPAAm over the 7 day culture period. Statistical significance between groups was determined by ANOVA. *p < 0.05; **p < 0.005. Three images were taken from each of at least three different samples from three separate RGC isolations. The scale bar is 200 pms. Day 5 cultures are shown to provide examples for quantification (A-E).
Conclusion
This work aimed to test the efficacy of this 3D scaffold through in vitro studies with RGCs. By studying the survival, axon extension, and morphology of the cells grown in this 3D scaffold, we can determine whether this polymer system is a good model for growing RGCs in vitro as well as a promising scaffold for use in cell replacement therapies. In their native environment, RGCs develop in a single monolayer with long axons extending horizontally toward a single point in the back of the eye, eventually joining together to form the optic nerve. However, current in vitro
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culture conditions (2D) have been able to produce only RGCs with short axons or axons extending in all directions. Although certain groups have been able to produce RGCs showing a unilateral morphology using 3D scaffolds, these studies have used an electrospun scaffold that requires implantation and were unable to seed cells directly within the scaffold uo. To provide a more suitable 3D culture and induce greater axon extension, we engineered a polymer system with properties similar to that of retinal tissue structure 131 4 Retinal tissue is laminar in structure with RGC axons extending horizontally toward the optic cup. Therefore, recreating this sheet structure could provide topographical cues and, along with biomolecular cues of RGD, guide horizontal extension of RGC axons.
Collectively, we created a material that is similar to an RGC’s native environment, which could induce the same morphology and axon extension seen in the retinal ganglion layer. Through the use of this retina-like scaffold, we were able to mimic a cell’s native environment and provide improved RGC culture conditions in vitro. Culture of RGCs in this 3D polymer system exhibited long, prominent axons, similar to what is seen in the native retina (Figure 3.2). In addition, the average axon length and ratio of neurite-bearing cells increased steadily over the 7-day culture period (Figure 3.4).
In conclusion, we developed a 3D scaffold, providing both topographical and biochemical cues of a native retinal microenvironment, with great potential for RGC survival and extensive axon extension. We believe that this 3D polymer system could provide researchers with an improved model for in vitro RGC studies. This type of culture system could be utilized to investigate stratified structures, branching, and synaptogenesis in three dimensions while still preserving the appropriate cellular morphology 149. These results also show that this injectable scaffold could be promising for use in future cellular replacement studies aimed at treating optic neurodegenerative
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diseases.
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Chapter IV
AN INJECTABLE NEUROTROPHIC FACTOR DELIVERY SYSTEM SUPPORTS THE SURVIVAL AND REGENERATION FOLLOWING OPTIC NERVE CRUSH INJURY
In general, neurons belonging to the central nervous system (CNS), such as RGC do not regenerate. Due to this, strategies have emerged aimed at protecting and regenerating these cells. NTF supplementation has been a promising approach but is limited by length of delivery and delivery vehicle. For this study, we tested a polymeric delivery system (sulfonated reverse thermal gel or SRTG) engineered to deliver CNTF, while also being injectable. A rat ONC model was used to determine the neuroprotective and regenerative capacity of our system. The results demonstrate that one single intravitreal injection of SRTG-CNTF following ONC showed significant protection of RGC survival at both 1 week and 2 week time points, when compared to the control groups. Furthermore, there was no significant difference in the RGC count between the eyes that received the SRTG-CNTF following ONC and a healthy control eye. Intravitreal injection of the polymer system also induced noticeable axon regeneration 500 pm downstream from the lesion site compared to all other control groups. There was a significant increase in Muller cell response in groups that received the SRTG-CNTF injection following optic nerve crush also indicative of a regenerative response. Finally, higher concentrations of CNTF released from SRTG-CNTF showed a protective affect on RGCs and Muller cell response at a longer time point (4 weeks). In conclusion, we were able to show a neuroprotective and regenerative effect of this polymer SRTG-CNTF delivery system and the viability for treatment of neurodegenerations.
Introduction
The first question to ask when devising a treatment for optic neuropathies is, how do RGCs die as a result of optic nerve damage? The answer to this question is unfortunately extremely
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complex, with several molecular pathways likely contributing to RGC death \ However, recent progress in this field has increased our understanding of what leads to RGC degeneration following optic nerve (ON) injury. The answer to this question is now believed to lie with neurotrophic factors (NTFs) and their role in a healthy retina 28. NTFs are a family of small diffusible molecules that are strongly implicated in the control of adult neurogenesis, axon extension, proliferation, and cellular survival36. In a healthy eye, axon terminals are responsible for the uptake and transport of these NTFs to the RGC cell body 15°. However, when RGC axons become injured, as in glaucoma, they are unable to transport NTFs, leaving the RGCs susceptible to apoptotic signals and subsequent cell death \
These recent discoveries in the mechanisms behind RGC death following optic nerve damage have led to the formation of new potential treatments. One such example is treatments aimed at neuroprotecting RGCs from cell death and loss of function 18. NTF supplementation strategies have been one area of research extensively studied to deliver the necessary NTFs directly to RGCs and provide them with a neuroprotective effect 53 3 57 However preliminary studies using NTF supplementation typically involve delivering free NTFs unaccompanied through a simple bolus injection 151. Although somewhat promising, initial studies show that free NTFs administered through bolus injection present physiochemical instability, rapid diffusion, and a short half-life 3. Due to this, multiple high concentration injections would be required to realize the full effects of the NTFs, making the use of NTFs as it stands impractical and potentially dangerous in the clinical setting 115. To achieve effective and controlled therapeutic NTF levels, a delivery system sustaining NTF expression in the specific zone of interest while maintaining NTF bioactivity must be designed. For these reasons, research surrounding polymeric delivery systems has become of particular interest for making the use of NTF therapeutics a viable
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treatment option for glaucoma 152 153. An ideal polymeric delivery system would allow 1) simple injectable delivery of the NTF 2) NTF localization 3) functional properties to improve NTF stability, and 4) long-term stable release of the NTF into the surrounding tissue. To answer these requirements, we propose the use of a multi-component injectable reverse thermal gel (RTG) polymer system, poly(serinol hexamethylene urea) backbone (PSHU) conjugated to poly(N-isopropylacrylamide) PNIPAAm (PSHU-PNIPAAm), functionalized with sulfonate groups (Sul-PSHU-PNIPAAm or SRTG) (Fig SI). The first component, PNIPAAm, will provide the RTG properties and allow injectability of this polymer as well as entrapment of the NTF within the polymer matrix. The next component of this polymer system is a functionalizable PSHU backbone capable of attaching a large quantity of functional groups (18 potential linkages per molecule). In this case, the polymer backbone will be modified with PNIPAAm as well as negatively charged sulfonate groups to make the SRTG. Modeled off of interactions between native extracellular matrix (ECM) and NTFs (e.g. heparin interacting with NTFs), this electrostatic interaction between the sulfonate groups and the positively charged NTF can protect the proteins from proteolytic degradation, preserve their bioactivity, and increase their half-life 119. This SRTG system can be administered itravitreally and provide RGCs with localized and sustained release of NTFs, promoting substantial neuroprotection of RGCs following ONC damage. To analyze this delivery system, we will utilize cilliary neurotrophic factor (CNTF) as a model NTF. Early studies first identified increased CNTF as a response to disease conditions or injury of retinal neurons, after axotomy, ischemia, and experimental glaucoma 38 39. It was further discovered that CNTF works to provide neuroprotection to the RGCs as well as induce axon regeneration following injury 54
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The first aim of this work is to design and characterize an injectable polymer system that is capable of sustaining bioactive NTF release to the retinal ganglion cell layer. Lower critical solution temperature (LCST) and energy-dispersive X-ray spectroscopy (EDS) will be used to analyze gelling properties and quantify the functionalization process respectively. Next, a release test with IR-labeled CNTF and subsequent fluorescent analysis will be performed to determine the release profile of CNTF as released from the SRTG and RTG. Furthermore, the effects of the SRTG system on RGC neuroprotection and axon regeneration following ONC will be examined (with the appropriate controls) after 7-days, 14-days, and 28-days post-ONC. Bm3a, growth associated protein 43 (GAP-43), and glial fibrillary acidic protein (GFAP) staining will be used to study the neurprotective effects, axon regeneration, and Muller cell activation respectively.
Materials and Methods
Materials
N-BOC-Serinol, urea, hexamethylene diisocyanate (HDI), anhydrous chloroform, and anhydrous N,N-dimethylformamide (DMF), N-isopropylacrylamide (NIPAAm), paraformaldehyde (PFA), sucrose, bovine serum albumin (BSA), and 4,4‘-azobis(4- cyanovaleric acid) (ACA) were purchased from Sigma-Aldrich (St. Louis, MO, USA). N-(3-Dimethylaminopropyl)-N-ethylcarbodiimide hydrochloride (EDC), potassium tert-butoxide (t-BuOK), 1,3-propane sultone (PS), N-hydroxysuccinimide (NHS), and trifluoroacetic acid (TFA) were purchased from Alfa Aesar (Ward Hill, MA, USA). Anhydrous diethyl ether was purchased from Fisher Scientific (Pittsburgh, PA, USA). Anhydrous dichloromethane (DCM) was purchased from JT Baker (Phillipsburg, NJ, USA). The pentapeptide Gly-Arg-Gly-Asp-Ser (GRGDS) was purchased from Biomatik (Wilmington, DE). Dialysis tube (Spectra/ Por) was obtained from Spectrum Labs
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(Houston, TX). Phosphate Buffered Saline (PBS) was purchased from Thermo Scientific. Wistar rats were purchased from Charles River Laboratories (Wilmington, MA, USA). Sterile saline, isoflurane, ketoprofen, and xylazine were purchased from MWI Veterinary Supply (Boise, ID, USA). Optimal cutting temperature (OCT) compound was purchased from Sakura (Torrance,
CA, USA). Superfrost glass slides were obtained from Fisher Scientific (Chicago, IL). GFAP (mouse IgGl), GAP-43 (Rabbit IgG), Alexa Fluor 488 (goat anti-mouse IgG), Alexa Fluor 594 (goat anti-rabbit IgG), and SlowFade Diamond antifade mountant with DAPI were purchased from Life Technologies (Carlsbad, CA, USA). Triton X-100 was purchased from MP Biomedicals.
Equipment
Lower critical solution temperature (LCST) measurements were made on a Cary 100 UV-visible spectrophotomer (Agilent Technologies, Inc., Santa Clara, CA). Elemental analysis of non-coated samples was performed by scanning electron microscope (SEM) (JEOL JSAM-60101a analytical scanning electron microscope, Peabody, MA, USA) in low-vacuum (20 kV) mode, and the analysis was carried out by energy dispersive X-ray spectrometry (EDX system by ED AX). 32 gauge needles used for the injection (TSK Laboratory (Tochigi-shi, Tochigi-ken, Japan). Tissue was sectioned using a CryoStar NX70 Cryostat. Confocal images were collected using a Nikon Eclipse Ti C2 LUN-A microscope (Nikon, Tokyo) equipped with two C2-DU3 high sensitivity PMT detectors, 4 diode lasers (405/488/561/640 nm), and a motorized microscope stage with 3-axis navigation (X, Y, and Z) controlled by the NIS-Elements software package. Laser and software setting were kept constant between specimens and to allow for comparison of different image acquisitions.
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PSHU-PNIPAAm fabrication
To synthesize the RTG, We began by synthesizing a functionalizable biomimetic polymer backbone (PSHU, Mw: 10,500) using urea, N-BOC-serinol, and hexamethylene diisocyanate as described previously . PNIPAAm (Mw: 11,000) was synthesized as previously described and conjugated to PSHU to form the PSHU-PNIPAAm copolymer. To do this, PNIPAAm-COOH (0.75 g, 1.21 mmol), EDC (5x molar excess) and NHS (5x molar excess) were dissolved in 5 ml of anhydrous DMF and reacted for 24 h under a nitrogen atmosphere. Purified PSHU (0.125 g/ml) was then added to the reactant mixture and allowed to react for 48 h at room temperature. The product mixture was purified by precipitation in anhydrous diethyl ether twice followed by dialysis (MWCO: 12,000-14,000 Da) against ultra-pure water for 5 days. The final product was lyophilized at -45°C for 24 h and stored at room temperature.
Sul-PSHU-PNIPAAMfabrication
To synthesize the SRTG, sulfonation of the RTG was performed as previously described 154 In short, PS (0.034 g, 5 mmol) and t-BuOK (0.032 g, 5 mmol) were dissolved in 3 ml of anhydrous DMF. RTG (0.1 g/ml) dissolved in anhydrous DMF was slowly added to the flask and allowed to react for 3 days at 60°C under a nitrogen atmosphere. The product mixture was then precipitated in anhydrous diethyl ether twice and dialyzed (MWCO: 12,000-14,000 Da) against ultra-pure water for 48 h at room temperature. The final product was lyophilized at -45°C for 24 h and stored at room temperature.
Animals
All animal experiments were performed under a protocol approved by the Institutional Animal Care and Use Committee (IACUC) at the University of Colorado Anschutz Medical Campus. Male adult wistar rats (250-300 g) were allowed to acclimate for 1 week prior to surgical
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procedures. Rats were maintained on a 14/10-hour light/dark cycle with a continuous supply of fresh air and access to food and water ad libitum. Rats were anesthetized with 5% isoflurane in oxygen and maintained on 0.5-1% isoflurane in oxygen for the remainder of the surgery. To maintain body temperature, rats were placed on a warm recirculating water blanket. Rats were divided into four separate groups according to the experiment. The time course of the ONC and subsequent treatment included three different time points (1 week, 2 weeks, and 4 weeks). To determine the percentage of RGCs surviving post-ONC and axon regrowth, 3-5 rats were used per time-point.
Elemental analysis
Polymer solutions were prepared in ultra pure water and allowed to gel at 37 °C for 15 min. The gelled samples were then frozen quickly using liquid nitrogen, cut in half to expose the center structure, and rapidly transferred to a freeze-dryer for 24 h (-48 °C). An electron beam was scanned across a dried sample of both Sul-PSHU-PNIPAAm and PSHU-PNIPAAm. When the beam strikes the sample, signals are produced that are representative of a sample’s elemental composition.
Solution to gel phase transition.
Lower critical solution temperature (LCST) was used to analyze the gelling properties of the thermally reversible injectable scaffold. 1% (wt/v) PSHU-PNIPAAm and PSHU-PNIPAAm-RGD were be loaded in a temperature-controlled UV/visible spectrophotometer. The transmittance through the polymer solution at 480 nm was monitored as the temperature increases from 15 °C to 45 °C. The rate of temperature increase was held at 2°C/min. This constant rate allowed us to observe the gelation time of both SRTG and RTG for a better
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comparison. Gelation activity was indicated by an increase in opaqueness and therefore a decrease in transmittance.
IR-Dye conjugation to CNTF
CNTF was conjugated to IR Dye 800 CW NHS ester using a modified version of the IRDye 800 CW Microscale protocol provided by LI-COR. Briefly, lyophilized IR Dye was reconstituted in DMSO at a concentration of 15 mg/mL and then diluted to a working concentration of 0.15 mg/mL in PBS. Lyophilized CNTF was reconstituted in IX PBS at a concentration of 0.1 mg/mL. The IR dye solution and CNTF solution were added to a microcentrifuge tube in a 1:50 volume ratio, respectively, at pH 8.5 and reacted for 2 hours with shaking, protected from light. After the reaction, excess IR dye was removed using Zeba Spin Desalting Columns to purify the IR-labeled CNTF.
In-vitro CNTF Release Test
To a 1.5 mL centrifuge tube, 100 ng of IR-labeled CNTF was added to a solution of Sulfonated PSHU-PNIPAAm or PSHU-PNIPAAm for a final RTG concentration of 1% (mg/mL) with samples run in triplicate. This solution was allowed to dissolve overnight at 4°C protected from light. The following day, samples were gelled in an incubator (37°C, 5% CO2) for 5 minutes to allow stable gel formation, followed by the addition of 600 pL of IX PBS (37°C). At each time point, 300 pL of the release solution was gently mixed and removed from each sample and replenished with 300 pL of fresh PBS at 37°C. To a black, 96 well optical-bottom plate, 100 pi of each release sample was added. The fluorescence of each sample was measured using the LI-COR Odyssey Classic. Known concentrations of IR-labeled CNTF were added to the plate to generate a standard curve to calculate release sample concentrations.
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Optic nerve crush model
To begin, rats were given intraperitoneal injections of a solution (lml/kg) containing 1-2 mg/kg xylazine in sterile water. A lateral canthotomy was performed after anesthetizing the eye with proparicaine and then sterilizing the area surrounding the lateral comer of the eye with betadine. The eye was then irrigated using normal saline. Using a straight hemostat, the skin at the lateral corner of the eye was crimped all the way down to the orbit for 1 minute (to achieve hemostasis). The skin around the lateral orbit was lifted with forceps and a 0.5 cm incision was made with scissors. Blunt dissection was utilized to expose the optic nerve. Using a number 5 jeweler’s forceps, the optic nerve was crimped three times, 10 seconds each, approximately 2 mm behind the globe. Since a minimal incision was made, the procedure did not require sutures and healed properly on its own.
Intravitreal injection
Following the nerve crush procedure, animals received intravitreal injections depending on the group (saline, free CNTF (0.5 pg), RTG-CNTF (lwt/v% polymer) (0.5 pg CNTF), SRTG-CNTF (lwt/v% polymer) (0.5 pg CNTF), and SRTG-CNTF (lwt/v% polymer) (2.5 pg CNTF)). For each intravitreal injection, 5 pi of saline, 5 pi of the formulations was injected into the vitreous chamber using a 32-gauge, 4 mm long needle and a Hamilton glass syringe. When injected, the needle was inserted into the superior hemisphere of the eye at a 45° taking care to avoid any injury to the lens or retina. Following injection, the needle was held in the injection site for 30 seconds to prevent leakage of the treatment and then removed slowly.
Retina and optic nerve immunohistochemistry (IHC).
At the appropriate time points (1 week, 2 weeks, or 4 weeks), rats were euthanized using an unchanged cage and a flow rate of CO2 introduced to displace 20% of the cage volume per
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minute. A bilateral thoracotomy was performed as a secondary method of euthanasia. They eyes were then enucleated and fixed using 4 % PFA in PBS for 30 min, followed removal of the cornea and lens and an additional fixation period in 4% PFA for 30 min. The tissue was then cryoprotected with 10%, 20%, and 30 % sucrose in PBS for a total of 2 days, embedded in optimal cutting temperature (OCT) compound, and frozen at -80 °C. Longitudinal tissue sections for both analysis of the optic nerve (10 pm) as well as the retina (10 pm) were collected along the nasal-temporal plane using the optic cup as a reference point.
OCT was removed from the glass slides with a 5 min PBS wash. The slides were fixed in NBF for 10 min and washed 3 times in 0.01% tween in PBS for 3 minutes each. The sections were then permeabilized in 0.01% triton X-100 in PBS for ten minutes and washed 3 times in 0.01% tween in PBS 3 times for 3 minutes each. Non-specific binding was blocked by incubating sections in 2% BSA, 0.5% triton X-100, in PBS for 1 h at room temperature. Sections were then stained with the appropriate primary antibody GFAP (1:250), GAP-43 (1:250), or Brn3a (1:100) overnight at 4°C. All antibody dilutions were prepared in the blocking buffer. Slides were then washed with 0.01% tween in PBS 3 times for 3 minutes each and then incubated with the appropriate secondary (Alexa Fluor 488 (1:500) for GFAP, Alexa Fluror 546 (1:500) for GAP-43, and Alexa Fluror 594 (1:500) for Brn3a) prepared in the blocking buffer for 1 h at room temperature. Finally slides were and washed 3 times in 0.01% tween in PBS for 3 min each and then 3 times in dH20 for 3 min each. Stained slides were mounted with Fluoromount G with DAPI mounting medium. Slides were stored at -4°C.
Analyzing of neuronal survival
To determine the number of Bm3a positive cells following nerve crush and subsequent treatment, five images were taken from each area of each retina (two peripheral, two mid-peripheral, and
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optic nerve cup). As the number of RGCs may vary depending on the different areas of the retina, the region of the retina that was analyzed for Brn3a positive cells was held consistent by using the optic nerve cup as an anatomical marker. Only tissue sections taken from around the optic head were analyzed and used for Bm3a counting. Brn3a positive cell counts from all five images were performed using Fiji. The five cell counts from each retina were analyzed by averaging these values and comparing the cell count between each group (saline, free CNTF, PSHU-PNIPAAm-CNTF, Sul-PSHU-PNIPAAm-CNTF, and healthy eyes). Each data point represents the mean +/- SD of the surviving neurons post crush. Statistics were performed using a one-way ANOVA-Tukey post hoc test, * p<0.01, ** p<0.001, ***p<0.0001 n=3-5 eyes/group/time-point. Quantification of axonal growth
The extent of RGC axon growth was evaluated using GAP-43 immunostaining of longitudinal sections of the optic nerve (10 pm). Composite images of whole cross-sections of the retina and the optic nerve were imaged using lOx magnification (Nikon). Z-stack maximum intensity images of 5 total stacks were taken to encompass the full section height (10 pm). Axonal growth was quantified by determining the average level of fluorescence in regions of nerve 100 pm, 250 pm, 500 pm and 1000 pm downstream from the lesion site. During analysis, the injury site was determined in the same tissue section that had been stained with GAP-43 through dark field microscopy. Average pixel intensity for each retinal section was analyzed for n= 3-5 per group per time-point and expressed as mean ± S.E.M.
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Statistical analysis
All results are expressed as means ± standard error of the mean. Analysis of variance (ANOVA) was used to determine significant differences between groups. Statistical significance was considered when p < 0.05.
Results
Characterization of negatively charged SRTG
We began by characterizing PSHU. Both 1H nuclear magnetic resonance spectroscopy (NMR) and fourier transform infrared spectroscopy FT-IR (data previously reported) 23 154 were used to confirm the overall polymer structure and ensure the presence of free functionalizable amines on the backbone. Elemental analysis, through EDS, was used to confirm the sulfonation process and synthesis of SRTG 155. The approximate mass percent of sulfonate groups chemically conjugated to the SRTG was 0.09% as compared to 0.00% seen in RTG (Fig. S2). Next, we examined the gelling properties, through LCST, of the polymer system before and after sulfonation. The LCST describes the temperature that the polymer system transitions from solution to solid form and is therefore an extremely important characteristic for this application, allowing for facile injection into the intravitreal space. It is possible that the introduction of sulfonate groups on the polymer backbone could alter the gelling properties of the polymer system. Fig. S3 displays the LCST of both SRTG and RTG. We can see that the sulfonated polymer system still displayed very similar gelling properties to the RTG precursor, including a similar temperature of gelation and relative rate of gelation. This is likely because the addition of the small sulfonate groups to the polymer backbone is not sufficient to alter the large hydrophobic/hydrophilic properties of the PNIPAAm polymer that dictates the liquid-to-solid phase transition. Finally we analyzed the release profile of CNTF from each of the polymer types (SRTG and RTG). As discussed earlier, negatively
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charged sulfonate groups were conjugated to the polymer backbone to form the SRTG. The motivation behind this step was to encourage polymer-protein interaction between the negatively charged sulfonate groups and the large positively charged receptor binding site on CNTF 156. These interactions would not only preserve the integrity of the protein but also prolong the duration of release by dampening the burst release of CNTF from the polymer system. To compare the release profiles of CNTF released from the RTG to the release profile of CNTF released from the SRTG, we took samples of release buffer and analyzed these samples for fluorescently labeled CNTF using a fluorescent imaging system (Fig. S4). We can see from Fig S4 that the 1% SRTG polymer showed a decreased burst release of CNTF compared to the initial burst release from the 1% RTG polymer samples. This indicates that the increased charge diminished the initial expulsion of the CNTF from the polymer potentially leaving more CNTF entrapped within the polymer system. The decreased burst release could prolong the duration of CNTF release from the SRTG lending to the superior results during in vivo studies.
Delivery of CNTF from the SRTG promotes survival ofRGCs post ONC We first investigated the neuroprotective effects of the NTF delivery system. To do this, the polymer system loaded with CNTF or a control treatment was applied as an intravitreal injection following ONC and remaining RGCs were analyzed using immunostaining procedures. Experimental groups included animals with ONC that received: (i) a single intravitreal injection of saline (5 pi) (n=5, 1 week; n=5, 2 weeks); (ii) a single intravitreal injection of free CNTF (0.5 pg in 5 pi saline) (n=3, 1 week; n=5, 2 weeks); a single intravitreal injection of SRTG (5 pi) (n=5, 2 weeks), (iii) a single intravitreal injection of RTG-CNTF (5pl 1 wt/v% polymer solution loaded with 0.5 pg CNTF) (n=5, 1 week; n=5, 2 weeks); (iv) a single intravitreal injection of SRTG-CNTF (5 pi 1 wt/v% polymer solution loaded with 0.5 pg CNTF) (n=5, 1 week; n=5, 2
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weeks; n=5, 4 weeks), and (v) a single intravitreal injection of SRTG-CNTF (5 pi 1 wt/v%
polymer solution loaded with 2.5 pg CNTF) (n=3, 2 weeks; n=3, 4 weeks) (Fig. 1).
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Figure 4.1. Experimental protocol and animal groups. ONC was performed followed by the appropriate intravitreal injection. At one-week, two-weeks, or four-weeks post ONC, rats from each group were euthanized and the tissue was analyzed (A). Both 1 week and 2 weeks time points included the groups: saline, CNTF (0.5 pg), RTG-CNTF (0.5 pg), and SRTG-CNTF (0.5 pg) (n=3-5). The 2 week and 4 week time points also included groups: SRTG-CNTF (2.5 pg) (B).
At the correct time points, animals from each group were analyzed for RGC survival using Bm3a immunostaining followed by confocal imaging of the ganglion cell layer. At one week post-ONC, the group that received a single intravitreal injection of saline (Fig. 2A) and the group that received a single intravitreal injection of free CNTF (0.5 pg) (Fig. 2B) showed a decrease in the number of RGCs. In addition, quantitative analysis showed that a comparable number of RGCs survived post-ONC in the saline group (3.52 ± 1.945 RGCs/400 pm retina, mean ± SEM)
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compared to the free CNTF group (5 ± 2.511 RGCs/400 pm retina) (Fig. 21). These numbers are in fact consistent with previous studies showing survival rates of RGCs 1 week post-ONC with no treatment applied being around 37%-40% 157 158, indicating that both the saline and free CNTF injections provided no significant neuroprotective effect for RGCs 1 week post-ONC. We then examined if a single intravitreal injection of polymer systems (RTG-CNTF and SRTG-CNTF) loaded with 0.5 pg of CNTF, injected at the time of nerve injury could facilitate RGC survival 1 week post-ONC. Despite a slight increase, there was no statistically significant difference between the RTG-CNTF group (5.68 ± 1.945 RGCs/400 pm retina) when compared to both the saline treatment group as well as free CNTF (Fig. 21). However, there was a dramatic increase in the survival of RGCs in the SRTG-CNTF treatment group (9.56 ± 1.945 RGCs/400 pm retina) (Fig. 21). In addition, there was a statistically significant difference between the saline treatment group and the SRTG-CNTF (p-value: <0.0001) as well as between the free CNTF group and the SRTG-CNTF groups (p-value: 0.0082) (Fig. 21), indicating that the prolonged release of CNTF from the SRTG polymer can enhance the survival of RGCs one-week post-ONC.
Two weeks post-ONC, both groups that received a single intravitreal injection of saline (0.8 ±
1.526 RGCs/400 pm retina) or a single intravitreal injection of free CNTF (0.5 pg) (1.2 ± 1.526 RGCs/400 pm retina), showed similar numbers of surviving RGCs (Fig. 2E and 2F). In addition, both of these groups showed a number of surviving RGCs consistent with previous reports on the number of surviving RGCs two-weeks post-ONC with no applied treatment. These expected results indicate that both the saline and free CNTF injections had no marketable neuroprotective affect two-weeks post-ONC but provided good controls for comparison to the polymer treatment groups. In contrast, both polymer groups showed a dramatic increase in the number of surviving
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RGCs post-ONC (Fig. 2G and 2H). One single intravitreal injection of RTG-CNTF (0.5 pg) (6 ±
1.526 RGCs/400 pm retina) or one single intravitreal injection of SRTG-CNTF (0.5 pg) (7.8 ±
1.526 RGCs/400 pm retina) both displayed a significant difference in surviving RGCs when compared to both the saline and free CNTF groups; however the SRTG-CNTF group displayed the highest mean value of surviving RGCs (Fig. 2J). To determine if the SRTG itself was having any effect on RGC survival, we included an additional group that received a single injection of SRTG alone following the ONC. Results from this study show no significant difference between an injection of SRTG and saline at the two-week time point (Fig. S5), indicating that the previously described neuroprotective effects of the SRTG-CNTF system are likely not due to the polymer itself.
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CNTF (0.5 fig)
B
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Saline CNTF (0-5 MS) CNTF (0.5 pg) Saline CNTF (0.5 MS) CNTF (0.5 MS)
♦ Mean 3.52 5 5.68 9.56 ♦ Mean 0.8 1.2 6 7.8
Figure 4.2. Expression of Bm3a in the retina following ONC and the appropriate intravitreal injection. One week (A-D) and two weeks (E-H) retina sections from the four treatment groups (SRTG-CNTF (0.5 pg), RTG-CNTF (0.5 pg), CNTF (0.5 pg), and saline) are shown in the top two rows. Blue staining indicates DAPT Red staining indicates Bm3a expression. Scale bar = 50
pm. * p < 0.05, ** p < 0.01, *** p < 0.001.
Treatment with SRTG-CNTF following ONC shows preservation ofRGCs comparable to that of a healthy eye
Following the evaluation of each treatment group as shown above, we took the analysis one step further and compared eyes treated with SRTG-CNTF (0.5 pg) following ONC with eyes that have had no ONC and no intravitreal injection (i.e. healthy eyes). This analysis was done using a Brn3a immunostain and the same quantification process as described above. Fig. 3 shows immunostained representative zoomed images of retinal cross-sections from each group. Rats
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that had received the ONC followed by an intravitreal injection of SRTG-CNTF (0.5 jag) showed significant Brn3a expressing RGCs within the ganglion cell layer of the retina (Fig. 3A). Interestingly, healthy eyes (Fig. 3B) showed a comparable number of Brn3a expressing cells in the ganglion cell layer. Further quantification comparing the Bm3a cell count between these two groups shows no significant difference at both the 1 week and 2 week timepoints (Fig. 3C).
in
Healthy Eyes
♦ Mean 7.52
c
SRTG- SRTG-
CNTF CNTF
(0.5 fiU. 1 (0.5 HU. 2
weeks) weeks)
9.56 7.8
Figure 4.3. Expression of Bm3a in retinal cross sections following ONC and Sul-PSHU-PNIPAAm intravitreal treatment after two weeks (A) when compared to a healthy eye receiving no ONC and no treatment (B). There was no significant difference observed between these two groups (C). Blue staining indicates DAPI. Red staining indicates Brn3a expression. Scale bar = 50pm.
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Delivery of CNTF from SRTG promotes growth of injured RGC axons To investigate whether the increased RGC survival seen in the polymer treatment groups correlated to greater axon regeneration, we immunostained for damaged and regenerating RGC axons within ON cross sections using a GAP-43 immunostain 159. This stain was not only used to help deduce the crush site but also to elucidate any regenerating axons downstream from the injury. After two weeks, controls injected with saline or free CNTF (0.5 pg) following ONC showed minimal amounts of axons crossing over the glial scar (Fig. 4A and 4B). Additionally, animals that received an intravitreal injection of RTG-CNTF (0.5 pg) showed negligible RGC growth downstream from the lesion two-weeks post-ONC (Fig. 4C). In contrast, the group that received an intravitreal injection of SRTG-CNTF (0.5 pg) post-ONC, showed extensive upregulation of GAP-43 well beyond the injury site (Fig. 4D). Dark field imaging was used to elucidate the lesion site in the same slide that received the GAP-43 stain (Fig. 4E-H). Quantification analysis of axon growth showed that the expression of GAP-43 was upregulated for the SRTG-CNTF group at 100 pm, 250 pm, and 500 pm past the lesion site, when compared to control groups one week post ONC (Fig. 41). For example the average pixel intensity of the GAP-43 stain seen within the SRTG-CNTF group at 100 pm, 250 pm, and 500 pm was 2709.06 ± 409.2 (mean ± S.E.M), 2200.0 ± 125.3, and 1962.1 ± 108.1, respectively compared to the saline group values of 1251.7 ± 184.1, 1383.9 ± 124.1, and 1418.5 ± 192.9 respectively. These groups showed a statistically significant difference for distances 100 pm, 250 pm, and 500 pm with p-values: 0.0182, 0.0445, and 0.0371 respectively (Fig. 41). Quantification of two-week data also showed an upregulation of GAP-43 expression for the SRTG-CNTF group at 100 pm, 250 pm , and 500 pm past the lesion site, when compared to control groups. More specifically, the average pixel intensity of the GAP-43 stain seen within the SRTG-CNTF group at 100 pm,
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250 pm, and 500 (am was 2741.3 ± 523.1 (mean ± S.E.M), 2333.2 ± 384.2, and 1903.6 ±314.7, respectively compared to the saline group value of 1510.1 ± 467.9, 1409.2 ± 350.7, and 1183.4 ± 281.5 respectively. These groups showed a statistically significant difference for distances 100 pm, 250 pm, and 500 pm with p-values: 0.0135, 0.0135, and 0.0204 respectively (Fig. 4J). Furthermore, within the SRTG-CNTF group, there was no significant difference between the 1 and 2 week time points at all distances from the lesion (100 pm, 250 pm, 500 pm, and 1000 pm).
RTG-CNTF SRTG-CNTF
Saline CNTF (0.5 pg) (0.5 pg) (0.5 pg)
Distance from the lesion (pm) Distance from the lesion (pm)
Figure 4.4. Intravitreal treatment of SRTG-CNTF (0.5 pg) induced RGC axon growth following ONC. Saline, CNTF (0.5 pg), and RTG-CNTF (0.5 pg) groups showed minimal axon regrowth two weeks post-ONC (A-C). SRTG-CNTF (0.5 pg) showed robust RGC axon growth two-weeks post lesion (D). The injury site was distinguished using dark field microscopy (E-H). Triple asterisks represent the ONC site. Average pixel intensity (mean ± S.E.M.) of GAP-43 stain present past the lesion site following intravitreal injection of saline, CNTF (0.5 pg), RTG-CNTF (0.5 pg), and SRTG-CNTF (0.5 pg) at one-week post ONC (I) and two-weeks post ONC (J). Scale bar = 200 pm. * p < 0.05, ** p < 0.01, *** p < 0.001.
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Although there was no statistically significant difference between the groups at 1000 pm downstream from the lesion site for both 1-week and 2 weeks post-ONC, we decided to image cross-sections of the nerve at 40x to detect any regeneration (Fig. 51). After staining with GAP-43, nerve sections from each group were imaged at approximately 1000 pm downstream from the lesion site (as identified through dark field microscopy). We can see that there is minimal GAP-43 staining present in groups, saline, CNTF, and RTG-CNTF at both 1 and 2 week time points (Fig. 5A-C and 5E-G). However, the rats that received an intravitreal treatment of SRTG-CNTF showed a visible increase in GAP-43 expression 1000 pm down stream from the crush site at both time points (Fig. 5D and 5H). Quantification was done through measuring the average pixel intensity of each section. At the one week and two week time points there was a statistical difference between the saline and SRTG-CNTF group (p-values: 0.0181 and 0.0004 respectively) (Fig. 5J).
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Saline
CNTF (0.5 jig)
RTG-CNTF (0.5 fig)
SRTG-CNTF (0.5 Mg)
160 £» 140
1 i>2a
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i
* Saline
CNTF (0.5 Mg) RTG-CNTF (0.5 Mg)
3SRTG-CNTF (0.5 Mg)
I
1 week 2 weeks
Time Post Lesion
Figure 4.5. Longitudinal images of GAP-43 immunostained ON approximately 1000 pm downstream from the crush site. At both the 1 week and 2 week time points, groups saline, CNTF, and RTG-CNTF showed minimal GAP-43 expression (A-C, E-G). At both 1 week and 2 week time points, SRTG-CNTF showed increased GAP-43 expression at these downstream points (D, H). Quantification of the average pixel intensity showed a significant difference between SRTG-CNTF and other groups at both 1 and 2 week time points (J). Red staining indicates GAP-43 expression. Photomicrographs were captured at 40x magnification. Scale bar = 100 pm. * p < 0.05, ** p < 0.01, *** p < 0.001.
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Intravitreal injection of SRTG-CNTF alters the Muller cell response one week and two weeks post ONC
Next, we wanted to investigate the effects of each intravitreal treatment on Muller cell activation following ONC. Muller cells are the principal glial cells residing in the retina. It is thought that Muller cells, like other types of glial cells play important roles in the support and protection of retinal neurons following injury 160 161. For example, it has been previously reported that injury to the ON caused by ONC can trigger morphological and cellular changes in Muller cells, suggesting that Muller cells play a key role in the aiding the regeneration of RGC axons 162. Some of these changes seen from activated Muller cells can aid in the neuroprotection and regeneration of a damaged retina 163 164. However, it is important to note that the complete function of Muller cells may not be limited to a regenerative effect. Muller cell activation is also linked to the inflammatory response and while inflammation is part of the regenerative process, prolonged inflammation can lead to additional tissue damage 165 166. To see resulting Muller cell activation caused by ONC and the subsequent intravitreal treatment, we performed retinal immunostaining using an antibody against GFAP (Fig. 6A-H). At one and two weeks post-ONC, the treatment groups that received saline and free CNTF (0.5 pg) injections showed GFAP staining that was limited to the astrocytes and Muller cell end-feet present within the nerve fibre layer (Fig. 6A-B and Fig. 6E-F). The treatment group that received and intravitreal injection of RTG-CTNF (0.5 pg) showed an increase in Muller cell staining with expression moving down through the retina at both the one week and two week time points (Fig. 6C and 6G). However, the treatment group that received an intravitreal injection of SRTG-CNTF (0.5 pg) following ONC displayed robust GFAP labeling with a large number of Muller cell processes observed spanning the entire retina (Fig. 6D and 6H). We further quantified these results by analyzing the
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fluorescence of the stained and imaged section. Figure 61 shows a significant difference between the SRTG-CNTF group and the other three groups at the one week time point (p-value 0.025). Similarly, figure 6J shows a significant different at the two week time point between the SRTG-CNTF and the remaining three groups (p-value 0.0037). These results indicate that the release of CNTF from the SRTG system was able to induce a significant Muller cell response that was likely contributing to the neuroprotective and neuroregenerative effect we were seeing in this group.
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♦ Mean 15.VW504 127.4424* I26.2422K 26K.7V36
Figure 4.6. The response ot Muller glial cells to ONC and each subsequent intravitreal injection examined using a GFAP antibody. After one week post-ONC and appropriate intravitreal injection, GFAP staining in Muller cells is shown in saline, CNTF, RTG-CNTF, and SRTG-CNTF groups (A-D). Quantification shows an increase in Muller cell expression seen in the SRTG-CNTF group compared to the saline control at the one week time-point (I). At two-weeks post-ONC and appropriate intravitreal injection, there was an increased amount of Muller cell expression in all groups (E-H). There was a significant increase in the SRTG-CNTF group compared to all other groups (J). ONL: outer nuclear layer; INL: inner nuclear layer; IPL: inner plexiform layer; GCL: ganglion cell layer; FL: fibre layer. Staining indicates GFAP expression. Photomicrographs were captured at 20x magnification. Scale bar = 100 pm. * p < 0.05, ** p < 0.01, *** p < 0.001.
Increasing the amount of CNTF can prolong the neuroprotective effect of the polymer system
As mentioned earlier, we were not able to see a significant difference between the SRTG group and the remaining three groups 4-weeks post ONC when using a low CNTF amount
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(0.5 pg). To determine if the low amount of CNTF was the reason for the limited effect length, we increased the CNTF amount to a higher amount (2.5 jag) and analyzed the effects of SRTG-CNTF (2.5 jag) at two weeks and four weeks, and compared these results to corresponding SRTG-CNTF (0.5 pg) and saline control groups. Figure 7A shows the quantification following Brn3a staining to elucidate RGC survival post ONC. We can see from this figure that the groups SRTG-CNTF (0.5 pg, 2 weeks), SRTG-CNTF (2.5 pg, 2 weeks) and SRTG-CNTF (2.5 pg, 4 weeks) all showed a significant difference in the number of RGCs that survived post ONC (all p-values: < 0.0001). In addition, we analyzed the Muller cell response with this higher dose at extended time points. Figure 7B shows the quantification of fluorescence from the GFAP stained slides. We can see from these results that there was a significant difference between the SRTG-CNTF (2.5 pg, 2 weeks) and SRTG-CNTF (2.5 pg, 4 weeks) compared to the controls (p-values: < 0.0001), demonstrating that this increase in CNTF concentration caused a Muller cell response well into the 4-week time point indicating longer term neuroprotection and nueroregeneation.
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SRTG-CNTF (0.5 jig)
SRTG-CNTF
SRTG-CNTF (0.5 fig)
(2.5 jig) 10 a u E ♦ *** ***
B _c _ N S SE u U 6 E 3 c + 4 i > w w ♦ \ A **• ♦
D - 0
-2 Saline SRTG-CNTF (0.5 fig. 2 weeks) SRTG-CNTF (2.5 fig. 2 weeks) SRTG-CNTF (0.5 jig. 4 weeks) SRTG-CNTF (2.5 fig. 4 weeks)
♦ Mean 0 7.45 6.3 0.15 2.8
SRTG-CNTF (2.5 Jig) 800 41 B 700 1 600 500 •r. 2 _ 400 5 S 300 z ^ 200 * 100 J *** ♦ *** |
G ♦ ♦ ♦

I " " 04 0 ZI |i '-a Saline SRTG-CNTF (0.5 fig. 2 weeks) SRTG-CNTF (2.5 fig. 2 weeks) SRTG-CNTF (0.5 Mg. 4 weeks) SRTG-CNTF <2.5 Mg. 4 weeks)
♦ Mean 127.27 268.79 455.40 258.93 630.84
i
Figure 4.7. Larger amounts of CNTF loaded into the SRTG system increased the length of survival of RGCs and expression of GFAP to the 4 week time point. After 2 weeks post ONC, both low and high CNTF groups showed a significant difference in the number of remaining Brn3a positive cells to the saline control (A, B, E). After four weeks post ONC, only the high CTNF groups showed a significant difference in the number of Brn3a expressing cells compared to the control (C, D, E). For GFAP expression analysis, after 2 weeks post ONC, both low and high CNTF groups showed a significant difference between the saline control (F, G, J). After 4 weeks post ONC, there was no significant difference between the low CNTF group and saline (H, J). After four weeks there was a significant difference between the high CNTF group and saline group (I, J). Red staining of figures A-D indicates Brn3a expression. Green staining of figures F-I indicates GFAP expression. Photomicrographs were captured at 20x magnification. Scale bar = 50 pm. * p < 0.05, ** p < 0.01, *** p < 0.001.
Discussion
It has been proposed that RGC death in glaucoma may be due to axonal transport failure of NTFs from axon tip to the RGC cell body. As these NTFs are responsible for many crucial cell
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Full Text

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METHODS FOR PROMOTING RETINAL GANGLION CELL NEUROPROTECTION AND AXON REGENERATION by MELISSA LAUGHTER B.S., University of Colorado Boulder, 2012 M.S., University of Colorado Denver, 2014 A thesis submitted to the Faculty of the Graduate School of the University of Colorado in partial fulfillment of the requirements for the degree of Doctor of Philosophy Bioengineering Program 2019

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! ii This thesis for the Doctor of Philosophy degree by Melissa Laughter has been approved for the Bioengineering Program by Kendall Hunter , Chair Daewon Park , Advisor David Ammar Luisa Mestroni Danielle Soranno Date : May 1 8 th , 2019

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! iii Laughter, Melissa (PhD, Bioengineering Program) Engineering an Injectable Microenvironment to Promote Substantial RGC Survival and Axon Extension Thesis directed by A ssociate Professor Daewon Park ABSTRACT Retinal degenerations, such as glaucoma, affect millions of people worldwide and ultimately lead to death of retinal ganglion cells (RGCs) ultimately re sulting in blindness. Unfortunately RGCs do not regenerate on their own making cell transplantation necessary to restore function and vision to these patients. Cell transplantation therapies to treat these diseases are extremely complex; to begin, the cell s must first survive the implantation process in which they are injected into a diseased and damaged environment. Once implanted the replacement cells must then navigate their way into the ganglion cell layer and extend functional axons. During developmen t RGC axon extension is guided by a series of neurotrophic factors and guidance cues; however, these factors are typically not expressed during adulthood. Considering the unfavorable environment and lack of an adult modality, successful cell transplantatio n would require a multifunctional environment to be implanted along with the cells to promote survival and implantation. Herein, we developed a reverse thermal gel system capable of encapsulating cells while being functionalized with a wide variety of moie ties to cater to specific applications. We first functionalized this polymer with a cell binding motif (RTG) intended to promote survival and axon extension of encapsulated cells. Furthermore we tailored this polymer system to possess mechanical and morpho logical properties similar to that of native retinal tissue. 3D culture and analysis of RGCs grown within this polymer system showed increased survival and axon extension when compared to controls. Next we conjugated the reverse thermal gel system

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! iv system w ith charged sulfonate groups intended to promote prolonged release of NTFs from the polymer system. CNTF was loaded within the sulfonated polymer and injected into the vitreous of an optic nerve crush rat model. The sulfonated polymer system loaded with CN TF showed increased neuroprotection following optic nerve crush injury for up to 4 weeks when compared to the controls, indicating a prolonged release of CNTF from the polymer system. In the end, treatment for glaucoma will likely require a multifaceted ap proach due to the complexity of the disease state. However, in this work we were able to show successful in vivo and in vitro results through different manipulations of the same base reverse thermal gel system. Future work would involve combining these two functionalities in the same polymer for a multifaceted approach to the treatment of glaucoma. The form and content of this abstract are approved. I recommend its publication. Approved: Daewon Park

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! v ACKNOWLEDGEMENTS I would like to thank all the people that contributed to the development of this work. I am extremely grateful to all my mentors, professors, family and friends that helped me achieve this goal. This experience has helped me not only define my interests and passions but has been am azing in terms of personal development. First of all I would like to thank Dr. Daew on Park who has supported me throughout this whole project and from whom I ' ve learned a tremendous amount. I will be forever grateful for the opportunities you have given me and hope to continue to work with him in the future. At the same time, I would like to thank my committee members, Dr. David Ammar, Dr. Danielle Soranno, Dr. Luisa Mestroni, and Dr. Kendall Hu nter for their encouragement and advice. Similarly, I would like to thank Maria Bortot, Lindsay Hockensmith, Anna Laura Nelson, Madia Stein, James Bardill, and Adam Rocker for sharing this experience with me and always making lab a fun place to be . I would also like to thank my friends in the department of ophthalmology for their support throughout my work. Their guidance and advice were an invaluable and essential part of this work. Finally, I dedicate this thesis to my family, to my dad for being so insp iring and encouraging me everyday to have goals and motivate me to strive for greatness. To my mom who has been with me every step of the way and who has supported me with everything I do. To my sister, who I share every moment with and who has believed i n me immensely. She has always given me the confidence to continue pursuing my dreams. ! ! !

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! vi TABLE OF CONTENTS CHAPTER I. INTRODUCTION ................................ ................................ ................................ .............. É 2 II. SYNTHESIS AND CHARACTERIZATION OF A BIOMIMETIC REVERSE THERMAL GEL FUNCTIONALIZED WITH INTEGRIN BINDING RGD MOTIFS FOR 3 DIMENTIONAL RETINAL TISSUE ENGINEERING ................................ ................................ ................................ ................... 40 III. A SELF ASSEMBLING INJECTABLE BIOMIMETIC MICROENVIRONMENT ENCOURAGES RETINAL GANGLION CELL AXON EXTENSION IN VITRO .......... 60 IV. AN INJECTABLE NEUROTROPHIC FACTOR DELIVERY SYSTEM SUPPORTS THE SURVIVAL AND REGNERATION FOLLOWING O PTIC NERVE CRUSH INJURY ................................ ................................ ................................ ... 77 V. DISCUSSION LIMITATIONS AND FUTURE DIRECTIONS ................................ ....... 115 REFERENCE S ................................ ................................ ................................ ............................ 129 APPENDIX A. GPS analysis of PSHU ................................ ................................ .............................. 145 B. HPLC standard cu rve and quantification of RGD conjugation ................................ . 147 C. L oss modulus of RTG RGD and RTG ................................ ................................ ...... 148 D. M aximum intensity 3d fluorescent images of RGCs cultured in RTG ..................... 149 E. S ynthetic route of RTG and SRTG ................................ ................................ ........... 150 F. C haracterization of RTG and SRTG ................................ ................................ ......... 151 G. I ntravitreal inje ction cont r ol ................................ ................................ ...................... 154

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! vii LIST OF FIGURES FIGURE 1.1 S chematic showing the effects of rgd integrin binding on cellular behavior .............. 21 1.2 P olymer solubility behavior at the lower critica l solution temperature ...................... 24 2.1 S chematic synthesis of PSHU PNIPAAm RGD ................................ ........................ 41 2.2 NMR spectrum of PSHU ................................ ................................ ............................. 51 2.3 NMR spectrum of PSHU and deprotected PSHU ................................ ....................... 52 2.4 FT IR of DPSHU, PSHU, and PSHU RGD ................................ ................................ 53 2.5 T emperatur e dependent phase transition of PSHU PNIPAAM and PSHU PNIPAAM RGD ................................ ................................ ................................ ............. 55 2.6 T hermal g elling property of PSHU NIPAAM RGD ................................ ................. 56 2.7 SEM images of PSHU PNIPAAM RGD and PSHU PNIPAAM ............................... 57 2.8 R heological properti es of PSHU PNIPAAM RGD and PSHU PNIPAAM ............... 58 3.1 S chematic of 3D RGC culture process ................................ ................................ ........ 67 3.2 M axi mum intensity projections of RGC s cultured within 3D poly m er scaffold ........ 71 3.3 Li ve/dead staining of RGCs 3 days of culture on PDL laminin glass coverslips ....... 72 3.4 Q uantificatio n of average axon length for RGCs ................................ ........................ 73 4.1 E xperimental protocol and animal groups ................................ ................................ ... 91 4.2 E xpression of BRN3a in the retina following ONC ................................ .................... 94 4.3 E xpression of BRN3a in retina following ONC and SRTG intravitreal ..................... 99 4.4 I ntravitreal treatment of SRTG CNTF induced RGC axon growth .......................... 104 4.5 Images of GAP 43 immunostained on downstream from crush site ......................... 105 4.6 Response of Muller glial cells to ONC after intravitreal i njection ............................ 107 4.7 L arger amounts of SRTG CNTF system increased the survival of RGCs ................ 109 5.1 S ubretinal injection of RTG mixed with GFP ................................ ........................... 124

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! viii 5.2 F latmount retinas stained with BRN3, B Tubulin, and dapi ................................ ..... 127 5.3 C ross sections of retina showing integration of transplanted RGCS ........................ 128 A.1 GPC refractive index and light scattering curves for PSHU ................................ ..... 145 A.2 GPC refractive index and log mw indicating polymer size separation ..................... 145 B.1 HPLC curve of mola r ratio of RGD/free amine group ................................ .............. 147 C.1 R heological properties of PSHU PNIPAAM RGD and PSHU PNIPAAM ............. 148 D.1 M aximum intensity projections of representative 3d fluorescent images of RGCS cultured inside PSHU PNIPAAM ................................ ................................ . 149 E.1 S chematic of synthesi s process for SRTG ................................ ................................ . 150 F.1 E lemental analysis of SRTG and RTG ................................ ................................ ...... 151 F.2 T emperature dependent phase transiti on of SRTG and RTG ................................ ... 152 F.3. I n vitro releas e profile of CNTF released from SRTG and RTG .............................. 153 G.1 E xpression of BRN3A in retinal cross sections following ONC and SRTG ............ 154

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! ix ABBREVIATIONS ANOVA analysis of variance BDNF brain derived neurotrophic factor BRN3A brain specific homeobox/pou domain protein BSA bovine serum albumin CNS central nervous system CNTF ciliary neurotrophic factor DMF dimethylformamide DMSO dimethyl sulfoxide ECM extracellular matrix EDC n (3 dimethylaminopropyl) n ethylca rbodiimide hydrochloride FT IR fourier transform infrared GFAP glial fibrillary acidic protein GPC gel permeation chromatography HDI hexamethylene diisocyanate IOP intraocular pressure LCST lower critical solution temperature MMP matrix metalloprotein ase MW molecular weight MWCO molecular weight cut off NHS n hydroxysuccinimide NMR nuclear magnetic resonance NSC neural stem cells NTF neurotrophic factor ONC optic nerve crush

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! x OPP ocular perfusion pressure PBS phosphate buffered saline PNIPAAM poly( n isopropylacrylamide) PSHU poly(serinol hexamethylene urea) RGC retinal ganglion cells RTG reverse thermal gel SEM scanning electron microscopy SRTG sul pshu pnipaam

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! ! ! ! 2 CHAPTER I INTRODUCTION Overview Glaucoma, amongst other retinal degenerations and optic neuropathies, is as it stands an incurable disease. Glaucoma refers to a group of diseases characterized by neuropathy of the optic nerve and loss of retinal ganglion cells (RGCs), the projec tion neurons located in the retina with axons extending through the optic nerve 1 . The death of these cells causes loss of a patient's visual field and in the worst cases can leave a patient completely blind. The shock of losing vision can have devastating effects on a person's quality of life and with Glaucoma being the second leading cause of blindness worldwide 2 (second to cataracts), it is a problem in desperate need of a solution. This crux of this work focuses on taking key steps towards a cure for this disease. T here are many risk factors that increase the likelihood of developing glaucoma; however, increased intraocular pressure (IOP) is recognized as the principal risk factor 3 . This increa se in IOP can be caused by excessive aqueous production, inadequate aqueous drainage, certain medications, trauma, and other eye conditions. In addition, factors such as race, age, and family history can play a role in increased IOP. Treatments are availab le to lower IOP (medications, laser surgery, microsurgery) and currently these are the only proven treatments for glaucoma. However, the enigma behind glaucoma is that although raised IOP remains the most reliable risk factor, many glaucoma patients develo p this disease under normal IOP. Conversely, the majority of eyes with elevated IOPs do not develop glaucoma 4 . Furthermore, a significant number of glaucoma patients see disease progression despite early diagnosis and successful lowering of IOP. This means that patients can suffer significant vision loss even when the appropriate treatments

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! ! ! ! 3 are applied. In addition, there are some cases in which patients respond negatively to drugs or surgeries aimed at lowering IOP. In these cases, patients will see a more rapid disease progression and RGC loss 5 . The bottom line is that many patients still suffer significant vision loss with current glaucoma treatments and the RGCs that are lost are unable to regenerate meaning that the vision lost is currently irreversible. Patho physiology of glaucoma Glaucomatous optic neuropathy is characterized by damage to and subsequent loss of RGCs. This damage typically occurs slowly, taking many years to develop. Although there is much debate as to the mechanism of primary and secondary in sult, the end result is consistent, permanent loss of the neurons vital for sight. In this section, we will review the current topics in the pathophysiology of glaucoma, focusing on the steps leading to neuronal loss. The pathophysiology of glaucoma is mo st likely multifactorial with various risk factors leading to damage of both the RGC soma and RGC axons. The complex nature of this disease has left many unknowns in the exact development of RGC death. Theories for primary injury in glaucoma begin with var ious risk factors including, elevated IOP and fluid dysregulation. Both of these factors can contribute to the initial insult seen in glaucoma, damage to the RGCs through variations in fluid flow and connective tissue. The secondary insult can be thought o f as the damage resulting from injured neurons 6 . Once damaged, injured RGCs release a variety of cytokines and reactive species leading to a vicious cycle of cell death. 7 Following damage to RGCs in glaucoma, cell death occurs through the process of apoptosis 8 9 . Apoptosis, also known as programmed cell death, is a mechanism of cell death in the absence of inflammation. This is because in apoptosis, cell death occurs without lysing of the cell and thus without the release of inflammatory factors. Instead apo ptotic cells die through characteristic

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! ! ! ! 4 DNA fragmentation, chromatin condensation, formation of apoptotic bodies, loss of mitochondrial membrane potential, and the breakdown of the cell into multiple smaller membrane bound vesicles 6 . These vesicles are then cleared or engulfed fr om the area by neighboring phagocytic cells. Although there is agreement that RGC death in glaucoma occurs mostly through apoptosis, there is some evidence that shows necrosis plays a part in the late stages of this disease 10 . The process of apoptosis is heavily regulated and progresses in an orderly fashion through a series of signals. This signally cascade is modulated by caspases, a family of aspartate specific cysteine proteases and members of the interleukin 1 B converting enzyme family. Due to the destructive capabilities of these enzymes, caspases initially exist as inactive zymogens. Once cleav ed, they work to initiate, propagate, and amplify the apoptotic process. There are three major pathways of apoptotic progression in mammals: the death receptor pathway (extrinsic), the apoptosome pathway (intrinsic), and the cytotoxic lymphocyte initiated granzyme B pathway 11 . The extrinsic pathway is initiated when the Fas ligand (FasL) binds to the Fas receptor (death receptor) present on the cell's outer membrane. This binding triggers a signally cascade ultimately leading to the activation of caspase 8 and caspase 10. These downstream caspases work to digest the cellular contents and begin apoptosis 12 . There is ev idence that in humans, glaucoma progresses mostly through the extrinsic pathway. An increase in the expression of FasL and Fas receptor are seen throughout glaucoma progression and are also associated with increased IOP (a strong risk factor for glaucoma) 13 14 . Elevated IOP can also cause an increase in TNF alpha lending to further proof that the extrinsic pathway of apoptosis is highly involved in glaucoma. Due to this, researchers have begun to develop methods to intercept t his pathway. Specifically antibodies against TNF alpha have been shown to block some of the subsequent

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! ! ! ! 5 RGC damage in mice with ocular hypertension 15 . In addition, one study delivered caspase inhibitors to rats that had undergone axotomy of the optic nerve. This treatment recused 34% of RGCs that would have otherwise died 14 days post injury 16 . Elevated intraocular pressure as a risk factor in glaucoma RGC death in glaucoma commonly occurs in the presence of increased IOP, making this a very important risk factor for disease progression 17 . Several studies have shown the association between elevated IOP and RGC loss. More specifically, investigators have shown a positive correlation between a change in IOP and RGC death. Furthermore, it has been shown that the duration of increase in IOP can determin e the significance of RGC death 18 . However, among glaucoma patients, only one third of patients have increased intraocular pressure a t the early stages. Furthermore 30 40% of patients experiencing visual fie ld loss due to glaucoma, have an intraocular pressure within normal range 19 . Thus, while increased IOP is still considered an important risk factor, it is by no means the only player. Three factors that determine IOP are: 1. The degree of venous pre ssure 2. The rate of aqueous production by cilliary bodies, and 3. The rate of aqueous drainage across the trabecular meshwork schlemm's canal system. Generally elevated IOP results from the third mechanism, impaired drainage of aqueous humor due to chang es in the trabecular meshwork. Most of the aqueous humor in the eye drains through the trabecular meshwork into the schlemm's canal and empties into the collecting veins on the scleral surface. The trabecular meshwork is made of layers with one layer compo sed of connective tissue and the other of endothelial cells. As we age, the endothelial layer begins to diminish and the connective tissue layer begins to thicken. Thickening of the connective tissue increases resistance and impedes aqueous drainage thereb y prompting an increase in IOP. By

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! ! ! ! 6 definition, elevated resistance to outflow at the level of the trabecular meshwork is primary open angle glaucoma where the iridocorneal angle remains anatomically open. On the contrary, angle closure glaucoma occurs as a result of elevated resistance building between the front of the lens and the pupillary margin. This increase in pressure pushes the iris up blocking the iridocorneal angle and aqueous drainage. Mechanisms of retinal ganglion cell apoptosis in response to increased IOP Unfortunately, it is still not well understood how this increase in IOP leads to RGC apoptosis. One proposed mechanism is that increased IOP alters the cellular ECM through extensive collagen remodeling and the presence of matrix metalloprot einase (MMP) 1 20 21 22 . MMPs are the enzymes responsible for the degradation of ECM. In a recent study, the exposure of eyes to increased IOP caused an increase in the release of MMPs into the ECM. Digestion and damage to the ECM by these MMPs can lead to disruption of RGCs and subsequent cell death. It is known that cell ECM and cell biomolecule interactions are crucial for the survival of most cells. Specifically inte grin binding motifs present in many components of the ECM, are involved in supporting attachment, spreading, and survival of neural cell types 23 . Thus, the degradation of the ECM caused by the release of digesting MMPs, could compromise these crucial cellular attachments leading to RGC apoptosis. The second explanation is that elevated IOP increases the mechanical force exerted on RGC axon s passing through the lamina cribrosa of the optic nerve. Damaged RGCs begin to secrete MMPs, which in turn causes digestion and damage to the ECM further perpetuating this disease state 4 . Finally, the damage of RGC axons can hinder retrograde transport of critical growth factors. When these growth factors are no longer able to make it through the RGC axon and up to the cell body, the cell may no longer be able to regulate

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! ! ! ! 7 metabolism and cell survival leading to apoptosis (this will be discussed further in subsequent sections) 24 . Vascular insufficiency Undoubtedly, elevated IOP likely plays a huge role in the development of glaucoma. However, this theory does not account for the large number of patients that develop glaucoma under the normal range of IOP. In addition one study reported that over 90% of people that had high IOP failed to develop glaucoma ov er a 5 year period. Similarly, there are many patients that develop glaucomatous neuropathy with a normal IOP. It was also noted that patients that have glaucoma and increased IOP continued to show disease progression even when their IOP was reduced to a n ormal range 25 . Thus, there are se emingly other risk factors involved that contribute to the development of this disease. Recent works point towards a correlation between the presence of vascular insufficiency and glaucoma. Vascular insufficiency includes reduced ocular perfusion pressure and defective vascular regulation. Ocular perfusion pressure (OPP) describes the relationship between blood pressure (BP) and IOP as shown in the following equation: OPP = BP Ð IOP Studies have shown that a decrease in OPP is a risk factor for glaucoma a nd that systemic hypertension may in fact be protective against the development of glaucoma (likely because this would increase the OPP) 26 . Defective vascular regulation can also contribute to the formation of glaucoma. Constant blood flow is necessary to provide oxy gen and nutrients to the eye, most importantly the retina and optic nerve. In a healthy eye, the maintenance of constant blood flow is controlled through autoregulatory mechanisms existing within the blood vessels. Unfortunately, as you age, these mechanis ms worsen and the ability to adapt to fluctuations in

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! ! ! ! 8 pressure decreases. In one study, researchers showed that older rats were unable to adapt to changes in ocular pressure as compared to younger rats 27 . Thus, these defects in autoregulation that occur with age could lead to ischemia within the eye and damage to the RGCs. T reatments for glaucoma The first question to ask when devising a treatment for glaucoma is, how do RGCs die in glaucoma? It is now agreed that RGC death in glaucoma occurs via apoptosis or necrosis 28 (as discussed above) . Apoptosis is the active process of programmed cell death. This process may initiate when the cell is deemed no longer necessary or b ecomes damaged in some way 29 . On the contrary, RGC necrosis occurs when surrounding toxins injure the cellular membrane and destroy the cell. This process is pa ssive in nature and can lead to an inflammatory response, which can contribute to more cell death. Although both methods of cell death can occur during glaucoma, apoptosis is known to be the primary method of cell death with necrosis occurring during the l ate phase of the disease 30 . New therapeutic approaches Until recently, most therapeutic approaches for glaucoma were focused on lowering IOP. However, as mentioned above, a large portion of these patients continue to show disease progression even when their IOP is brought down to normal conditions. In addition, there is a significant portion of patients that develop glaucoma under normal IOP ra nges. As new information on the pathophysiology of glaucoma began to emerge, new treatments were developed that focused on these mechanisms instead of solely lowering the IOP. One new group of therapeutic agents are aimed at preventing apoptosis of the RG Cs. As discussed above, the process of apoptosis is mediated through enzymatic proteins called caspases. Caspase inhibitors, such as erythropoietin, have shown efficacy in preventing apoptosis in a glaucoma rat model 31 .

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! ! ! ! 9 However, apoptosis is the final step in the cellula r progression of glaucoma and thus these treatments act far downstream to where the problems is initiated. Other agents, such as N acetyl L cysteine, NOS inhibitors, and N methyl D aspartate (NMDA) antagonists, have shown promising results at preventing RG C death in glaucoma animal models 32 16 . NMDA antagonists showed a lot of potential at neuroprotecting RGCs in animal models and were subsequently taken to clinical trials. NMDA is a glutamate receptor that, once activated, leads to the activation of ion channels and the influx of Ca 2+ and Na + into the neuron. The influx of these extracellular ions works to trigger a signaling cascade that leads to cell death 33 34 . Thus, NMDA antagonists prevent the activation of the NMDA receptors and the resulting cell death. However, once in clinical trials, it was quickly noted that these antagonists not only block the NMDA receptors in the eye but all NMDA receptors essential for neuronal function throughout the body 35 . Neurotrophic factor treatments for glaucoma The second question to ask when investigating RGC death as a result of glaucoma is, what causes the RGCs to die? The answer to this question is unfortunately much more complex, with likely several molecular pathways contributing to RGC cell loss. However, recent progress in this field has increased our underst anding of the pathways that lead to RGC degeneration following optic nerve injury. The answer to this question is now believed to lie with neurotrophic factors (N T F) and their role in a healthy retina. Neurotrophins are a family of small diffusible molecul es that are implicated in the control of adult neurogenesis, axon extension, proliferation, and cellular survival 36 . Two NF exhibiting key roles in the survival of RGCs are brain derived neurotrophic factor (BDN F) and ciliary neurotrophic factor (CNTF). In healthy eyes, BDNF is produced in the lateral geniculate body where it binds to its receptor on the RGC axon and moves to the RGD cell body through retrograde transport 37 . When BDNF reaches the RGC cell

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! ! ! ! 10 body , it binds to its corresponding receptors to affect survival of RGCs 24 . On the contrary, CNTF is released as a response to diseased conditions. Early studies first identified i ncreased CNTF following injury of retinal neurons, a fter axot omy, ischemia, and experimental glaucoma 38 39 . However, CNTF works to provide a similar result to BDNF by induci ng neuroprotection of the RGCs. As mentioned above RGC axons are responsible for transmitting signals from the photoreceptors to the cortex of the brain; however, RGC axons are also responsible for the uptake and retrograde transport of distant N T F to the cell body 1 . Although cells within the retina also produce N T F, studies have shown that both sources of N T F may be important in preserving the survival and function of RGCs 40 . The hypothe sis that RGC death occurs due to Ôaxonal trans port failure' derives from this knowledge. As glaucoma proceeds, RGC axons get damaged decreasing axonal transport of vital N T F s . Insufficient or unbalanced levels of N T F can cause RGCs to go into an apoptotic cascade. Evidence for this hypothesis stems from studies showing decreased in retrograde axonal transport following glaucoma like optic nerve damage 41 42 . Additionally, increased IOP a nd subsequent optic nerve damage has been correlated with retinal accumulation of dynein, a motor prote in required for retrograde axonal transport 43 . Although the Ôaxon al transport failure' hypothesis and Ôneurotrophic factor deprivation ' hypothesis have not been completely proven, significant evidence points to these reasons for the loss of RGCs in glaucoma. Neuroprotection of retinal ganglion cells Recent discoveries i n the mechanisms behind RGC death in glaucoma have le d to new potential treatmen ts. The first line of treatment for glaucoma have been aimed at neuroprotection of RGCs. Neuroprotection is a therapeutic strategy aimed at preventing RGC death and loss of function 18 .

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! ! ! ! 11 Instead of traditional preventative measures to treat glaucoma, such as decreasing IOP, neuroprotective strategies focus on guarding RGC survival and function. Neurotrophin su pplementation strategies have been one area of research extensively studied for RGC neuroprotection. BDNF and CNTF have been the primary N T Fs used for these neuroprotection strategies. Initial data supporting these therapies have involved optic nerve inju ry models (e.g. optic nerve crush or transection model) 18 . Although optic nerve crush (ONC) in a rodent shows a different progression from glaucoma in a human , it does provide researchers with the ability to study RGC death during optic neuropathy. In addition, this animal model has been able to provide an effective model for optic nerve trauma as well as regeneration failure. Furthermore, the optic nerve is highly accessible increasing the consistency of the model and making it an effective tool to investigate new treatments. Studies have been able to show repeatable results in RGC survival following nerve crush with a predictable number of RGC s remaining at various time points. For instance, during the first week following optic nerve crush there is a small difference in RGC cell count (depending on distance between crush site and optic cup) 44 . After t he first week post nerve crush, changes in RGCs occur much more rapidly. A significant number of RGCs die after the first week reducing the number of RGCs in the retinal layer to 10%. This abrupt cell death allows researchers to see differences in RGC survival between treatment and control groups. For instance, studies using an optic nerve crush or transection animal model followed by intraocular injec tion of BDNF or viral mediated BDNF gene transfer using adenovirus have been able to show an increase in RGC survival at just one week post lesion (100% survival of RGCs in treatment group compared to 50% in control groups). Two weeks post lesion, similar treatments have been able to rescue 48% of RGCs compared to the 10% survival rate in the control eyes 45 46 47 . Although BDNF does show a

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! ! ! ! 12 profound effect on RGC survival following injury, it unfortunately has little effect on regenerating axons. Previous studies h ave shown that intravitreal injection of BDNF following optic nerve crush did not increase RGC axon regeneration past the lesion site. In fact, it is now believed that BDNF may have an inhibitory effect on RGC axon regeneration 48 49 . Pernet et al showed that not only did the application of endogenous BDNF fail to stimulate axon regeneration, but it also led to hypertrophic axonal swelling potentia lly causing further damage to segments of the optic nerve 48 . Furthermore, neuroprotective strategies using BDNF (either via recombinant protein or g ene transfer) have failed to extend the time course of RGC survival. Groups have attempted repeated intravitreal injections of BDNF as well as minipumps to provide sustained N T F delivery, but have not seen RGC survival past certain time points 50 . This result stems from decreased expressio n of the BDNF receptor (TrkB) following RGC injury. The decreased amounts of the TrkB receptor from cellular internalization limit s the ability of RGCs to respond to neurotrophin stimulation 51 45 . Ciliary neurotrophic factor Similar to BDNF, CNTF has a neuroprotective effect on RGCs following optic nerve injury. Administration of exogenous CNTF following optic nerve axonomy displayed increas ed survival of RGCs although not as increased as BDNF 52 . However, therapies with CNTF using gene transfer appear to be more effective at preserving RGC survi val post injury. These studies showed a modest 25% to 50% survival of RGCs (compared to 10% control eyes) two weeks post lesion 53 . Although CNTF doesn't have as great an effect on RGC survival as BNDF, CNTF has been shown to stimulated RGC axon regeneration post injury. Studies using AAV mediated CNTF gene transfer following optic nerve crush stimulated significant RGC axon regeneration 54 . This information indicates that BNDF and CNTF operate using different signaling pathways and

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! ! ! ! 13 therefo re stimulated different cellular responses such that BDNF is more effective at stimulated neuroprotection and CNTF is more effective at stimulating axon growth. Neuroprotective strategies involving N T F can be divided into three categories, ones that supply RGCs with exogenous N T Fs, ones that use genetic modifications to stimulate the cellular expression of N T Fs, and finally ones that use genetic modifications to increase the expression of N T F receptors (such as TrkB). As mentioned above, N T F receptors can become downregulated following optic nerve lesion 51 . This can lead to RGC death by making the cells less responsive to N T F survival signaling. Different strategies have been used to upregulate these receptors in hopes of increasing the potency of surrounding N FT s and thus promoting RGC cell survival 55 . Most strategies use vectors to transfer g enetic information coding various receptors. Additionally, strategies have been investigated to increase endogenous expression of N T F from the RGCs themsel ves. Similar to strategies used to increase receptor expression, these strategies include viral vecto r transfection of genes coding N T F into RGCs 56 . Although promising results have been seen in optic nerve crush animal models, both of these strategies include variable genetic alterations that may be too complicated for current clinical applications. Lastly, direct injection into the vitreous has been used to deliver N T F s to damaged RGCs. As mentioned above, these strategies have shown prolonged survival of RGCs as well as some axon regeneration after optic nerve crush or axonomy 57 58 . Although this approach is minimally invasive and does not require any genetic manipulations, this method pr ompts diffusion of the protein out of the intended treatment area and quick proteolytic degradation. Therefore, applying this method in the clinical setting would require multiple and frequent d oses of the GF but would still not likely result in the full e ffect 59 .

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! ! ! ! 14 Cellular scaffolds for retinal ganglion cell regeneration As it stands, RGC loss from any optic neuropathy is irreversible. This means that for many patients suffering from glaucoma, the loss of one's visual field is permanent and cannot be improved. The only treatments that are available are preventative measures aimed at reducing IOP and protecting surviving cells from death. For many individuals, these preventative treatments are not sufficient and do nothing to imp rove their quality of life. For these cases the only hope of restoring vision is optic nerve regeneration; however, this goal presents immense challenge. Firstly stem cells induced to differentiate into RGCs must be sour ced to replace the depleted cells . I t is important that these cells are delivered to the appropriate area and integrate successfully into the surrounding tissue. Unfortunately, transplantations consisting solely of cells will be hindered by a diseased microenvironment characterized by growth inhibitory signals. In glaucoma, the diseased state consists of myelin associate molecules, scar formation, and lack of passage across a lesion, all which need to be mitigated to allow successful cell transplantation 60 . The environment of transplantation needs to be made permissive to cell survival and axon regeneration. Peripheral nerve grafts Peripheral nerve (PN) grafts have been used to promote survival of injured RGCs (when transplanted into the vitreous) or to prom ote RGC axon regeneration through the lesion (when grafted to the optic nerve). Studies have shown that implanted fragments of PN into the vitreous can stimulate RGCs to regenerate axons across the lesion site 61 . However, these results are mostly attributed to the release of N T F from the Schwann cells and the presence of macrophages that can enhance axon regeneration. Furthermore, the use of Ôconditioned' or pre crushed grafts increase RGC survival significantly more than normal grafts due to the increase in release of

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! ! ! ! 15 N T F 62 . PN s can also be grafted to the end of a transected optic nerve 63 . Studies have shown that RGCs are capable of regenerating axons through the grafted PN extending all the way to the superior colliculus and forming correct synaptic connections 64 65 . The ability of RGCs to regenerate axons through the grafted PN is likely due to the increased amount of cell adhesion molecules and decreased amoun t of growth inhibitory molecules present within the graft 66 . Stem cell replacement therapies Although regenerating the optic nerve seems like a very distant goal, stem cells and tissue engineering hold great potential for achieving this goal. As mentioned above, the loss o f RGCs due to glaucoma or other optic neuropathies are currently permanent. Therefore, it is necessary to treat such diseases with cellular replacement therapies to replace the depleted RGCs and possibly reverse the vision lost. Stem cells have garnered si gnificant attention due to their ability to self renew and differentiate into multiple cell types, making them capable of replacing specific types of tissue. For this work, we are concerned with the differentiation of stem or precursor cells into RGC like cells. Preliminary studies have investigated whether implanted cells can differentiate into an RGC like phenotype 67 68 . To do this, groups have used histological markers specific for RGCs to determine effective differentiation of the stem cells. Two of the most common immune markers are B 3 tubulin and brain specific homeobox/POU domain protein 3 (Brn3). Furthermore, studies have been conducted to determine the best place for cell transplantation. The first and most simple injection site investigated was the vitreous cavity. Studies have shown varia ble results depending on the type of stem cell transplanted. One study showed that neural stem cells (NSCs) injected into the vitreous were unable to integrate into the retinal layer and efficiently differentiate into RGCs 69 . However, other groups have demonstrated that retinal stem cells (RSCs) are more capable of incorporating into the retinal layer following intravitreal

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! ! ! ! 16 injectio n. Regardless, few stem and precursor cell types injected intravitreally differentiate efficiently into RGCs and incorporate into the retinal layer 70 . Transplantation can also be done using subretinal implantation. Such studies have shown that RSCs integrate better when implanted into the subretinal layer when compared to intravitreal injection 71 . However promising these early results are, cell transplant therapies are still plagued by a low rate of cell survival and limited integration of cells into target tissue. These results are likely caused by the injured microenvironment the cells are exposed to upon injection. The microenvironme nt or Ô niche' can have a severe effect on cellular behavior by exposing cells to various spatially and temporally controlled biochemical cues, cellular attachment ligands, and extracellular matrix (ECM) molecules. In addition, even the microenvironment's m echanical properties and topographical cues can illicit specific cellular behavior. Accordingly, cell implantation into an injured microenviroment that has been compromised by growth inhibiting myelin associate d molecules, absence of growth permissive mole cules (e.g, laminin), scar formation, and lack of passage across a lesion is extremely detrimental to implanted cells 72 . One way to mitigate these limitations is through the use of biomaterials to ac t as two dimensional 2D or three dimensional 3D cellular scaffold . These scaffolds not only provide physical support for the implanted cells but also can be engineered to possess specific mechanical properties and biochemical cues. Therefore, these biomaterials provide the exciting possibility of constructing an alternative microenvironment that can instead encourages cell survival, growth, and in our case axon extension. Biomaterials for controlled delivery of alternative microenvironment Not only can biomaterials be used to aid in the delivery of replacement cells, but they can also be injected into a diseased environment to provide an alternate bioactive niche. In this case the

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! ! ! ! 17 biomaterial would be used to alter the disease microenvironment to augment endogenous stem cell activity and subsequent regeneration. This can be achieved by incorporating biomolecules within the scaffold that can be released into the surrounding area or through the chemica l conjugation of biomolecules. Both methods can lead to activation of endogenous stem cell populations and potentially to the self renewal and differentiation into a replacement cell population. Studies with transplantation of melanoma cell adhesion molecu les (MCAM) have been able to show the activation of endogenous human bone marrow stroma and subsequent differentiation into bone cells 73 . Furthermore, in previous studies we have show n that a novel synthetic biomimetic polymer (PSHU RGD) can induce significant differentiation of human neural stem cells (hNSCs) into a population of human motor neurons (hMNs) in 2D. Designing biomaterials for a 3D scaffold/microenvironment Analysis of c ellular behavior on 2D materials has been the first step to engineering an appropriate cellular scaffold. 2D materials allowed for controlled experimentation to determine how cells interact with individual components of a cellular niche 23 . However, 2D approaches are limited to preliminary in vitro studies in that they are unable to represent natural tissue. Unlike 2D systems, 3D materials are more in tune with the cell cell interactions, cellular organization, and microenvironment s een in native tissue. In 3D systems, cells will be fully encapsulated by a solid microenvironment and thus be more likely to respond to the s caffold specifically engineered to induce positive cellular behavior. Fabricating a 3D scaffold is much more than just creating a 3D structure. Instead, t here are many considerations depending on the desired cell behavior. Each characteristic and component of the 3D scaffold will have some sort of effect on cellular behavior so each property must be considered and eng ineered appropriately. Not only do 3D scaffold s require the same

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! ! ! ! 18 considerations as 2D scaffold s (e.g. biomolecule density, materia l wettability, material charge) but also 3D scaffolds bring new complications such as mechanical properties (elastic ity) and m orphology ( porosity). In addition, practical considerations need to be taken when designing a 3D scaffold. First, the biomaterial must allow embedding of cells and ideally encapsulate the cells in situ. Typically, this is done through crosslinking of the mater ial after implantation. However, this has to be well controlled to ensure harmful side products produced from the reaction do not damage the encapsulated cells or the surrounding tissue. Second, how the biomaterials will be delivered to the diseased a rea is an other important consideration. Perhaps the ultimate goal is to create an injectable material that allows facile encapsulation of cells and gelation in situ. In addition, this characteristic would provide a more minimally invasive method less likel y to further dam age already damaged and possibly necrotic tissue present in a diseased state 74 . Biomolecule modification of biomaterials and cell behavior Surface modifications of biomaterials provi de researchers with a way to fine tune and manipulate materials to possess certain characteristics. One of the main goals of surface modifications is to make a material biomimetic. Biomimetic polymers combine the controllability of a synthetic polymer with the biocompatibility of natural material found in the body. Biomimetic polymers can be manipulated to prompt a specific cellular response or direct new tissue formation. Extensive work has been conducted to create biomimetic polymers suitable for cell tr ansplantation therapies. One of the most prominent methods to create a biomimetic material is through surface modification of common ECM proteins such as fibronectin (FN), vitronectin (VN), and laminin (LN) 75 . However, it was discovered that a small sequence of amino acid s within the larger proteins is responsible for integrin interaction and

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! ! ! ! 19 cellular attachment. It is now more common to simply incorporate the short peptide fragments to the material surface. This is because the larger proteins can present steric hindrance that may prevent cellular attachment. In addition, using smaller peptides can increase the amount of biomolecule that can be attached to the polymer surface thus increasing the cellular response. Finally, s mall peptide sequences are relatively more stable than the full protein and can thus better withstand the chemical conjugation process. Arg Gly Asp (RGD) is the cell binding motif found in fibronectin (FN) and laminin (LN). Because of its extensive presen ce and its highly effective cell binding capacity, RGD has become the most commonly employed peptide used to enhance cell attachment to a polymer's surface 76 . RGD is c apable of binding to a plethora of cell adhesion receptors found throughout the body; however, the largest and most common group of RGD binding cell adhesion receptors is the integrin family. This group of adhesion molecules is not only responsible for cel l attachment, but they also play an important role in cell differentiation, embryogenesis, proliferation, and gene expression 77 . Binding of the RGD peptide to integrins causes a cascade of signals that can influence cell behavior in different ways (Figure 1.1 ) 76 . Re search has shown that focal adhesion associated proteins trigger the expression of anti apoptotic protein Bcl 2 and the focal adhesion kinase pathway 78 . These two pathways along with focal adhesion formation play an important role in cell survival, proliferation, cell spreading, and neurite extension 23 . It was also determined that the lack of this interaction between integrins and cell binding motifs can induce apoptosis and cell death. This Ôintegrin mediated death' is caused by the lack of a substrate to bind the ligand leadin g to the recruitment of caspases and the cell death cascade. Relating this finding to a

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! ! ! ! 20 cell's interaction with the ECM, we can see that the absence of integrin binding ligands in a cell's environment can cause the cell to undergo apoptosis 79 . Mechanical prop erties of biomaterials and cell behavior In glaucoma, the diseased tissue is not only associated with cell death and release of toxic cytokines, but also with altered organization of the ECM. This disturbance in the tissue can lead to changes in the mechanical properties or stiffness of the matri x, which most likely contributes to the progression of the disease. It was recently discovered that the mechanical properties of the ECM, in which cells reside, can have an extreme effect on cell behavior and cell fate. In one of the first studies displayi ng this finding, human mesenchymal stem cells grown on polyacrylamide with varying stiffness (typical of brain, muscle, and bone) began to differentiate into the corresponding tissue specific cells 80 . This innovative study highlighted the strong Figure 1.1 . Schematic showing the effects of RGD integrin binding on cellular behavior. Integrin binding can induce cell survival, proliferation, and axon extension.

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! ! ! ! 21 influence scaffold mechanica l properties can have on cells and marked the beginning of many future studies relating matrix stiffness and cell behavior. Following suit, other groups were able to show that mechanical properties have significant control over proliferation and differenti ation of other types of stem cells. One such study showed that adult neural stem cells (NSCs) grown on relatively stiff substrates triggered differentiation into glial cells, whereas adult NSCs grown on softer substrates (more closely resembling the stiffn ess of brain tissue) triggered differentiation into mostly neurons 81 . Compared to these 2D studies, examining mechanical properties in a 3D scaffold i s significantly more challenging. This is due to the presence of other factors (physical, chemical, and mechanical) that are in play in a 3D matrix. Therefore, constructing a suitable 3D scaffold with the appropriate characteristics is complex. With this said, substrate elasticity and materials with mechanical properties closest to that of the native tissue are the most desirable and will likely have the most beneficial effect on surrounding cells 82 . For our purposes, it was important to create a material with similar mechanical properties to native retinal tissue, as this w ould likely improve RGC culture in vitro as well as in vivo results. Rheology studies have been done to assess the mechanical properties of retinal tissue. Results from these experiments show that native retina l tissue is soft with a storage modulus not e xceeding 100 Pa 83 . Thus, a matrix that possesses softer mechanical propert ies and a modulus that does not exceed a couple hundred Pascals would be beneficial to RGC growth and survival. In situ polymer gelling systems Perhaps the ultimate goal of cellular scaffolds is to create a multicomponent, injectable material that can con form to the target tissue and provide a replacement niche. This characteristic is especially im portant when it comes to engineering neural tissue. As stated above, it is important

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! ! ! ! 22 that scaffold s be tailored to match the mechanical properties of the native tissue. However, neural tissue has an extremely low mechanical modulus making implantation of a scaffold nearly impossible. To combat this obstacle, researchers have turned to in situ forming polymers. These polymer systems work by transitioning from a low viscosity fluid to a solid gel upon application of various stimuli. This characteristic allows the polymer system to be deployed through a minimally invasive injection into the target site and conform to the site upon gelation. Not only does this allow fo r the use of polymer systems with an extremely low modulus, but it also allows the entrapment of molecules or cells before application. Initially, researchers used polymer systems that relied on photo initiated free radial formation to provide this in situ gelling behavior. However, the side products of this reaction proved cytotoxic to cells and made this technique no longer feasible for tissue engineering or drug delivery application 84 . Avoiding polymer systems that produce harmful side products upon gelation, researchers turned to stimuli sensitive polymer s that undergo reversible volume phase transition or sol gel phase transition in response to various stimuli. These stimuli m ight include external physical or chemical stimuli such as pH, charge, light, or temperature. Unlike in situ gelling polymers formed by harmful chemical crosslinking reactions, stimuli sensitive polymer s can be reversibly transformed from solution to solid state by altering environmental conditions. Therefore, these stimuli sensitive polymer systems may be polymerized and purified befo re application providing a simpler and safer method for injectable materials. Thermally induced gelling systems Among stimuli sensitive polymer systems, temperature sensitive polymers (thermogels ) are the most widely used due to their facile control and practical advantage for both in vitro and in vivo purposes 85 . These systems work through an entropically driven phase transition upon reaching a

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! ! ! ! 23 specific temperature, also referred to as the lower critical solution temperature (LCST). The LCST of a polymer system is dete rmined by competing hydrophobic and hydrophilic interactions with the polymer backbone. More specifically, below the polymer's LCST hydrophilic interaction s between the backbone and water molecules prevail causing the polymer to go into solution. Above the polymer's LCST, the hydrogen interactions become weaker and hydrophobic interactions between the polymer backbone moieties prevail, causing the water to be expel led from the polymer and form a solid ge l. A Schematic of this phenomenon can be seen in Figure 1.2 . Because this process is controlled through hydrophilic and hydrophobic interactions and not through polymerization, this process is fully reversible. For this reason, these temperature sensitive sys tems are termed reverse thermal gels (RTGs). The general principle behind RTGs, is to construct a block copolymer that has both hydrophobic and hydrophilic segments, named an ampiphilic polymer. This study focuses on one ampiphilic polymer specifically, p oly (N isopropylacrylamide) (PNIPAAm). Figure 1.2 . Polymer solubility behavior at the Lower Critica l Solution Temperature (LCST). Left hand side shows hydrated polymer below LCST with entropic loss of water and right side shows chain collapse above LCST 208 .

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! ! ! ! 24 PNIPAam is a thermosensitive water soluble homopolymer that has garnered a lot of attention in the biomedical field. As described above, this ampiphilic polymer exhibits a re versible hydration dehydration solution t o gel (sol gel) change in response to temperature 86 . In addition, this pol ymer has been shown to exhibit a sharp, reversible sol gel transition sufficiently above room temperature (25 ¡ C ) and sufficiently below body temperature ( 32 ¡ C) 87 making it extremely useful for many biomedical applications 88 . When conjugated with other polymers, PNIPAAm can impart its thermosensitive behavior on otherwise non stimuli sensitive polymers or molecules. For these reasons, PNIPAAm has been employed for many different applications including drug delivery, tissue engineering and, scaffolding applications 89 90 91 . From a drug delivery perspective, PNIPAAm offers significant advantage in that it shows good sensitivity, reversible transition, and low cytotoxicity when compared with other reversible thermal gels 92 . I n this sense, PNIPAAm provides a favorable environment for the entrapment of certain drugs and molecules. In addition, these molecules can be loaded into the polymer system through facile mixing below room temperature and delivered to the site through a mi nimally invasive injection, encapsulating the loaded molecule s and conforming to the injury site upon gelation. Furthermore, the degradation of this polymer can be altered through changing the chemical structure of PNIPAAm, providing control of the release profile of entrapped molecule s 92 . Snowden et al first demonstrated the potential of PNIPAAm as a drug delivery system through investigating the release of fluorescein labeled dextran from the PNIPAAm based microgels 93 . Research ers then moved to show the sustained delivery of other molecules such as insulin 94 and bovine serum albumin (BSA) 95 from PNIPAAm systems . However, studies soon found that PNIPAAm, used as a drug delivery system , is limited by a quick burst release. In one case encapsulated drugs were released from the polymer within the first 24 hrs of deployment 96 .

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! ! ! ! 25 Groups such as Zhang et al . mitigate d this obstacle through manipulating the PNIPAAm based systems. They were able to show controlled releas e of water soluble protein BSA for an extended period of time without a significant burst release. This study gave rise to further studies using PNIPAAm as a drug delivery system for similar water soluble growth factors and neurotrophic factors. Some group s showed the benefits of prolonged release of neurotro phic factors from a PNIPAAm drug delivery system in the regeneration of motor neurons and treatment of spinal cord injury (SCI). One particu lar study showed that six weeks of release of BDNF from a PNIP AAm PEG injectable scaffold encourage d survival and attachment of motor neurons 90 . PNIPAAm has also b een used for many tissue engineering efforts. Typically polymer scaffold s used for tissue regeneration have been comprised of biodegradable polymer systems such as poly (lactic acid) PLA, PLA based copolymers, alginate or collagen. Unfortunately these pol ymers degrade quickly and th us may not provide sustained mechanical support 97 . Disappearance of the scaffold before injured axons are able to regenerate can halt regeneration across the lesion as well as cause an increase in inflammatory response and glial scar formation . Furthermore, the polymer systems listed above all require s urgical implantation. This process can lead to significant damage of an already damaged site. For these reasons, researchers have begun to investigate injectable scaffold s that can conform to the irregular shaped injury site and be deployed through facile injection. Unfortunately most of these polymer systems require crosslinking following injection and can release harmful unreacted chemicals into the surrounding tissue. PNIPAAm was the clear alternative to other injectable polymer systems in that it does not chemica l l y crosslink upon injection . Furthermore PNIPAAm degrades slowly and can thus provide prolonged support for encapsulated cells once implanted 98 99 .

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! ! ! ! 26 Bi omimetic scaffolds for retinal tissue engineering In the final two sections of this introduction, we will discuss the use of injectable materials in both retinal tissue engineering as well as neurotrophic factors/drug delivery for retinal regeneration appl ications. These sections will include the current state of research, notable pitfalls, and an introduction to the research performed in the subsequent chapters. We will focus this discussion on the use of synthetic polymers. With the recent advances in st em cell technology, cell based therapies have an increased potential to provide curative treatments for a wide variety of diseases. For retinal tissue engineering, transplantation of NSCs, RGCs, or RGC progenitors are of particular interest. Solely the del ivery of cell suspensions into the vitreal space have been shown to provide some amount of neuroprotection in neurodegenerative disease models 100 101 , likely due to the release of NTF from the implanted cells 102 . However, the integration of cells injected in this fashion is erratic with most cells incorporating randomly within the target tissue. Thus, the efficacy of transplantation as a t reatment is hindered by poor survival and improper integration of freely injected cells into the damaged tissue 103 . A suitable vehicle to deliver replacement cells to the target tissue can mitigate these issues by providing appropriate biochemical cues and mechanical support to the encapsulated cells. However, traditional cellular scaffolds require surgical implantat ion, which can incur further damage to already, damaged tissue 104 . For this reason, attention has turned to injectable materials that avoid surgical im plantation while retaining all the benefits of a cellular scaffold. Various forms of ECM technology have emerged capable of delivering transplanted cells in an injectable scaffold. One commonly used platform for neural tissue engineering has been injectabl e hydrogels. These materials are highly tunable, capable of providing various mechanical and biochemical cues to encourage cell survival, deliver drugs, or

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! ! ! ! 27 provide specific mechanical support. Although promising, many of the injectable hydrogel scaffolds r equire in vivo polymerization giving way to potentially harmful side products that may negatively affect implanted cells and surrounding tissue. Furthermore, in situ polymerization limits the controllability of each polymerization reaction because the body can be an extremely variable environment. Polymers that remain injectable but can be polymerized prior to injection would provide more control and limit any unwanted side effects. In addition, researchers have focused on creating biomimetic scaffolds capa ble of providing appropriate biochemical cues to encapsulated cells. A number of previous works have created injectable materials that possess specific biomimetic properties for neural and retinal tissue engineering. In one such study, synthetic hydrogels were engineered to specifically mimic the biochemical composition and neurotrophic potential of native brain ECM. The constructed hydrogels were able to promote 3D neurite outgrowth in in vivo studies 105 . In another study, researchers designed an injectable and bioresorbable hydrogel polymer composed of hyaluronan and methy cellulose (HAMC). This hydrogel platform was shown to enhance cell survival and integration of transplanted retinal stem cell (RSC) derived rods in the retina 106 . However, limitations of the chemical conjugation process can constrain the amount and type of biomolecules conjugated to the polymer backbone. Researchers have also begun altering the mechanical properties of injectable scaffolds to induce specific cellular behavior of implanted cells. Matrix stiffness varies with respect to the different organ systems in the body. For example retinal tissue show s a very low modulus of around 100 Pa 83 while bone has a modulus closer to 50 kPa 107 . Cells of different lineages naturally reside in tissue types with specific mechanical properties. Proving this point, researchers have shown that stem cells grown on stiff materials have a higher propensity to differentiate into bone cell

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! ! ! ! 28 progenitors 107 while stem cells grown on softer material are more likely to differen tiate into neural cell lineages 108 . The different mechanical properties of each tissue type will alter how they interact with cells. To encourage incorporation and prevent damage, polymer materials used to replace native tissue should possess similar mechanical properties to the tissues they are aimed at replacing. Neural tissue has an extremely low modulus which makes surgical implantation of a material soft enough to replace it relatively impossible 83 . For this reason, injectable scaffold are again beneficial when trying to replace damaged neural tissue specifically. Finally, researchers have also begun to modify the properties of scaf folds to better mimic the morphological structure of native tissue. In native retinal tissue specifically, RGCs grow in a laminar formation with axons extending in a unilateral direction towards the optic nerve cup 109 . However previously developed polymer scaffolds used in retinal tissue engineering simply form a porous structure forcing axons to extend in all directions. To better mimic native retinal tissue, researchers have begun to use electros pun scaffolds 110 111 and retinal sheets to enhance transplantation into the retinal space 112 . One research group was able to show axon growth in a pattern consistent with the structure of native retinal tissue through guidance of embedded RGCs using biochemical and morphological cues 111 . In another study, retinal sheets taken from rats were used to construct a scaffolding system that better mimics the sheet formation of native retinal tissue 112 . However, once again these materials require surgical implantation and likely do not possess a comparable modulus to soft neural tissue. The ideal material would be one that could be injected and take the shape of native retinal tissue while possessing the appropriate mechanical properties. As you can see, much of the focus of retinal tissue engineering has been to construct a material that can mimic native tissue in the clo sest way. In this work, the goal was to engineer a injectable

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! ! ! ! 29 material that can mimic the biochemical, mechanical, and even the morphological properties of native retinal tissue. The first component of this polymer system is a previously described highly f uncationalizable backbone with a protein like structure aimed at mimicking a cell's native extracellular matrix 23 . T he second component of this polymer system is a rginine glycine aspartic acid (RGD), an integrin/cell binding motif foun d in many components of the ECM . It has been well established that cellular adhesions made through RGD integrin binding can pr omote cell survival, cell spreading, proliferation, and neurite extension. RGD was chemically conjugated to the polymer backbone, in order to provide a stable linkage capable of withstanding the cellular contractile forces and promote strong cellular adhes ion 113 . Lastly, we modified this polymer system with poly (N isopropylacrylamide) (PNIPAAm) to allow for an injectable cellular scaffold. PNIPAam, as discussed above is a thermosensitive water soluble homopoly mer that exhibit s a sharp, reversible sol gel transition point sufficiently above room temperature and sufficiently below body temperature ( 32 ¡ C) 114 maki ng it extremely useful for many biomedical applications 88 . When conjugated with other polymers, PNIPAAm may impart its therm osensitive behavior on otherwise non stimuli sensitive polymers or molecules. Following synthesis and characterization of this polymer system (Chapter 2) , we investigated the 3D growth of RGCs cultured within this polymer scaffold (Chapter 3) . The results of this work provide insight into the complex mechanical and chemical environment that support RGC growth in 3D and could have a direct application in cellular replacement therapies for the treatment of glaucoma and other neurodegenerative diseases. Inje ctable systems for neurotrophic factor delivery In this final section we will discuss the use of injectable materials for the delivery of protein drugs (such as NFs) in the setting of retinal tissue engineering. Importantly, the end goal of this work would be a polymer system that may be used for concurrent NF delivery and cell transplantation to maximize cell survival, integration, and function. Initial studies to administer

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! ! ! ! 30 NF consisted of direct injection of the NF to the treatment site. This method pres ents with physiochemical instability, rapid diffusion, and a short half life. Due to this, multiple high concentration doses would be required to r ealize the full effects of the N F, making the use of NFs impractical and potentially dangerous in the clinica l setting 115 . For this reason, an appropriate delivery system that can protect protein drugs from enzymatic release while controlling release is an encouraging approach to improve protein stability and control release within the body. In this section we will focus on synthetic, injectable materials. A number of different polymer system have been investigated for controlled release of protein drugs. Amongst them, hydrogels have become a popular material for protein drug delivery with an injectable material. In one such study, investigators used photoinitiated polymerization to contruct hydrogels based on poly(ethylene glycol). They then used the degree of polymerization to control the neurotrophic diffusion from the hydrogel system. The release of NF (specifically CNTF) was assessed using a neurite outgrowth assay with cells taken from retinal explants. CTNF released from the hydrogel polymer triggered significantly more neurite outgrowth compared to controls 116 . In a similar study, researchers studied PEGPLA photocrosslinkable hydrogels that release BDNF. They were able to show an increase in neurite survival and neurite extension when the NF eluting polymer was cultured with retinal explants. However, these results were absent by day 14 lending to the limitations of the hydrogel systems 117 . There are many more variations of polymer based protein drug delivery systems such as microspheres, nanoparticles, etc; howev er, none of these can simultationously be used as cell transplantation systems. Furthermore, the hydrogel systems that may also be used as scaffolds remain limited by the length of NF delivery and the setbacks discussed above.

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! ! ! ! 31 In this work, we utilized th e same polymer system backbone discussed in the previous section, PSHU, capable of attaching a large quantity of functional groups. In this case, the polymer backbone will be further modified with PNIPAAm as well as negatively charged sulfonate groups simi lar to those seen in Heparin. Heparin, a naturally sulfated biopolymer with an intrinsic negative charge, exists in the ECM where it interacts with the overall positive charge of N Fs 118 . This electrostatic interaction betw een the polysaccharide and the N F protects the N F from proteolytic degradation, preserves its bioactivity, and increases its half life 119 . However the use of heparin itself as a N F delivery system has serious drawbacks includ ing, batch batch variability in structure and biocompatibility as well as unwanted biological interactions with non target tissue 59 120 . Therefore, it is more advantageous to incorporate solely the functional elements of heparin (negative su lfonate groups) into the polymer system itself. The biomimetic polymer backbone and conjugated sulfonate groups are aimed at the improvement of N F stability and long term stable release of the GF. Following characterization of this polymer system (Chapter 2 and Chapter 4), we investigated the release of NF from the polymer system using an optic nerve rat glaucoma model (Chapter 4).

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! ! ! ! 32 CHAPTER II SYNTHESIS AND CHARACTERIZATION OF A BIOMIMETIC REVERSE THERMAL GEL FUNCTIONALIZED WITH INTEGRIN BINDING RGD MOTIFS FOR 3 DIMENSIONAL RETINAL TISSUE ENGINEERING Cell replacement therapies have become a promising area of research for the treatment of many diseases including neurodegenerative diseases such as glaucoma . Although this type of research has shown some promise, there are still many complicatio ns to overcome. One obstacle is the lack of an implantable cell scaffold that can provide transplanted cells with a growth permissive environment following transplantation. We aimed to overcome this hurdle by s ynthesizing a bioactive injectable material that may be injected in conjunction with the replacement cells to provide mechanical and bioactive cellular support. We initially synthesized poly(serinolhexamethylene urea) co pol y(N isopropylacrylamide) (PSHU P NIPAAm) and subsequently conjugated a pentapeptide, Gly Arg Gly Asp Ser (G RGDS) (PSHU NIPAAm GRGDS). In this work, the polymer synthesis was first optimized an d each component characterized. Next we analyzed the gelling properties to show ideal temperatur e dependent thermal gelling properties. Finally, the physical structure of this polymer was analyzed to show the morphology once gelled. The results of this section show that this polymer system has the potential to be manipulated immensely and is highly s uited for retinal tissue engineering. Introduction Substantial research has been conducted to investigate the signals responsible for promoting neural cell survival and axon extension in hopes of incorporating these signals to improve cellular scaffolds. Unfortunately, this task has proved to be extremely challenging due to the complex progression of signals required to induce specific cellular behavior 121 . In the human

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! ! ! ! 33 body , cells reside in cellular niches that control cell behavior through various cell signals 122 123 . These signals are complex and multifac eted including cell cell interactions, cell biomolecule interactions, and ECM interactions. For this reason, great interest has been given to synthetic scaffolds that mimic the ECM environment 124 . Synthetic polymers have easy to control properties such as stiffness or morphology as well as allowing the attachment of functional biomolecules. Thus polymer systems may be t ailored to promote cell adhesion, proliferation, and survival of the encapsulated cells giving potential for use with cellular transplantation. We propose the use of a synthetic polymer tha t can be fine tuned and manipulated to mimic the signals found spec ifically in the retinal ECM. PSHU (Figure 2.1) was employed due to its protein like backbone structure and its potential to attach a large quantity of biomolecules (18 potential linkages per molecule). The protein like backbone structure of the polymer ma y provide a cellular environment more similar to the naturally occurring proteins within the extracellular matrix. The functionalizable aspect of this polymer is extremely beneficial for these purposes because achieving a high concentration of biomolecules for cell biomolecule interactions plays a crucial role in cell survival and axon extension 125 . The RGD sequence, an integrin binding motif found in fibronectin and laminin (major components of the ECM), was found to be implicated in outside inside cell signaling that can affect cell proliferation, migration, cell survival, and axon extension in the case of neurons 126 . Therefore , incorporating enough of this RGD sequence into synthetic polymers to produce a synthetic scaffold has the potential to increase cell attach ment and health 127 . Not only do cell biomolecule interactions play a role in directing stem cell fate, but also cell ECM interactions can help to modulate neural stem cell behavior an d differentiation 128 . To cater to these

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! ! ! ! 34 interactions, we engineered this polymer to possess multiple peptide mimicking bonds to increase biocompatibility of the polymer. Furthermore, we ai med to create a material that is mechanically and morphologically similar to native retinal tissue. The use of a retina like scaffold could induce the same morphology and axon extension seen in the retinal ganglion layer. Finally the polymer system was inc orporated with PNIPAAm to allow delivery through simple injection . Perhaps the ultimate goal of tissue engineering is to deliver scaffolding through a minimally invasive injection 129 . This way, damaged and diseased niches may be replaced with cell loaded injectable synthetic niches without causing additio nal damage through surgical or implantation procedures. PNIPAAm is a thermo responsive polymer with a LCST of around 32 ¡ C. After conjugation of PNIPAAm to the polyurea backbone (PSHU PNIPAAm), the entire polymer system will possess these RTG properties. The first goal of this work was to synthesize and characterize an appropriate injectable scaffold. To confirm and characterize the polymer backbone we employed nuclear magnetic magnetic resonance ( NMR ) and fourier transform infrared spectroscopy ( FT IR ) . I n addition, LCST and gelling studies were used to investigate the temperature and concentration required to form a stable gel. Finally, scanning electron microscopy (SEM) was used to investigate the morphology of the polymer once gelled and rheological st udies were done to determine the storage modulus of the polymer once gelled.

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! ! ! ! 35 Figure 2.1. Schematic synthesis of PSHU PNIPAAm RGD. The first stage is PSHU synthesis followed by the removal of the N B OC groups from the PSHU backbone and finally RGD conjug ation. Materials and Methods Materials N BOC Serinol, urea, hexamethylene diisocyanate (HDI), anhydrous chloroform, and anhydrous N,N dimethylformamide (DMF) were purchased from Sigma Aldrich (St. Louis, MO, USA). N (3 Dimethylamino propyl) N ! ethylcarbodiimide hydrochloride (EDC), N hydroxysuccinimide (NHS), 2,2,2 trifluoroethanol (TFE), and trifluoroacetic acid (TFA) were purchased from Alfa H N O O O O H N NH 2 N H N H N H O O N H O n H N O O O O H N NH N H N H N H O O N H O 1-n O C N N C O H 2 N O NH 2 HO HO NH O O Hexamethylene Diisocyanate Urea N-Boc Serinol + + H N O O O O H N NH 2 N H N H N H O O N H O n H N O O O O H N NH N H N H N H O O N H O 1-n Step 1 O O H N NH O O N m Step 2 i. PSHU Synthesis ii. PSHU Deprotection iii. Conjugation of PNIPAAm-COOH H N O O O O H N NH N H N H N H O O N H O n H N O O O O H N NH N H N H N H O O N H O 1-n H N NH O O N m Step 3 iv. Conjugation of GRGDS-COOH O GRGDS

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! ! ! ! 36 Aesar (Ward Hill, MA, USA). Anhydrous diethyl ether was purchased from Fisher Scientific (Pittsburgh, PA, USA). Anhydrous dichloromethane (DCM) was purchased from JT Baker (Phillipsburg, NJ, USA). The pentapeptide Gly Ð Arg Ð Gly Ð Asp Ð Ser (GRGDS) was purchased from Biomatik (Wil mington, DE). Dialysis tubing (Spectra/Por) was obtained from Spectrum Labo ratori es (Houston, TX). Equipment. Equipment Gel permeation chromatography (GPC) was recorded on a Viscotek GPC Max with 270 Dual Detector with right angle light scattering (RALS) and VE 3580 refractive index (RI) detector from Malvern Instruments (Houston, TX USA). Each sample's number average molecular mass (M n ) and weight average molecular mass (M w ) was calculated using OmniSEC 5.02 software. 1 H NMR spectra were recorded on a Varian 500 NMR (500 MHz, Varian) with CDCl 3 as the solvent at 25¡C. Chemical shifts are in ppm using the solvent peak as the internal reference. Fourier transform infrared (FT IR) spectra were recorded on a Nicolet 6700 (Thermo Fisher Scientific, Waltham, MA) using polyethylene windowed equipment. High performance . LCST measurements were made with a Cary UV Vis s pectrophotometer using quartz cuvettes with 1 wt% solutions of polymer. Polymer morphology was imaged using a JEOL (Peabody, MA) JSAM 6010la analytical scanning electron microscope. Rheological measurements w ere performed on a stress controlled rheometer (Rheostress Haake RS 150) using a cone and plate geometry (angle of 1¡, diameter of 60 mm) and a solvent trap to prevent evaporation of the polymer solution. Synthesis of poly (serinol hexamethylene urea) (PSHU) PSHU was synthesized by combining HDI (1.928ml, 12mmol), N BOC Serinol (1.147g, 6mmol), and urea (0.360g, 6mmol) in 6ml of anhydrous DMF. This reaction was maintained at 90 o C for 7

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! ! ! ! 37 days. After the 7 day reaction period, the solution was coole d to room temperature and rotary evaporated at 70 o C to remove the DMF. Next, the product mixture was re dissolved in a small amount of DMF (2 3ml) and precipitated into cooled anhydrous diethyl ether (100ml). The purification process was repeated twice and the solution was washed overnight in excess ether (100ml) to remove any unreacted reactants and remaining solvent. Finally, the product was rotary evaporated at 45 o C until dry; the final product was stored at room temperature until conjugation with RGD. De protection of PSHU In order to expose free amine groups on the PSHU backbone and subsequently conjugate PNIPAAm and RGD, the tert Butyloxycarbonyl (BOC) groups must first be removed. To do this PSHU (1.0g) was dissolved in equal parts of methylene chlor ide (15ml) and trifluoroacetic acid (TFA) (15ml) in a round bottom flask. This reaction was held at room temperature with gently stirring for 30 min. Rotary evaporation was used to remove the solvent and the product m ixture was dissolved in DMF (1ml ). The solution was precipitated in cold diethyl ether and rotary evaporation was used to remove the residual ether. This precipitation and evaporation method was carried out once more and the deportected PSHU (dPSHU) was dissolved in 1ml of TFE and precipitated again in cold diethyl ether. After final rotary evaporation, the purified polymer was stored at room temperature. Synthesis of poly (N isopropylacrylamide) (PNIPAAm) PNIPAAm COOH was prepared as previously described 130 . In brief, NIPAAm (44.19mmol) and ACA (0.22mmol) were dissolved in 25ml anhydrous methanol. Next, this solution was bubbled with nitrogen for 30 min and mixed for 3h at 68 ¡ C. This product solution was precipitated in hot water (60 ¡ C) to purify the PNIPAAm COOH. The product was washed twice in hot water (60 ¡ C)

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! ! ! ! 38 and further purified by dialysis (MWCO: 3500Da) for 2 days. This final product was freeze dried and stored at room temp erature (yield: 92%, Mw 11,000) Conjugation of PNIPAAm to PSHU PNIPAAM was conjugated to 25% of the free amine groups on the PSHU backbone using EDC/NHS chemistry. To begin, PNIPAAm COOH (0.04mmol) was dissolved in 2 ml anhydrous DMF. EDC (0.048mmol) and N HS (0.048mmol) were dissolved in 1 ml anhydrous DMF and added drop wise to the PNIPAAM COOH dissolved in DMF. This activation reaction was allowed to proceed for 24hrs. Next, deprotected PSHU (0.1956mmol amine groups) dissolved in 1 ml DMF was added to the activated PNIPAAM and allowed to conjugate for 24 hrs. The resulting polymer was precipitated three times in cool diethyl ether and the solvent was removed each time with rotary evaporation. The dried polymer was dissolve in 5ml miliQ water at 4¡C, dialyz ed (MWCO: 12,000Da) for three days, and finally lyophilized). Conjugation of RGD to PSHU PNIPAAm copolymer A similar conjugation protocol was used to conjugate RGD to PSHU as well as conjugating RGD to PSHU PNIPAAm. RGD (0.125g, 0.25mmol), EDC (0.143g, 0. 75mmol), and NHS (0.86g, 0.75mmol) were dissolved in 1 ml ultra pure water for conjugation to PSHU PNIPAAm. For conjugation to PSHU, 1ml of anhydrous DMF was used instead of ultra pure water. This solution was reacted at room temperature for 2 h. PSHU PNI PAAm (0.10g) or PSHU (0.10g) was dissolved in ultra pure water or DMF respectively and added drop wise to the appropriate activated RGD solution and reacted for 24 h at room temperature, protected from light. The PSHU PNIPAAM RGD solution mixture was plac ed in dialysis tubing (MWCO: 3500) and dialyzed against 1 liter of ultra pure water for 24 h, with one water change. After dialysis, the solution was freeze dried, resulting in a white, flaky precipitate, which was stored at room

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! ! ! ! 39 temperature. However, the PSHU RGD solution mixture precipitated three times in excess ether, dried using rotary evaporator, and stored at room temperature. Gel permeation chromatography (GPC) (Malvern Instruments, Houston, TX USA) was used to determine the molecular weight distri bution of the synthesized copolymers. Analysis was performed on each sample using a 100 " l injection into a single Viscotek D6000M column and 270 Dual Detector with right angle light scattering with DMF as the system solvent. The column and detector temperatures were kept constant at 45¡C. The instrument was calibrated with polystyrene standards (MW: 105,000, dn/dc: 0.185 ml /g). ! Solution to gel phase transition using LCST To examine the thermal gelling temperature, the lower critical solution temperature (LCST) of PSHU PNIPAAm and PSHU PNIPAAm RGD was measured by loading a 1% (wt/v) solution in deionized water into a UV/visib le spectrophotometer fitted with a temperature controlled cell and reading percent transmittance at 480 nm at temperatures between 20¡C and 45¡C. Solution to gel phase transition using gelling tests The sol gel phase transition of the PSHU PNIPAAm RGD sol utions was determined by a test tube inversion method. The polymer was dissolved in PBS with various concentrations and each polymer solution (2ml) was placed in a glass vial. The initial temperature of a water bath was set to 25¡C and heated up to 50¡C wi th 1¡C/min. The sol gel phase transition was determined by inverting the vial horizontally at each temperature. Scaffold morphology using SEM Polymer solutions of 2.5, 5, and 10% were prepared in ultra pure water and allowed to gel at 37 ¡ C for 15 min. The gelled samples were then frozen quickly using liquid nitrogen, cut in half to

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! ! ! ! 40 expose the center structure, and rapidly transferred to a freeze dryer for 24 h ( 48 ¡ C , 38*10 3 ). Liquid nitrogen was used to freeze the samples rapidly (~3 sec ) to avoid de gelling of the polymer and preserve its 3D structure. The samples were then sputter coated with gold for 30 s and the cross section of the gel was analyzed using SEM. Mechanical properties using rheology First, polymers (PSHU PNIPAAm RGD and PSHU PNIPAAm) were dissolved in RGC media at concentrations of 2.5, 5 and 10 wt%. Temperature sweep tests composed of heating ramps (at 5 ¡C/min) were conducted at constant frequency (1 Hz) and stress (0.05 Pa) between 25 ¡C and 45 ¡C. Statistical analys is Statistical significance between three or more data sets will be determined by ANOVA, while the t test will be used to compare significance between 2 groups. A p value of < 0.05 will be considered statistically significant. Results and Discussion !"#$%&'( Characterization ( The molecular weight distribution of PSHU was determined by GPC. GPC analysis showed molecular weights for PSHU, number average molecular mass (M n ): 1,610 Da, weight average molecular mass (M w ): 3,354 Da, and PI (M w /M n ): 2.083 (See appendix A ). NMR and FT IR were both used to confirm the overall polymer structure and to detect any remaining BOC protecting groups. The NMR spectrum of PSHU confirmed the expected copolymer structure, with peaks at 1.3 ( Ð CH2 Ð ), 1.5 ( Ð NH Ð CH2 Ð CH2 Ð ), a nd 3.2 ( Ð NH Ð CH2 Ð ) associated with HDI, at 1.4 ( Ð C Ð (CH3)3), and 4.1 ( Ð CH Ð NH Ð ) associated with N BOC serinol (Figure 2.2). NMR

PAGE 50

! ! ! ! 41 analysis was also able to show complete deprotection of the BOC protecting groups from the PSHU backbone as well as the expected po lymer structure. This can be seen by the absence of a peak at 4.1 (corresponding to N BOC serinol) in the NMR spectrum for deprotected PSHU (Figure 2.3). Figure 2. 2 . 1 H NMR (500 MHz, CDCl 3 ) spectrum of PSHU to confirm overall structure of polymer chain. ! ! ! ! ! ! ! ! ! ! ! ! !

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! ! ! ! 42 Figure 2. 3 . 1 H NMR (500 MHz, CDCl 3 ) spectrum of PSHU and deprotected PSHU to confirm removal of the BOC protecting groups during deprotection process. FT IR spectroscopy was used to confirm the conjugation of RGD to the polymer backbone by viewing the region of 1630# 1680 cm# 1 (Figure 2.4). This region is associated with the carbonyl groups found within the polymer backbone as well as the carbonyl groups found within peptide bonds of RGD. The wavenumbers correlated to the carbonyl groups o f RGD are slightly lower than those of the carbonyl groups in the polymer backbone. In the PSHU# PNIPAAm# RGD spectrum, we can observe an obvious shift in this carbonyl peak toward the lower end of the spectrum, indicating the presence of carbonyl groups in the RG D peptide.

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! ! ! ! 43 Figure 2. 4 . FT IR of dPSHU, PSHU, and PSHU RGD. Confirmation of free amine groups on dPSHU after deprotection (B). Shift in carbonyl absorbance to confirm attachment of RGD to the polymer backbone (A). High performance liquid chromatography (HPLC) was also used to quantify the amount of GRGDS COOH that was successfully conjugated to the free amine groups on the polymer backbone. A calibration curve was first constructed using known concentrations of GRGDS COOH (0.78# 200 µg/mL) ( See appendix B ). Each of these concentrations produced a corresponding HPLC peak area. Using the calculated area beneath the peaks and the corresponding calibration curve, we were able to determine the amount of RGD within each sample and thus the corresp onding RGD conjugation efficiency. Results showed 93% conjugation of RGD to the free amine gro ups on the polymer backbone.

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! ! ! ! 44 Analysis of gelling properties LCST was used to analyze the gelling properties of both PSHU PNIPAAm and PSHU PNIPAAM RGD. This data gave us valuable information on not only the temperature of sol gel transition, but also the relative rate of gelation. As shown in Figure 2.5 , the transmittance of both aqueous solutions of PSHU PNIPAAm and PSHU PNIPAAM RG D decreased slowly upon heating from 20¡C to 31¡C, reached almost zero at 32¡C, and turned to an opaque solid upon further heating over 33¡C, indicating that the aqueous solution turns to a physical gel as the temperature increases. PSHU PNIPAAm RGD exhibi ted a LCST and phase transition profile very similar to PSHU PNIPAAm, remaining in solution state at temperatures below 32 ¡ C and rapidly undergoing a phase transition to a physical gel upon reaching body temperature. These unique characteristics will allo w PSHU PNIPAAm RDD to be administered through a minimally invasive injection at the desired location. Figure 2. 5 . Temperature dependent phase transition of PSHU PNIPAAm (orange) and PSHU PNIPAAm RGD (blue). Both polymers display similar gelling temperature while the incorporation of the RGD peptide slightly increased the gelling time. !

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! ! ! ! 45 S ince the PSHU PNIPAAm RGD was designed to show temperature dependent solution to gel phase transition, its thermal gelli ng properties were examined. Although not significant, the gell ing temperature of PSHU PNIPAAm RGD decreased from 32 ¡ C to 31 ¡ C as the co ncentration of aqueous solution increa sed from 3 to 24% (wt) (Figure 2.6 ). More importantly all aqueous solutions remained a gel at 37 ¡ C indicating that PSHU PNIPAAm RGD is a promising temperature dependent injectable material. Moreover, the gel status was maintained up t o the highest temperature (50 ¡ C) with no phase separation. Figure 2. 6 . Thermal gellin g property of PSHU NIPAAm GRGDS. All solutions turned to physical gel upon temperature increase and maintained gel status in a broad range (blue area) of temperature . Morphology and mechanical properties SEM was used to investigate the morphology of the 3D polymer scaffold after gelation at body t emperature. As shown in Figure 2.7 , higher concentrations (5 and 10 wt /v %) of both PSHU# PNIPAAm# RGD and PSHU# PNIPAAm assembled into a laminar sheet like

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! ! ! ! 46 conformation upon gelling. However, at a low concentration (2.5 wt /v %), both polymer systems formed a heterogeneous structure, implying unstable gelling conditions. The sheet like conformation adopted by this polymer system is very similar to that of retinal tissue. Retinal tissue has a laminar organization in w hich different cell types are layered within specific retinal strata. To provide a more suitable 3D culture and induce greater axon extension, we engineered a polymer system with properties similar to t hat of retinal tissue structure. Retinal tissue is lam inar in structure with RGC axons extending horizontally toward the optic cup. Therefore, recreating this sheet structure could provide topographical cues and, along with biomolecular cues of RGD, guide horizontal extension 131 4 . Figure 2.7 . Scaffold morphology. Representative SEM images of PSHU PNIPAAm RGD (bottom row) and PSHU PNIPAAm (top row) at various concentrations. Concentrations above 2.5 wt /v % show a laminar sheet like formation of the gelled polymer. Scale bar is 50 µm.

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! ! ! ! 47 Rheology was used to assess the mechanical properties of the polymer scaffold as compared to native retinal tissue. Figure 2.8 shows the storage modulus (G!) plotted against increasing temperature. All six polymer samples showed a G! that was greater than the loss modulus (G$)(See a ppendix A ), indicating the dominant elastic beh avior of the polymer scaffold 131 . The value for PSHU# PNIPAAm# RGD and PSHU# PNIPAAm both increased with increasing polymer concentration and temperature. In addition, the incorporation of RGD showed a slight increase in G! at all three polymer concentrations. In addition, we were able to create a polymer syst em with a G! similar to that of retinal tissue. At a concentration of 5 wt /v %, the polymer system possessed the laminar sheet like morphology seen in the native tissue 132 a nd had a G! of 200 Pa , which is comparable to the valu e of 100 Pa of retinal tissue, 83 making this system the best choice for subsequent studies.

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! ! ! ! 48 Figure 2.8 . Rheological properties of polymer scaffold. The storage modulus was plotted vs. temperature for various concentrations of each polymer system (PSHU PNIPAAm RGD and PSHU PNIPAAm). Increasing either polymer concentration or temperature leads to an increase in the storage modulus of the material. Conclusion In this work, we have demonstrated that potential PSHU NIPAAm RGD has for retinal tissue engineering. The polymer backbone, PSHU, has the capacity to be modified with various moieties making it easy to manipulate for various tissue engineering applications. In this chapter we used RGD as a biomolecular conjugate; ho wever, this polymer may be conjugated with a variety of molecul es and chemical groups depending on the purpose. Furthermore, once conjugated to PNIPAAm, the PSHU polymer backbone possessed appropriate reverse thermal gel properties for injection and gelati on at body temperature. In addition, the reverse thermal gelling property of this polymer system will allow facile encapsulation of cells in a three dimensional (3D) scaffold. Finally, this polymer system was created to possess both topographical and mecha nical cues that mimic the native retinal microenvironment. In the

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! ! ! ! 49 subsequent section, we will access the potential of this scaffold for R GC survival and axon extension.

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! ! ! ! 50 Chapter III A SELF ASSEMBLING INJECTABLE BIOMIMETIC MICROENVIRONMENT ENCOURAGES RETINAL GANGLION CELL AXON EXTENSION IN VITRO Sensory somatic nervous system neurons, such as retinal ganglion cells (RGCs), are typically thought to be unable to regenerate. However, it is now known that these cells may be sti mulated to regenerate by providing them with a growth permissive environment. We have engineered an injectable microenvironment designed to provide growth stimulating cues for RGC culture. Upon gelation, this injectable material not only self assembles int o laminar sheets, similar to retinal organization, but it also possesses a comparable storage modulus to retinal tissue. Primary rat RGCs were grown, stained, and imaged in this 3D scaffold. We were able to show that RGCs grown in this retina like structur e exhibited characteristic long, prominent axons. In addition, RGCs showed a consistent increase in average axon length and neurite bearing ratio over the 7 day culture period indicating this scaffold is capable of supporting substantial RGC axon extension . Introduction Glaucoma, amongst other optic neuropathies, leads to the neurodegeneration of retinal ganglion cells (RGCs), the projection neurons located in the retina with axons extending through the optic nerve 1 . These cells play a crucial role in sight by transmitting visual information from the bipolar, amacrine, and interplexiform cells of the retina to the visual cortex of the brain 133 . Due to the inability of these cells to regenerate in the normal human disease condition, the loss of these cells is permanent 134 . Additionally, clinical therapies for glaucoma are currently limited to treatments that p revent or limit further damage to the RGCs 135 , 5 . Future interventions that seek to regain or improve visual function must not only inclu de mechanisms for RGC neuroprotection

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! ! ! ! 51 but also methods to facilitate the survival and axon regeneration of damaged RGCs and eventually methods to replace dead RGCs. It was previously believed that RGCs, like many central nervous system neurons, do not pos sess the ability to regenerate following injury or death. However, it is now known that the limited regeneration of axonal regrowth of these cells is possible but inhibited due to the injured microenvironment (myelin associate molecules) 136 , 137 , 138 , scar formation 139 , 140 , and lack of passage across a lesion 141 , 142 , 143 . Therefore, the regenerative capacity of RGCs may be stimulated by creating an alternate extracellular microenvironment that will instead activate RGC growth, maintain RG C viability, and counteract the inhibitory signals of the injured nerve 72 . To alter the fate of damaged RGCs, the cells must be encapsulated i n a growth permissive microenvironment, protected from the diseased environment, presented with cell binding molecules, and exposed to appropriate mechanical properties to induce and cue growth. Here, we have developed an injectable biomimetic three dimen sional (3D) scaffold with similar mechanical and morphological properties to native retinal tissue. The first component of this polymer system is a previously described, with highly functionalizable backbone aimed at mimicking a native ECM 23 . The second component of this polymer system is the small peptide RGD , an integrin/cell binding motif found in many components of the ECM. It has been well established that cellular adhesions made through RGD integrin binding can promote cell s urvival, cell spreading, proliferation, and neurite extension 23 . In this study the peptide sequence GRGDS was used instead of RGD in order to preserve the integrity of the entire RGD binding motif. Studies have shown that the RGD tripeptide has little effect on cell attachment; however surrounding the RGD motif with flanking amino acids according to the natural sequence (GRGDS) can preserve activity of this integrin binding motif 144 . Lastly, we modified this

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! ! ! ! 52 polymer system with PNIPAA m to allow for an injectable cellular scaffold. PNIPAAm, a thermosensitive water soluble homopolymer, has garnered a lot of attention in the biomedical field 145 114 146 147 . This polymer ha s been shown to exhibit a sharp, reversible sol gel transition point at 32 ¡ C , sufficiently above room temperature and sufficiently below body temperature making it extremely useful for many biomed ical applications 88 . Following synthesis and characterization of this polymer system (as described in the Chapter II) , we investigated the 3D growth of RGCs cultured within this polymer scaffold. The results of this work provide insight into the complex mechanical and chemical environment that support RGC growth in 3D and could have a direct application in cellular repla cement therapies for the treatment of RGC associated ocular neurodegeneration and other neurodegenerative diseases. Material and Methods Materials Dimethyl sulfoxide (DMSO) and bovine serum albumin (BSA ) were purchased from Sigma Aldrich (St. Louis, MO, USA). Succinic acid, tetrahydrofuran (THF), anhydrous methanol, 4 Dimethylaminopyridine (DMAP), N,N' Dicyclohexylcarbodiimide (DCC), N (3 Dimethylamino propyl) N' ethylcarbodiimide hydrochloride (EDC), ch loroform d, and N hydroxysuccinimide (NHS) were purchased from Alfa Aesar (Ward Hill, MA, USA). Anhydrous N, N Dimethylformamide (DMF) was purchased from EMD Millipore (Billerica, MA, USA). N Isopropylacrylamide (NIPAAm) was purchased from TCI Chemicals (Portland, OR, USA). Dichloromethane (DCM), Acetone, and 30% hydrogen peroxide were purchased from BDH Chemicals through VWR (Radnor, PA, USA). Anhydrous diethyl ether was purchased from Fisher Scientific (Pittsburgh, PA, USA). Gly Ð Arg Ð Gly Ð Asp Ð Ser (RGD) was purchased from Biomatik (Wilmington, DE, USA) and Cellmano Biotech Limited (Hefei, AnHui,

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! ! ! ! 53 China). Sprague Dawley rats were purchased from Charles River Laboratories (Wilmington, MA, USA). Optimal cutting temperature (OCT) compound was purchased from S akura (Torrance, CA, USA). MTT Cell Proliferation Assay Kit was purchased from Invitrogen Molecular Probes (Carlsbad, CA, USA). Goat serum, GFAP (mouse IgG1), GAP 43 (Rabbit IgG), Alexa Fluor 488 (goat anti mouse IgG), Alexa Fluor 594 (goat anti rabbit Ig G), and SlowFade Diamond antifade mountant with DAPI were purchased from Life Technologies (Carlsbad, CA, USA). %III tubulin (goat anti rabbit, IgG) was purchased from abcam (Cambridge, MA, USA). Fluoromount G with DAPI was purchased from Electron Micros copy Sciences (Hatfield, PA, USA). Triton X 100 was purchased from MP Biomedicals. Retinal Ganglion Cell Isolation Kits (rat) were obtained from MACS Miltenyl Biotec (San Diego, CA, USA). Equipment Confocal images were collected using a Nikon Eclipse Ti C 2 LUN A microscope (Nikon, Tokyo) equipped with two C2 DU3 high sensitivity PMT detectors, 4 diode lasers (405/488/561/640 nm), and a motorized microscope stage with 3 axis navigation (X, Y, and Z) controlled by the NIS Elements software package. Laser and software setting were kept constant between specimens and to allow for comparison of different image acquisitions. Glass bottom culture dishes for RGC growth were purchased from MatTek Corporation (Ashland, MA, USA). Tissue was sectioned using a CryoStar NX70 Cryostat. Confocal images were taken 3i Marianas Spinning Disk microscope. Synthesis of PSHU PNIPAAm RGD See previous chapter for details on polymer synthesis and characterization

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! ! ! ! 54 Animal Procedures All animal experiments were performed in accordance with procedures approved by the Institutional Animal Care and Use Committee (IACUC) at the University of Colorado Denver Anschutz Medical Campus. All experiments were performed in accordance with IACUC gu idelines and regulations. A total of two breeding pairs (Wistar rats male: 350 400 g, Wistar rats female: 250 300 g) were required to complete this project. These animals were used to obtain rat pups by housing one male and one female rat per cage. Once th e rat pups reach postnatal day 5 7, the pups were euthanized by prolonged exposure to CO 2 (~50 min) followed by a secondary form of euthanasia (decapitation). Retinal ganglion cell isolation RGCs were purified from rat pups (postnatal day 5 7) as these cel ls show higher survival rates following the separation process. To begin, eyes were carefully enucleated and transferred to a petri dish filled with D PBS. Using a dissecting microscope, a small incision was made along the anterior part of the eye (behind the lens and cornea). Tweezers were inserted into the small incision and the eye was carefully pulled along this incision line to maintain the integrity of the retina. At this point, the retina was peeled away from the sclera and was moved with a transfer pipette to a 15 ml tube. The dissected retinas were then dissociated using the Neural Tissue Dissociation Kit for Postnatal Neurons from Miltenyi Biotec per the manufacturer's instructions. Following this, the RGCs were purified using the Retinal Ganglion Cell Isolation Kit from Miltenyi Biotec per the manufacturer's instructions. The isolated cells containing mostly RGCs were then resuspended in prewarmed RGC growth medium containing MACS NeuroMedium (130 093 570), NeuroBrew 21 (1:50 dilution, 130 093 56 6 ), sodium pyruvate (Sigma, 1 mm ),

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! ! ! ! 55 BDNF (Peprotech, 25 ng/ml), CNTF (Peprotech, 10 ng/ml), N acetylcystein (50 µ g/ml), insulin (Sigma, 5 µ g/ml), Forskolin (Sigma, 10 µ M), glutamine (2 mM), triiodothyronin (40 ng/ml), streptomycin sulfate (100 µ g/ml), penici llin (100 U/ml). Cells were then plated on PDL Laminin coated coverslips (Data not shown) or seeded in 3D polymer scaffolds. Three separate RGC isolations were completed on three different rat pup litters. 3D culture of Retinal ganglion cells RGCs (8x10 3 ) were suspended in a solution of 5 wt /v % PSHU PNIPAAm RGD or 5 wt /v % PSHU PNIPAAm in complete media. 50 µl of the cell suspension in the polymer solutions were pipetted into each 35 mm glass bottom dish (Mat Tek, Ashland, MA, USA) (Figure 3.1 ) and placed i n a 37 ¡C incubator for 10 min to allow polymer gelation and RGC encapsulation. After incubation, 1 ml of warm, RGC media was added to the culture dish, using a hotplate set at 37 ¡C to maintain gel stability when removed from the incubator. Cells were cu ltured for 3, 5, and 7 days, with media changes each day. Figure 3. 1. Schematic showing 3D culture process. RGCs were first mixed in with the polymer solution and loaded into a glass bottom petri dish. After gelling at 37¡C, RGCs became encapsulated in the polymer scaffold and culture for future analysis.

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! ! ! ! 56 Immunostaining within the 3D scaffold Cells were immunostained while remaining encapsulated in the 3D polymer scaffold. It is important to note that all steps of the staining process were conducted at 37 ¡C to prevent de gelling and disruption of the polymer sample. First samples were washed twice with PBS 1x and fix ed with 4% PFA in PBS for 15 min at 37 ¡C. Next, samples were permeabilized with 1% Triton X (in PBS) for 90 min, followed by a PBS wash overnight. Blocking buffer composed of 2% bovine serum albumin (BSA) in PBS was added to the cells for 90 min. After blocking, cells were incubated overnight with the first primary antibody Brn3a (1:200, prepared in blocking buffer). Cells were then washed with 1% Triton X, 3x for 3 min each. The secondary antibody, anti goat Alexa 594 (1:500) was added to each sample and incubated for 45 min. Cells were washed with PBS Tween (0.002% in PBS) for 3 min and washed twice with PBS, 3 min each. Cells were then incubated with the second primary antibody %III tubulin (1:100, prepared in blocking buffer) overnight at 37 ¡C. C ells were then washed with 1% Triton X, 3x for 3 min each. The secondary antibody, anti rabbit Alexa 488 (1:500) was added to each sample and incubated for 45 min. Cells were washed with PBS Tween (0.002% in PBS) for 3 min and washed twice with PBS, 3 mi n each. Hoechst 33342 (1:2000, in PBS), a DAPI stain, was added to each sample and incubated for 5 10 min, followed by 3 washes in PBS, 3 min each. RGCs were imaged within the polymer scaffold using fluorescence confocal microscopy. Analysis of average n eurite length, average branchpoint, and neurite bearing cells Using a 20& objective, z stack projections of 4 µm thickness were sampled from three random visual fields in each sample. At least nine samples taken from three different RGC isolations were an alyzed for each time point and for each polymer (PSHU# PNIPAAm# RGD and

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! ! ! ! 57 PSHU# PNIPAAm). The simple neurite tracer plugin from FIJI (Longair, Baker, and Armstrong, 2011) was used to analyze the length of each process. Each neurite was traced starting from the cell body extending out into the image frame. The total length of the neuritis was divided by the cell count determined using DAPI to obtain the average neurite length. Statistical analysis Statistical significance between three or more data sets will be d etermined by ANOVA, while the t test will be used to compare significance between 2 groups. A p value of < 0.05 will be considered statistically significant. Results and Discussion 3D Culture of RGCs in an Injectable Polymer Scaffol d The reverse thermal gelling property of this polymer system will allow simple mixing of the RGCs with the polymer at room temperature and encapsulation of the cells within the 3D polymer scaffold upon reaching physiological temperatures. Cells can then be stained and imaged within the scaffold to approximately the center of the gel (250 µ m) (Figure 3.1 ). Unlike two dimensional (2D) systems, this injectable 3D scaffold is more attune with the cell cell interactions, cellular organization, and microenvironmen t seen in native tissue. In 3D , cells can be fully encapsulated in a solid microenvironment rather than solely exposed to one flat 2D surface 129 . Along with being injectable, this polymer system is functionalized with cell binding peptides to improve cellular localization and axon extension. For RGCs specifically, the L1 integrin binding molecule that contains the peptide sequence RGD was shown to be present i n axonal regeneration of RGCs 148 . Because integrins modulate how cells interact with their substrate, increasing the number of L1 binding sites (RGD) within the polymer scaffold could increase the

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! ! ! ! 58 level of cell adhesion, cell growth, and extension of axons within the 3D matrix. RGD was chemically conjugated to the p olymer backbone, to provide a stable linkage capable of withstanding the cellular contractile forces and prom oting strong cellular adhesion 113 . The aim of this work was to determine how RGCs would behave once encapsulated in the 3D polymer scaffold. Following optimization and characterization of the injectable material, we were able to select the polymer system that was most similar to native retinal tissue and would likely support the growth and axon extension of RGCs. RGCs from rat pups postnatal day 5 7 were purified, and seeded simply through mixing with the polymer solution. Once the cell/polymer mixture was placed in the incubator, the cells became encapsulated in the 3D scaffold. Using confocal microscopy , we were able to visualize RGC axons growing in 3D t hroughout the scaffold. Figure 3.2 displays maximum intensity images from 200 µm thick z stacks with 4 µm intervals. Cells grown within P SHU PNIPAAm RGD (5 wt /v %) showed robust axon extension whereas cel ls grown in PSHU PNIPAAm (5 wt /v %) showed m inimal axon extension (See Appendix D ). Furthermore, RGCs grown on the PDL Laminin coated cover slips displayed a Ôstar like' morphology extending shorter axons in all directions (Figure 3.3 ). In contrast RGCs in their native environment in the retina typically e xtend fewer and longer axons in a unilateral direction. Interestingly, the RGCs grown within the polymer scaffold presented a morphology similar to RGCs in vivo 109 and less like RGCs grown in 2D culture dishes.

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! ! ! ! 59 Figure 3.2 . Maximum intensity projections of representative 3D fluorescent images. 3D images of RGCs (red) cultures inside of 5 wt /v % PSHU PNIPAAm RGD were taken using a confocal micr oscope after 3, 5, and 7 days in culture. RGCs (red) show long and mostly planar axon extension (green) for the 7 day culture period.

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! ! ! ! 60 Figure 3.3 . Live/dead (green/red) staining of RGC 3 days of culture. RGCs were cultured on PDL Laminin glass coverslips and stained with live/dead. RGCs grown on this coverslips display a s tar shaped morphology typical of RGCs grown in vitro. Image analysis of 3D RGC cultures Average neurite length was determined by measuring the total length of all axons (in 3D) within a visual field and dividing this number by the total cell count within that fie ld. The panels A C in Figure 3.4 show the image an alysis process, where Figure 3.4 A is the maximum intensity image of a visual field and Figure 3.4 B shows the measured axon length in the X Y, X Z, and Y Z frames. Figure 3.4 C graphs the average axon length vs. culture period. We can see that there is a significant difference between the mean axon length at day 3 and day 5 (p value < 0.001) as well as a significant difference between day 5 and day 7 (p value < 0.01). Ratio of neurite bearing cells was calculated by determining the number of neurite bearing cells per frame (only RGCs beari ng neurites with a length greater than two cell bodies were counted) and dividing this number by the total cell count (determined by nuclear DAPI stain). Figure 3.4 D shows the total

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! ! ! ! 61 cell coun t in a visual field and Figure 3.4 E is an example of a maximum in tensity image used for neurite bear ing cell counting. From Figure 3.4 F we can see that there is no significant different between the ratio of neurite bearing cell from day 3 to day 5 or from day 5 to day 7; however, there is a significant difference betwee n the ratio of neurite bearing cells from day 3 to day 7 (p value < 0.01).

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! ! ! ! 62 Figure 3. 4 . Average neurite length and ratio of neurite bearing cells. Quantification of average axon length was performed by analyzing maximal intensity images (A) in all three planes (B) for both PSHU# PNIPAAm# RGD and PSHU# PNIPAAm cultures at 3, 5, and 7 day time points (C). The ratio of neurite bearing cells (F) was quantified by calculating the number of cells bearing neurites in a visual frame (E) and dividing by the total number of cells within that frame (D). Numbers are plotted for both PSHU# PNIPAAm# RGD and PSHU# PNIPAAm over the 7 day culture period. Statistical significance between groups was determined by ANOVA. *p < 0.05; **p < 0.005. Three images were taken fro m each of at least three different samples from three separate RGC isolations. The scale bar is 200 µms. Day 5 cultures are shown to provide examples for quantification (A# E). Conclusion This work aimed to test the efficacy of this 3D scaffold through in vitro studies with RGCs. By studying the survival, axon extension, and morphology of the cells grown in this 3D scaffold, we can determine whether this polymer system is a good model for growing RGCs in vitro as well as a promising scaffold for use in cell replacement therapies. In their native environment, RGCs develop in a single monolayer with long axons extending horizontally toward a single point in the back of the eye, eventually joining together to form the optic nerve. However, current in vitro

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! ! ! ! 63 cult ure conditions (2D) have been able to produce only RGCs with short axons or axons extending in all directions. Although certain groups have been able to produce RGCs showing a unilateral morphology using 3D scaffolds, these studies have used an electrospun scaffold that requires implantation and were unable to seed cells directly within the scaffold 110 . To provide a more suitable 3D culture and induce greater axon extension, we engineered a polymer system with propertie s similar to tha t of retinal tissue structure 131 4 Retinal tissue is laminar in st ructure with RGC axons extending horizontally toward the optic cup. Therefore, recreating this sheet structure could provide topographical cues and, along with biomolecular cues of RGD, guide horizontal extension of RGC axons. Collectively, we created a m aterial that is similar to an RGC's native environment, which could induce the same morphology and axon extension seen in the retinal ganglion layer. Through the use of this retina like scaffold, we were able to mimic a cell's native environment and provid e improved RGC culture conditions in vitro. Culture of RGCs in this 3D polymer system exhibited long, prominent axons, similar to wha t is seen in the native retina (Figure 3.2 ). In addition, the average axon length and ratio of neurite bearing cells increa sed steadily over th e 7 day culture period (Figure 3.4 ). In conclusion, w e developed a 3D scaffold, providing both topographical and biochemical cues of a native retinal microenvironment, with great potential for RGC survival and extensive axon extension. We believe that this 3D polymer system could provide researchers with an improved model for in vitro RGC studies. This type of culture system could be utilized to investigate stratified structures, branching, and synaptogenesis in three dimensions while st ill preserving the ap propriate cellular morphology 149 . These results also show that this injectable scaffold could be promising for use in future cellular replacement studies aimed at treating optic neurodegenerative

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! ! ! ! 64 diseases.

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! ! ! ! 65 Chapter IV AN INJECTABLE NEUROTROPHIC FACTOR DELIVERY SYSTEM SUPPORTS THE SURVIVAL AND REGENERATION FOLLOWING OPTIC NERVE CRUSH INJURY In general, neurons belonging to the central nervou s system (CNS), such as RGC do not regenerate. Due to this, strategies have emerged aimed at protecting and regener ating these cells. NTF supplementation has been a promising approach but is limited by length of delivery and delivery vehicle. For this study, we tested a polymeric delivery system (sulfonated reverse thermal gel or SRTG) engineered to deliver CNTF , while also being injectable. A rat ONC model was used to determine the neuroprotective and regenerative capacity of our system. The results demonstrate that one single intravitreal injection of SRTG CNTF following ONC showed significant protection of RGC survi val at both 1 week and 2 week time points, when compared to the control groups. Furthermore, there was no significant difference in the RGC count between the eyes that received the SRTG CNTF following ONC and a healthy control eye. Intravitreal injection of the polymer system also induced noticeable axon regeneration 500 µ m downstream from the lesion site compared to all other control groups. There was a significant increase in Muller cell response in groups that received the SRTG CNTF injection following optic nerve crush also indicative of a regenerative response. Finally, higher concentrations of CNTF released from SRTG CNTF showed a protective affect on RGCs and Muller cell response at a longer time point (4 weeks). In conclusion, we were abl e to show a neuroprotective and regenerative effect of this polymer SRTG CNTF delivery system and the viability for treatment of neurodegenerations. Introduction The first question to ask when devising a treatment for optic neuropathies is, how do RGCs die as a result of optic nerve damage ? The answer to this question is unfortunately extremely

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! ! ! ! 66 complex, with several molecular pathways likely contributing to RGC death 1 . However, recent progress in this field has increased our understanding of what l eads to RGC degeneration following optic nerve (ON) injury. The answer to this question is now believed to lie with neurotrophic factors (NTFs) and their role in a healthy retina 28 . NTFs are a family of small diffusible molecules that are strongly implicated in t he control of adult neurogenesis, axon extension, proliferation, and cellular survival 36 . In a healthy eye, axon terminals are responsible for the uptake and transport of these NTFs to the RGC cell body 150 . However, when RGC axons become injured, as in glaucoma, they are una ble to transport NTFs, leaving the RGCs susceptible to apoptotic signals and subsequent cell death 1 . T hese recent discoveries in the mechanisms behind RGC death following optic nerve damage have led to the formation of new potential treatments. One su ch example is treatments aimed at neuroprotecting RGCs from cell death and loss of function 18 . NTF supplementation strategies have been one area of research extensively studied to deliver the necessary NTFs directly to RGCs and provide them with a neuroprotective effect 53 3 57 . However preliminary studies using NTF supplementation typically involve delivering free NTFs unaccompanied through a simple bolus injection 151 . Although somewhat promising, i nitial studies show that free NTFs administered through bolus injection present physiochemical instability, rapid diffusion, and a short half life 3 . Due to this, multiple high concentration injections would be required to realize the full effects of the NTFs, making the use of NTFs as it stands impractical and potentially dangerous in the clinical setting 115 . To achieve effective and controlled therapeutic NTF levels, a delivery system sustain ing NTF expression in the specific zone of interest while maintainin g NTF bioactivity must be designed. For these reasons, research surrounding polymeric delivery systems has become of particular interest for making the use of NTF therapeutics a viable

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! ! ! ! 67 treatment option for glaucoma 152 153 . An ideal polymeric delivery system would allow 1) simple injectable delivery of the NTF 2) NTF localization 3) functional properties to improve NTF stab ility, and 4) long term stable release of the NTF into the surrounding tissue. To answer these requirements, we propose the use of a multi component injectable reverse thermal gel (RTG) polymer system, poly(serinol hexamethylene urea) backbone (PSHU) conju gated to poly(N isopropylacrylamide) PNIPAAm (PSHU PNIPAAm), functionalized with sulfonate groups (Sul PSHU PNIPAAm or SRTG) (Fig S1). The first component, PNIPAAm, will provide the RTG properties and allow injectability of this polymer as well as entrapment of the NTF within the polymer matrix. The next component of this polymer system is a functionalizable PSHU backbone capable of attaching a large quantity of functional groups (18 potential linkages per molecule). In t his case, the polymer backbone will be modified with PNIPAAm as well as negatively charged sulfonate groups to make the SRTG. Modeled off of interactions between native extracellular matrix (ECM) and NTFs (e.g. heparin interacting with NTFs), this electros tatic interaction between the sulfonate groups and the positively charged NTF can protect the proteins from proteolytic degradation, preserve their bioactivity, and increase their half life 119 . This SRTG system can be administered itravitreally and provide RGCs with localized and sustain ed release of NTFs , promoting substantial neuroprotection of RGCs following ONC damage. To analyze this delivery system, we will utilize cilliary neurotrophic factor (CNTF) as a model N TF. Early studies first identified increased CNTF as a response to disease conditions or injury of retinal neurons, after axotomy, ischemia, and experimental glaucoma 38 39 . It was further discovered that CNTF works to provide neuroprotection to the RGCs as well as induce axon regeneration following injury 54 .

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! ! ! ! 68 The first aim of this work is to design and characterize an injectable polymer system that is capable of sustaining bioactive NTF release to the retinal ganglion cell layer. Lower critical solution t emperature (LCST) and energy dispersive X ray spectroscopy (EDS) will be used to analyze gelling properties and quantify the functionalization process respectively . Next, a release test with IR labeled CNTF and subsequent fluorescent analysis will be perfo rmed to determine the release profile of CNTF as released from the SRTG and RTG. ! Furthermore, the effects of the SRTG system on RGC neuroprotection and axon regeneration following ONC will be examined (with the appropriate controls) after 7 days, 14 days, and 28 days post ONC. Brn3a, growth associated protein 43 (GAP 43), and glial fibrillary acidic protein (GFAP) staining will be used to study the neurprotective effects, axon regeneration, and M Ÿ ller cell activation respectively. Materials and Methods Materials N BOC Serinol, urea, hexamethylene diisocyanate (HDI), anhydrous chloroform, and anhydrous N,N dimethylformamide (DMF), N isopropylacrylamide (NIPAAm), paraformaldehyde (PFA), sucrose, bovine serum albumin (BSA), and 4,4Ô azobis(4 cyanovaleric a cid) (ACA) were purchased from Sigma Aldrich (St. Louis, MO, USA). N (3 Dimethylaminopropyl) N ethylcarbodiimide hydrochloride (EDC), potassium tert butoxide (t BuOK), 1,3 propane sultone (PS), N hydroxysuccinimide (NHS), and trifluoroacetic acid (TFA) we re purchased from Alfa Aesar (Ward Hill, MA, USA). Anhydrous diethyl ether was purchased from Fisher Scientific (Pittsburgh, PA, USA). Anhydrous dichloromethane (DCM) was purchased from JT Baker (Phillipsburg, NJ, USA). The pentapeptide Gly Arg Gly Asp Ser (GRGDS) was purchased from Biomatik (Wilmington, DE). Dialysis tube (Spectra/ Por) was obtained from Spectrum Labs

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! ! ! ! 69 (Houston, TX). Phosphate Buffered Saline (PBS) was purchased from Thermo Scientific. Wistar rats were purchased from Charles River Laborato ries (Wilmington, MA, USA). Sterile saline, isoflurane, ketoprofen, and xylazine were purchased from MWI Veterinary Supply (Boise, ID, USA). Optimal cutting temperature (OCT) compound was purchased from Sakura (Torrance, CA, USA). Superfrost glass slides w ere obtained from Fisher Scientific (Chicago, IL). GFAP (mouse IgG1), GAP 43 (Rabbit IgG), Alexa Fluor 488 (goat anti mouse IgG), Alexa Fluor 594 (goat anti rabbit IgG), and SlowFade Diamond antifade mountant with DAPI were purchased from Life Technologies (Carlsbad, CA, USA). Triton X 100 was purchased from MP Biomedicals. Equipment Lower critical solution temperature (LCST) measurements were made on a Cary 100 UV visible spectrophotomer (Agilent Technologies, Inc., Santa Clara, CA). Elemental analysis o f non coated samples was performed by scanning electron microscope (SEM) (JEOL JSAM 6010la analytical scanning electron microscope, Peabody, MA, USA) in low vacuum (20 kV) mode, and the analysis was carried out by energy dispersive X ray spectrometry (EDX system by EDAX). 32 gauge needles used for the injection (TSK Laboratory (Tochigi shi, Tochigi ken, Japan). Tissue was sectioned using a CryoStar NX70 Cryostat. Confocal images were collected using a Nikon Eclipse Ti C2 LUN A microscope (Nikon, Tokyo) equi pped with two C2 DU3 high sensitivity PMT detectors, 4 diode lasers (405/488/561/640 nm), and a motorized microscope stage with 3 axis navigation (X, Y, and Z) controlled by the NIS Elements software package. Laser and software setting were kept constant b etween specimens and to allow for comparison of different image acquisitions.

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! ! ! ! 70 PSHU PNIPAAm fabrication To synthesize the RTG, We began by synthesizing a functionalizable biomimetic polymer backbone (PSHU, M w : 10,500) using urea, N BOC serinol, and hexamethylene diisocyanate as described previously 23 91 . PNIPAAm (M w : 11,000) was synthesized as previously described 91 and conjugated to PSHU to form the PSHU PNIPAAm copolymer. To do this, PNIPAAm COOH (0.75 g, 1.21 mmol ), EDC (5x molar excess) and NHS (5x molar excess ) were dissolved in 5 ml of anhydrous DMF and reacte d for 24 h under a nitrogen atmosphere. Purified PSHU (0.125 g/ml) was then added to the reactant mixture and allowed to react for 48 h at room temperature. The product mixture was purified by precipitation in anhydrous diethyl ether twice followed by dial ysis (MWCO: 12,000 14,000 Da) against ultra pure water for 5 days. The final product was lyophilized at 45¡C for 24 h and stored at room temperature. Sul PSHU PNIPAAM fabrication To synthesize the SRTG, sulfonation of the RTG was performed as previously described 154 . In short, PS (0.034 g, 5 mmol) and t BuOK (0.032 g, 5 mmol) were dissolved in 3 ml of anhydrou s DMF. RTG (0.1 g/ml) dissolved in anhydrous DMF was slowly added to the flask and allowed to react for 3 days at 60¡C under a nitrogen atmosphere. The product mixture was then precipitated in anhydrous diethyl ether twice and dialyzed (MWCO: 12,000 14,000 Da) against ultra pure water for 48 h at room temperature. The final product was lyophilized at 45¡C for 24 h and stored at room temperature. Animals All animal experiments were performed under a protocol approved by the Institutional Animal Care and Use Committee (IACUC) at the University of Colorado Anschutz Medical Campus. Male adult w istar rats (250 300 g ) were allowed to acclimate for 1 week prior to surgical

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! ! ! ! 71 procedures. Rats were maintained on a 14/10 hour light/dark cycle with a continuous supply of fresh air and access to food and water ad libitum. Rats were anesthetized with 5% isoflurane in oxygen and maintained on 0.5 1% isoflurane in oxygen for the remainder of the surgery. To maintain body temperature, rats were placed on a warm recirculating water blanket. Rats were divided into four separate groups according to the experiment. The time course of the ONC and subsequent treatment included three different time points (1 week, 2 weeks, and 4 weeks). To determine the percentage of RGCs surviving post ONC and axon regrowth, 3 5 rats were used per time point. Elemental analysis Polymer solutions were prepared in ultra pure water and allowed to gel at 37 ¡ C for 15 min. The gelled samples were then frozen quickly using liquid nitrogen, cut in half to expose the center structure, and rapidly transferred to a freeze dryer for 24 h ( 48 ¡ C ). An electron beam was scanned across a dried sample of both Sul PSHU PNIPAAm and PSHU PNIPAAm. When the beam strikes the sample, signals are produced that are representative of a sample's elemental composition. Solution to gel phase transition. Lo wer critical solution temperature (LCST) was used to analyze the gelling properties of the thermally reversible injectable scaffold. 1% (wt/v) PSHU PNIPAAm and PSHU PNIPAAm RGD were be loaded in a temperature controlled UV/visible spectrophotometer. The tr ansmittance through the polymer solution at 480 nm was monitored as the temperature increases from 15 ¡ C to 45 ¡ C. The rate of temperature increase was held at 2 ¡ C/min. This constant rate allowed us to observe the gelation time of both SRTG and RTG for a b etter

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! ! ! ! 72 comparison. Gelation activity was indicated by an increase in opaqueness and therefore a decrease in transmittance. IR Dye conjugation to CNTF CNTF was conjugated to IR Dye 800 CW NHS ester using a modified version of the IRDye 800 CW Microscale prot ocol provided by LI COR. Briefly, lyophilized IR Dye was reconstituted in DMSO at a concentration of 15 mg/mL and then diluted to a working concentration of 0.15 mg/mL in PBS. Lyophilized CNTF was reconstituted in 1X PBS at a concentration of 0.1 mg/mL. The IR dye solution and CNTF solution were added to a microcentrifuge tube in a 1:50 volume ratio, respectively, at pH 8.5 and reacted for 2 hours with shaking, protected from light. After the reaction, excess IR dye was removed using Zeba Spin Desalting Columns to purify the IR labeled CNTF. In vitro CNTF Release Test To a 1.5 mL centrifuge tube, 100 ng of IR labeled CNTF was added to a solution of Sulfonated PSHU PNIPAAm or PSHU PNIPAAm for a final RTG concentration of 1% (mg/mL) with samples run in triplicate. This solution was allowed to dissolve overnight at 4¡C protected from light. The following day, samples were gelled in an incubator (37¡C, 5% CO 2 ) for 5 minutes to allow stable gel formation, followed by the addition of 600 "L of 1X PBS (37¡C). At each time point, 300 "L of the release solution was gently mixed and removed from each sample and replenished with 300 "L of fresh PBS at 37¡C. To a black, 96 w ell optical bottom plat e, 100 "l of each release sample was added. The fluorescence of each sample was measured using the LI COR Odyssey Classic. Known concentrations of IR labeled CNTF were added to the plate to generate a standard curve to calculate release sample concentrat ions.

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! ! ! ! 73 Optic nerve crush model To begin, rats were given intraperitoneal injections of a solution (1ml/kg) containing 1 2 mg/kg xylazine in sterile water. A lateral canthotomy was performed after anesthetizing the eye with proparicaine and then steriliz ing the area surrounding the lateral corner of the eye with betadine. The eye was then irrigated using normal saline. Using a straight hemostat, the skin at the lateral corner of the eye was crimped all the way down to the orbit for 1 minute (to achieve he mostasis). The skin around the lateral orbit was lifted with forceps and a 0.5 cm incision was made with scissors. Blunt dissection was utilized to expose the optic nerve. Using a number 5 jeweler's forceps, the optic nerve was crimped three times, 10 sec onds each, approximately 2 mm behind the globe. Since a minimal incision was made, the procedure did not require sutures and healed properly on its own. I ntravitreal injection Following the nerve crush procedure, animals received intravitreal injections depending on the group (saline, free CNTF ( 0.5 µg ), RTG CNTF (1wt /v % polymer) ( 0.5 µg CNTF), SRTG CNTF (1wt /v % polymer) ( 0.5 µg CNTF), and SRTG CNTF (1wt /v % polymer) (2.5 µg CNTF) ). For each intravitreal injection, 5 µl of saline, 5 µl of the f ormulations was injected into the vitreous chamber using a 32 gauge, 4 mm long needle and a Hamilton glass syringe. When injected, the needle was inserted into the superior hemisphere of the eye at a 45 ¡ taking care to avoid any injury to the lens or retin a. Following injection, the needle was held in the injection site for 30 seconds to prevent leakage of the treatment and then removed slowly. Retina and optic nerve immunohistochemistry (IHC). At the appropriate time points (1 week, 2 weeks, or 4 weeks), rats were euthanized using an unchanged cage and a flow rate of CO 2 introduced to displace 20% of the cage volume per

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! ! ! ! 74 minute. A bilateral thoracotomy was performed as a secondary method of euthanasia. They eyes were then enucleated and fixed using 4 % PFA in PBS for 30 min, followed removal of the cornea and lens and an a dditional fixation period in 4% PFA for 30 min. The tissue was then cryoprotected with 10%, 20%, and 30 % sucrose in PBS for a total of 2 days, embedded in optimal cutting temperature (OCT) compound, and frozen at 80 ¡C. Longitudinal tissue sections for b oth analysis of the optic nerve (10 µ m) as well as the retina (10 µm ) were collected along the nasal temporal plane using the optic cup as a reference point. OCT was removed from the glass slides with a 5 min PBS wash. The slides were fixed in NBF for 10 m in and washed 3 times in 0.01% tween in PBS for 3 minutes each. The sections were then permeabilized in 0.01% triton X 100 in PBS for ten minutes and washed 3 times in 0.01% tween in PBS 3 times for 3 minutes each. Non specific binding was blocked by incub ating sections in 2% BSA, 0.5% triton X 100, in PBS for 1 h at room temperature. Sections were then stained with the appropriate primary antibody GFAP (1:250), GAP 43 (1:250), or Brn3a (1:100) overnight at 4¡C. All antibody dilutions were prepared in the b locking buffer. Slides were then washed with 0.01% tween in PBS 3 times for 3 minutes each and then incubated with the appropriate secondary (Alexa Fluor 488 (1:500) for GFAP, Alexa Fluror 546 (1:500) for GAP 43, and Alexa Fluror 594 (1:500) for Brn3a) pre pared in the blocking buffer for 1 h at room temperature. Finally slides were and washed 3 times in 0.01% tween in PBS for 3 min each and then 3 times in dH2O for 3 min each. Stained slides were mounted with Fluoromount G with DAPI mounting medium. Slides were stored at 4¡C. Analyzing of neuronal survival To determine the number of Brn3a positive cells following nerve crush and subsequent treatment, five images were taken from each area of each retina (two peripheral, two mid peripheral, and

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! ! ! ! 75 optic nerve c up). As the number of RGCs may vary depending on the different areas of the retina, the region of the retina that was analyzed for Brn3a positive cells was held consistent by using the optic nerve cup as an anatomical marker. Only tissue sections taken fro m around the optic head were analyzed and used for Brn3a counting. Brn3a positive cell counts from all five images were performed using Fiji. The five cell counts from each retina were analyzed by averaging these values and comparing the cell count between each group (saline, free CNTF, PSHU PNIPAAm CNTF, Sul PSHU PNIPAAm CNTF, and healthy eyes). Each data point represents the mean +/ SD of the surviving neurons post crush. Statistics were performed using a one way ANOVA Ð Tukey post hoc test, * p<0.01, ** p<0.001, ***p<0.0001 n=3 5 eyes/group/time point. Quantification of axonal growth The extent of RGC axon growth was evaluated using GAP 43 immunostaining of longitudinal sections of the optic nerve (10 µ m). Composite images of whole cross sections of th e retina and the optic nerve were imaged using 10x magnification (Nikon). Z stack maximum intensity images of 5 total stacks were taken to encompass the full section height (10 µm). Axonal growth was quantified by determining the average level of fluoresce nce in regions of nerve 100 µm, 250 µ m, 500 µ m and 1000 µm downstream from the lesion site. During analysis, the injury site was determined in the same tissue section that had been stained with GAP 43 through dark field microscopy. Average pixel intensity for each retinal section was analyzed for n= 3 5 per group per time point and expressed as mean ± S.E.M.

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! ! ! ! 76 Statistical analysis All results are expressed as means ± standard error of the mean. Analysis of variance (ANOVA) was used to determine significant differences between groups. Statistical significance was co nsidered when p < 0.05. Results Characterization of negatively charged SRTG We began by characterizing PSHU. Both 1H nuclear magnetic resonance spectroscopy (NMR) and fourier transform infrared spectroscopy FT IR (data previously reported) 23 154 were used to confirm the overall polymer structure and ensure the presence of free functionalizable amines on the backbone. Elemental analysis, through EDS, was used to confirm the sulfonation process and synthesis of SRTG 155 . The approximate mass percent of sulfonate groups chemically conjugated to the SRTG was 0.09% as compared to 0.00% seen in RTG (Fig. S2). Next, we examined the gelling properties, through LCST, of the polymer system before and aft er sulfonation. The LCST describes the temperature that the polymer system transitions from solution to solid form and is therefore an extremely important characteristic for this application, allowing for facile injection into the intravitreal space. It is possible that the introduction of sulfonate groups on the polymer backbone could alter the gelling properties of the polymer system. Fig. S3 displays the LCST of both SRTG and RTG. We can see that the sulfonated polymer system still displayed very similar gelling properties to the RTG precursor, including a similar temperature of gelation and relative rate of gelation. This is likely because the addition of the small sulfonate groups to the polymer backbone is not sufficient to alter the large hydrophobic/ hydrophilic properties of the PNIPAAm polymer that dictates the liquid to solid phase transition. Finally we analyzed the release profile of CNTF from each of the polymer types (SRTG and RTG). As discussed earlier, negatively

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! ! ! ! 77 charged sulfonate groups were conjugated to the polymer backbone to form the SRTG. The motivation behind this step was to encourage polymer protein interaction between the negatively charged sulfonate groups and the large positively charged receptor binding site on CNTF 156 . These interactions would not only preserve the integrity of t he protein but also prolong the duration of release by dampening the burst release of CNTF from the polymer system. To compare the release profiles of CNTF released from the RTG to the release profile of CNTF released from the SRTG, we took samples of rele ase buffer and analyzed these samples for fluorescently labeled CNTF using a fluorescent imaging system (Fig. S4). We can see from Fig S4 that the 1% SRTG polymer showed a decreased burst release of CNTF compared to the initial burst release from the 1% RT G polymer samples. This indicates that the increased charge diminished the initial expulsion of the CNTF from the polymer potentially leaving more CNTF entrapped within the polymer system. The decreased burst release could prolong the duration of CNTF rele ase from the SRTG lending to the superior results during in vivo studies. Delivery of CNTF from the SRTG promotes survival of RGCs post ONC We first investigated the neuroprotective effects of the NTF delivery system. To do this, the polymer system loaded with CNTF or a control treatment was applied as an intravitreal injection following ONC and remaining RGCs were analyzed using immunostaining procedures. Experimental groups included animals with ONC that received: (i) a single intravitreal injection of s aline (5 µ l) (n=5, 1 week; n=5, 2 weeks); (ii) a single intravitreal injection of free CNTF (0.5 µ g in 5 µ l saline) (n=3, 1 week; n=5, 2 weeks); ! a single intravitreal injection of SRTG (5 µ l) (n=5, 2 weeks), (iii) a single intravitreal injection of RTG CN TF (5 µ l 1 wt /v % polymer solution loaded with 0.5 µ g CNTF) (n=5, 1 week; n=5, 2 weeks); (iv) a single intravitreal injection of SRTG CNTF (5 µ l 1 wt /v % polymer solution loaded with 0.5 µ g CNTF) (n=5, 1 week; n=5, 2

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! ! ! ! 78 weeks; n=5, 4 weeks), and (v) a single intravitreal injection of SRTG CNTF (5 µ l 1 wt /v % polymer solution loaded with 2.5 µ g CNTF) (n=3, 2 weeks; n=3, 4 weeks) (Fig. 1). Figure 4.1 . Experimental protocol and animal groups. ONC was performed followed by the appropriate intravitreal injection. At one week, two weeks, or four weeks post ONC, rats from eac h group were euthanized and the tissue was analyzed (A). Both 1 week and 2 weeks time points included the groups: saline, CNTF ( 0.5 µg) , RTG CNTF ( 0.5 µg) , and SRTG CNTF ( 0.5 µg) (n=3 5). The 2 week and 4 week time points also included groups: SRTG CNTF ( 2.5 µg) (B). At the correct time points, animals from each group were analyzed for RGC survival using Brn3a immunostaining followed by confocal imaging of the ganglion cell layer. At one week post ONC, the group that received a single intravitreal injection of saline (Fig. 2A) and the group that received a single intravitreal injection of free CNTF ( 0.5 µg ) (Fig. 2B) showed a decrease in the number of RGCs. In addition, quantitative a nalysis showed that a comparable number of RGCs survived post ONC in the saline group (3.52 ± 1.945 RGCs/400 µ m retina, mean ± SEM)

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! ! ! ! 79 compared to the free CNTF group (5 ± 2.511 RGCs/400 µ m retina) (Fig. 2I). These numbers are in fact consistent with previous studies showing survival rates of RGCs 1 week post ONC with no treatment applied being around 37% 40% 157 158 , indicating that both the saline and free CNTF injections provided no significant neuroprotective effect for RGCs 1 week post ONC. We then examined if a si ngle intravitreal injection of polymer systems (RTG CNTF and SRTG CNTF) loaded with 0.5 µg of CNTF, injected at the time of nerve injury could facilitate RGC survival 1 week post ONC. Despite a slight increase, there was no statistically significant differ ence between the RTG CNTF group (5.68 ± 1.945 RGCs/400 µ m retina) when compared to both the saline treatment group as well as free CNTF (Fig. 2I). However, there was a dramatic increase in the survival of RGCs in the SRTG CNTF treatment group (9.56 ± 1.945 RGCs/400 µ m retina) (Fig. 2I). In addition, there was a statistically significant difference between the saline treatment group and the SRTG CNTF (p value: <0.0001) as well as between the free CNTF group and the SRTG CNTF groups (p value: 0.0082) (Fig. 2I ), indicating that the prolonged release of CNTF from the SRTG polymer can enhance the survival of RGCs one week post ONC. Two weeks post ONC, both groups that received a single intravitreal injection of saline (0.8 ± 1.526 RGCs/400 µ m retina) or a single intravitreal injection of free CNTF (0.5 µ g) (1.2 ± 1.526 RGCs/400 µ m retina), showed similar numbers of surviving RGCs (Fig. 2E and 2F). In addition, both of these groups showed a number of surviving RGCs consistent with previous reports on the number of surviving RGCs two weeks post ONC with no applied treatment. These expected results indicate that both the saline and free CNTF injections had no marketable neuroprotective affect two weeks post ONC but provided good controls for comparison to the polymer treatment groups. In contrast, both polymer groups showed a dramatic increase in the number of surviving

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! ! ! ! 80 RGCs post ONC (Fig. 2G and 2H). One single intravitreal injection of RTG CNTF (0.5 µ g) (6 ± 1.526 RGCs/400 µ m retina) or one single intravitreal injec tion of SRTG CNTF (0.5 µ g) (7.8 ± 1.526 RGCs/400 µ m retina) both displayed a significant difference in surviving RGCs when compared to both the saline and free CNTF groups; however the SRTG CNTF group displayed the highest mean value of surviving RGCs (Fig . 2J). To determine if the SRTG itself was having any effect on RGC survival, we included an additional group that received a single injection of SRTG alone following the ONC. Results from this study show no significant difference between an injection of S RTG and saline at the two week time point (Fig. S5), indicating that the previously described neuroprotective effects of the SRTG CNTF system are likely not due to the polymer itself.

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! ! ! ! 81 Figure 4.2 . Expression of Brn3a in the retina following ONC and the appropriate intravitreal injection. One week (A D) and two weeks (E H) retina sections from the four treatment groups (SRTG CNTF ( 0.5 µg) , RTG CNTF ( 0.5 µg) , CNTF ( 0.5 µg), and saline) are shown in th e top two rows. Blue staining indicates DAPI. Red staining indicates Brn3a expression. Scale bar = 50 µm. * p < 0.05, ** p < 0.01, *** p < 0.001. ! ! Treatment with SRTG CNTF following ONC shows preservation of RGCs comparable to that of a healthy eye Foll owing the evaluation of each treatment group as shown above, we took the analysis one step further and compared eyes treated with SRTG CNTF ( 0.5 µg) following ONC with eyes that have had no ONC and no intravitreal injection (i.e. healthy eyes). This analys is was done using a Brn3a immunostain and the same quantification process as described above. Fig. 3 shows immunostained representative zoomed images of retinal cross sections from each group. Rats

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! ! ! ! 82 that had received the ONC followed by an intravitreal inje ction of SRTG CNTF ( 0.5 µg) showed significant Brn3a expressing RGCs within the ganglion cell layer of the retina (Fig. 3A). Interestingly, healthy eyes (Fig. 3B) showed a comparable number of Brn3a expressing cells in the ganglion cell layer. Further quan tification comparing the Brn3a cell count between these two groups shows no significant difference at both the 1 week and 2 week timepoints (Fig. 3C). Figure 4. 3. Expression of Brn3a in retinal cross sections following ONC and Sul PSHU PNIPAAm intravitreal treatment after two weeks (A) when compared to a healthy eye receiving no ONC and no treatment (B). There was no significant difference observed between these two groups (C). Blue staining indicates DAPI. Red staining indicates Brn3a expressio n. Scale bar = 50µm.

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! ! ! ! 83 Delivery of CNTF from SRTG promotes growth of injured RGC axons To investigate whether the increased RGC survival seen in the polymer treatment groups correlated to greater axon regeneration, we immunostained for damaged and regenera ting RGC axons within ON cross sections using a GAP 43 immunostain 159 . This stain was not only used to help deduce the crush site but also to elucidate any regenerating axons downstream from the injury. After two weeks, controls injected with saline or free C NTF ( 0.5 µg) following ONC showed minimal amounts of axons crossing over the glial scar (Fig. 4A and 4B). Additionally, animals that received an intravitreal injection of RTG CNTF ( 0.5 µg) showed negligible RGC growth downstream from the lesion two weeks p ost ONC (Fig. 4C). In contrast, the group that received an intravitreal injection of SRTG CNTF ( 0.5 µg) post ONC, showed extensive upregulation of GAP 43 well beyond the injury site (Fig. 4D). Dark field imaging was used to elucidate the lesion site in th e same slide that received the GAP 43 stain (Fig. 4E H). Quantification analysis of axon growth showed that the expression of GAP 43 was upregulated for the SRTG CNTF group at 100 µ m, 250 µ m, and 500 µ m past the lesion site, when compared to control groups one week post ONC (Fig. 4I). For example the average pixel intensity of the GAP 43 stain seen within the SRTG CNTF group at 100 µ m, 250 µ m, and 500 µ m was 2709.06 ± 409.2 (mean ± S.E.M), 2200.0 ± 125.3, and 1962.1 ± 108.1, respectively compared to the sal ine group values of 1251.7 ± 184.1, 1383.9 ± 124.1, and 1418.5 ± 192.9 respectively. These groups showed a statistically significant difference for distances 100 µ m , 250 µ m , and 500 µ m with p values: 0.0182 , 0.0445 , and 0.0371 respectively (Fig. 4I). Quantification of two week data also showed an upregulation of GAP 43 expression for the SRTG CNTF group at 100 µ m, 250 µ m , and 500 µ m past the lesion site, when compared to control groups. More specifically, the average pixel inte nsity of the GAP 43 stain seen within the SRTG CNTF group at 100 µ m,

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! ! ! ! 84 250 µ m, and 500 µ m was 2741.3 ± 523.1 (mean ± S.E.M), 2333.2 ± 384.2, and 1903.6 ± 314.7, respectively compared to the saline group value of 1510.1 ± 467.9, 1409.2 ± 350.7, and 1183.4 ± 2 81.5 respectively. These groups showed a statistically significant difference for distances 100 µ m , 250 µ m , and 500 µ m with p values: 0.0135, 0.0135, and 0.0204 respectively (Fig. 4J). Furthermore, within the SRTG CNTF group, there was no significant diffe rence between the 1 and 2 week time points at all distances from the lesion ( 100 µ m, 250 µ m, 500 µ m, and 1000 µ m). Figure 4. 4. Intravitreal treatment of SRTG CNTF ( 0.5 µg) induced RGC axon growth following ONC. Saline, CNTF ( 0.5 µg) , and RTG CNTF ( 0.5 µg) groups showed minimal axon regrowth two weeks post ONC (A C). SRTG CNTF ( 0.5 µg) showed robust RGC axon growth two weeks post lesion (D). The injury site was distinguished using dark field microscopy (E H). Triple asterisks represent the ONC site. Aver age pixel intensity (mean ± S.E.M.) of GAP 43 stain present past the lesion site following intravitreal injection of saline, CNTF ( 0.5 µg) , RTG CNTF ( 0.5 µg) , and SRTG CNTF ( 0.5 µg) at one week post ONC (I) and two weeks post ONC (J) . Scale bar = 200 µm. * p < 0.05, ** p < 0.01, *** p < 0.001.

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! ! ! ! 85 Although there was no statistically significant difference between the groups at 1000 µ m downstream from the lesion site for both 1 week and 2 weeks post ONC, we decided to image cross sections of the nerve at 40x to detect any regeneration (Fig. 5I). After staining with GAP 43, nerve sections from each group were imaged at approximately 1000 µ m downstream from the lesion site (as identified through dark field microscopy). We can see that there is minimal GAP 43 stain ing present in groups, saline, CNTF, and RTG CNTF at both 1 and 2 week time points (Fig. 5A C and 5E G). However, the rats that received an intravitreal treatment of SRTG CNTF showed a visible increase in GAP 43 expression 1000 µ m down stream from the crus h site at both time points (Fig. 5D and 5H). Quantification was done through measuring the average pixel intensity of each section. At the one week and two week time points there was a statistical difference between the saline and SRTG CNTF group (p values : 0.0181 and 0.0004 respectively) (Fig. 5J).

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! ! ! ! 86 Figure 4. 5 . Longitudinal images of GAP 43 immunostained ON approximately 1000 µm downstream from the crush site. At both the 1 week and 2 week time points, groups saline, CNTF, and RTG CNTF showed minimal GAP 43 expression (A C, E G). At both 1 week and 2 week time points, SRTG CNTF showed increased GAP 43 expression at these downstream points (D, H). Quantification of the average pix el intensity showed a significant difference between SRTG CNTF and other groups at both 1 and 2 week time points (J). Red staining indicates GAP 43 expression. Photomicrographs were captured at 40x magnification. Scale bar = 100 µm. * p < 0.05, ** p < 0.01 , *** p < 0.001.

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! ! ! ! 87 Intravitreal injection of SRTG CNTF alters the MŸller cell response one week and two weeks post ONC Next, we wanted to investigate the effects of each intravitreal treatment on MŸller cell activation following ONC. MŸller cells are the principal glial cells residing in the retina. It is thought that MŸller cells, like other types of glial cells play important roles in the support and protection of retinal neurons fo llowing injury 160 161 . For example, it has been previously reported that injury to the ON caused by ONC can trigger morphological and cellular changes in MŸller cells, suggesting that MŸller cells play a key role in the aiding the regeneration of RGC axons 162 . Some of these changes seen from activated MŸller cells can aid in the neuroprotection and regeneration of a damaged retina 163 164 . However, it is important to note that the complete function of MŸller cells may not be limited to a regenerative effect. MŸller cell activation is also linked to the inflammatory response and while inflammation is part of the regenerative process, prolonged inflammation can lead to additional tissue damage 165 166 . ! To see resulting MŸller cell activation caused by ONC and the subsequent intravitreal treatment, we perform ed retinal immunostaining using an antibody against GFAP (Fig. 6A H). At one and two weeks post ONC, the treatment groups that received saline and free CNTF ( 0.5 µg) injections showed GFAP staining that was limited to the astrocytes and MŸller cell end fee t present within the nerve fibre layer (Fig. 6A B and Fig. 6E F). The treatment group that received and intravitreal injection of RTG CTNF ( 0.5 µg) showed an increase in MŸller cell staining with expression moving down through the retina at both the one we ek and two week time points (Fig. 6C and 6G). However, the treatment group that received an intravitreal injection of SRTG CNTF ( 0.5 µg) following ONC displayed robust GFAP labeling with a large number of MŸller cell processes observed spanning the entire retina (Fig. 6D and 6H). We further quantified these results by analyzing the

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! ! ! ! 88 fluorescence of the stained and imaged section. Figure 6I shows a significant difference between the SRTG CNTF group and the other three groups at the one week time point (p valu e 0.025). Similarly, figure 6J shows a significant different at the two week time point between the SRTG CNTF and the remaining three groups ( p value 0.0037). These results indicate that the release of CNTF from the SRTG system was able to induce a signifi cant MŸller cell response that was likely contributing to the neuroprotective and neuroregenerative effect we were seeing in this group.

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! ! ! ! 89 Figure 4.6. The response of M Ÿ ller glial cells to ONC and each subsequent intravitreal injection examined using a GFAP antibody. After one week post ONC and appropriate intravitreal injection, GFAP staining in M Ÿ ller cells is shown in saline, CNTF, RTG CNTF, and SRTG CNTF groups (A D). Quantification shows an increase in M Ÿ ller cell express ion seen in the SRTG CNTF group compared to the saline control at the one week time point (I). At two weeks post ONC and appropriate intravitreal injection, there was an increased amount of M Ÿ ller cell expression in all groups (E H). There was a significan t increase in the SRTG CNTF group compared to all other groups (J). ONL: outer nuclear layer; INL: inner nuclear layer; IPL: inner plexiform layer; GCL: ganglion cell layer; FL: fibre layer. Staining indicates GFAP expression. Photomicrographs were capture d at 20x magnification. Scale bar = 100 !m. * p < 0.05, ** p < 0.01, *** p < 0.001. Increasing the amount of CNTF can prolong the neuroprotective effect of the polymer system As mentioned earlier, we were not able to see a significant difference between the SRTG group and the remaining three groups 4 weeks post ONC when using a low CNTF amount

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! ! ! ! 90 ( 0.5 !g). To determine if the low amount of CNTF was the reason for the limited effect length, we increased the CNTF amount to a higher amount ( 2.5 !g) and analyze d the effects of SRTG CNTF (2.5 !g) at two weeks and four weeks, and compared these results to corresponding SRTG CNTF (0.5 !g) and saline control groups. Figure 7A shows the quantification following Brn3a staining to elucidate RGC survival post ONC. We ca n see from this figure that the groups SRTG CNTF (0.5 !g, 2 weeks), SRTG CNTF (2.5 !g, 2 weeks) and SRTG CNTF (2.5 !g, 4 weeks) all showed a significant difference in the number of RGCs that survived post ONC (all p values: < 0.00 0 1). In addition, we analy zed the MŸller cell response with this higher dose at extended time points. Figure 7B shows the quantification of fluorescence from the GFAP stained slides. We can see from these results that there was a significant difference between the SRTG CNTF (2.5 !g , 2 weeks) and SRTG CNTF (2.5 !g, 4 weeks) compared to the controls (p values: < 0.00 0 1), demonstrating that this increase in CNTF concentration caused a MŸller cell response well into the 4 week time point indicating longer term neuroprotection and nueror egeneation.

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! ! ! ! 91 Figure 4.7. Larger amounts of CNTF loaded into the SRTG system increased the length of survival of RGCs and expression of GFAP to the 4 week time point. After 2 weeks post ONC, both low and high CNTF groups showed a significant difference in the number of remaining Brn3a positive cells to the saline control (A, B, E). After four weeks post ONC, only the high CTNF groups showed a significant difference in the number of Brn3a expressing cells compared to the control (C, D, E). For GFAP expres sion analysis, after 2 weeks post ONC, both low and high CNTF groups showed a significant difference between the saline control (F, G, J). After 4 weeks post ONC, there was no significant difference between the low CNTF group and saline (H, J). After four weeks there was a significant difference between the high CNTF group and saline group (I, J). Red staining of figures A D indicates Brn3a expression. Green staining of figures F I indicates GFAP expression. Photomicrographs were captured at 20x magnificati on. Scale bar = 50 !m. * p < 0.05, ** p < 0.01, *** p < 0.001. Discussion It has been proposed that RGC death in glaucoma may be due to axonal transport failure of NTFs from axon tip to the RGC cell body. As these NTFs are responsible for many crucial cell

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! ! ! ! 92 functions, deprivation of NTFs can lead cells into subsequent apoptotic d egeneration 24 167 . In addition, recent developments indicate that damage to CNS neurons, such as RGCs, can cause them to re express NTF markers commonly found during development in an attempt to bolster NTF pro duction and begin to regenerate 168 169 170 . These discoveries turned the focus to new treatment strategies aimed at neuroprotecting and regenerating RGCs by delivering NTFs to injured cells. NTF supplementation therapies have been tes ted using many different and complicated methods including viral gene transfers to retinal cells causing them to overexpress NTFs 3 56 , intraocular transplantation of genetically modified cells 171 , etc. Perhaps the simplest and most straightforward method of delivering NTFs to RGCs is through a fa cile intravitreal bolus injection. However, single intravitreal injections of NTFs result in limited neuroprotection and axon regeneration due to rapid clearance, degradation, and short half life of the NTF 172 . Polymeric delivery systems could potentially mitigate the pitfalls of repeated NTF injections by providing continuous delivery of NTF to the appropriate areas. Currently there are a diverse range of polymeric materials that could aid in the delivery of drugs/NTFs to the eye 173 174 175 . I n the present study, we investigated the use of an injectable polymer system to deliver CNTF to RGCs following ONC. CNTF was chosen for this analysis because it has been shown to have a neuroprotective effect o n injured RGCs in various pathological conditions 52 176 177 . In addition, unlike brain derived neurotrophic factor (BDNF) or fibroblast growth factor (FGF), CNTF was also shown to stimulate RGC axon regeneration fol lowing injury 178 179 . CNTF was intermixed within the injectable polymeric delivery system or unconfined in free bolus and delivered intravitreally following ONC. The ONC model has been used extensively to not only study the mechanisms of RGC death but also to devise new neuroprotective and repair strategies for the injured ON 180 181 . This model generates direct trauma to RGC axons leading to a severe injury

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! ! ! ! 93 and almost absolute depletion in RGCs after 2 weeks 46 . Although this model typically represents a more severe situation than seen in glaucoma, it provided us with a highly reliable and Ôextreme case' model to study the effects of the NTF delivery system. In other words, bec ause this model is so severe, any neuroprotection or regeneration seen would be due to a robust treatment effect and thus will be more easily detected with a smaller standard deviation. We found that a single intravitreal injection of SRTG CNTF resulted i n significant preservation of RGCs in the ganglion cell layer when compared to the control groups (saline or /free CNTF) at both 1 week and 2 week time points. We consider it remarkable that this CNTF delivery system was able to rescue such a large amount of RGCs from apoptosis following a severe injury in which control groups experienced almost complete depletion of RGCs after two weeks. Additionally, rats that received an intravitreal injection of SRTG CNTF following severe ON injury retained a comparable number of RGCs to a completely healthy eye that had no manipulations or ONC performed on it. Considering the robust effect of this polymer system following an ONC which induces severe and rapid depletion of RGCs, it would be interesting to evaluate the ne uroprotective capacity of the polymer system using an animal model that more closely resembles the slow and variable loss of RGCs seen with glaucoma 182 183 184 . Along with a neuroprotective effect, CNTF has been shown to induce axon regeneration following ON injury 176 185 179 . We observed this regenerative effect of CNTF in eyes that received an intravitreal injection of the polymer system following ONC. Specifically, at both the one week and two week time points post ONC, we observed robust and sig nificant axon regeneration in mature rat RGCs that received an intravitreal injection of SRTG CNTF, whereas eyes that received intravitreal injections of saline, free CNTF, or RTG CNTF showed minimal axon regrowth past the lesion. The separation seen with the SRTG CNTF group is likely due to

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! ! ! ! 94 the addi tion of sulfonate groups, as the difference was not seen with the un sulfonated polymer (RTG CNTF). This lends to the idea that the negatively charged sulfonate groups sequestered or stabilized the CNTF before r eleasing it into the surrounding retina. Recently, it has been shown that RGC regeneration is closely correlated with activation of retinal astrocytes and M Ÿ ller cells 186 . This is likely because activated M Ÿ ller cells express CNTF as a response to nerve injury 187 39 188 , aiding in protecting RGCs from apoptosis and promoting regeneration of damaged axons 52 189 187 185 . Even mor e interesting, new developments show that CNTF also leads to activation of M Ÿ ller cells 58 , meaning that CNTF contributes to a positive feedback loop lead ing to RGC protection and axon regeneration following injury. In this study, we saw the effects of prolonged exposure of CNTF on M Ÿ ller cell activation. At both one week and two week time points, there was a large increase in GFAP staining for M Ÿ ller cells in the groups that received intravitreal injections of SRTG CNTF, extending processes spanning the entire width of the retina. In contrast, there was minimal GFAP staining present in the three other groups (saline, free CNTF, and RTG CNTF) at both time po ints. This result indicates that prolonged exposure of CNTF led to activation of M Ÿ ller cells, which could have bolstered neuroprotection of RGCs and axon regeneration by releasing additional CNTF produced by M Ÿ ller cells. Furthermore, we were able to indu ce a longer effect simply through increasing the amount of CNTF initially loaded into the polymer system. Using this higher amount of CNTF, we were able to see a significant increase in RGC survival and M Ÿ ller cell activation at the four week time point.

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! ! ! ! 95 Conclusion With this work, we aimed to show the viability of this polymer as a NTF delivery system for the treatment of retinal degenerations. However, in reality supplementation of NTFs to the injury site only resolves part of the disease state. The o ther aspect of glaucoma and other retinal degenerations is the depletion and permanent loss of RGCs. Since these cells do not regenerate, replenishing this population must involve some form of cell transplantation therapy. Due to the number of functional g roups available on the polymer backbone (PSHU), this polymer is highly modifiable and may be conjugated with more than one functional unit. For example, in a previous work, we have chemically conjugated biomolecules aimed at promoting cell survival and att achment of neurons (specifically RGCs), making this polymer system a viable scaffold for cell transplantation therapies 190 23 91 . Through this work and previous works, we have shown that this polymer system may be used simultaneously for NTF delivery and as a scaffold for cell transplantations giving a two fold approach to the treatment of retinal degenerations. In summary, we believe that this polymer system may be utilized as a multifaceted approach to prevent RGC degeneration (through NTF release) as well as provide a new RGC population (through cellular scaffolding).

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! ! ! ! 96 CHAPTER V DISCUSSION, LIMITATIONS AND FUTURE DIRECTIONS Discussion Currently, the loss of RGCs through uncontrolled glaucoma is irreversible. For many patients suffering from extensive and prolonged disease, cell replacement therapy may be the only hope of restoring lost vision. The inability of the mature CNS to regenera te and/or replace RGCs damaged by optic neuropathies, has encouraged research to look to cell replacement therapies as future treatment options 100 101 . However, cell transplantation involving RGCs is inherently challenging due to low cell transplant survival and even lower functional integration 112 191 . In order for cell transplantation t o be successful a few criteria need to be met: 1) Cells must survive the transplantation process without being rejected 2) the environment to which the cells are transplanted into must be permissive to axonal growth 3) the environment must also promote axo n growth and path finding. While considerable research has been conducted to develop a scaffold capable of encouraging cell survival post transplantation, the sensitivity of RGCs specifically and the target area of transplantation make developing a scaffo ld for these cells exceedingly difficult. As discussed above, naturally derived ECM scaffolds have been shown to have a wide array of application in tissue engineering. Their benefits stem from their ability to maintain many of the bioactive molecules that comprise native tissue, including collagens, laminins, and growth factors 105 192 . Furthermore, these scaffolds produce a much more muted immune response and many are already approved through the FDA 193 194 195 . However, it is now known that an increased immune response can actually promote optic nerve regeneration and cause surrounding ce lls to produce additional growth factors further aiding in the regeneration process 165 . In addition,

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! ! ! ! 97 naturally derived scaffolds are very difficult to chemically and mechanically modulate. Both the retina and optic nerve ECM are highly specialized tissues. Due to the lack of controllability of naturally derived scaffolds, it will be extremel y difficult to modulate them to possess the appropriate biochemical, mechanical, and morphological characteristics to match that of retinal tissue. To address the shortcomings of naturally derived cell scaffolds, ECM hydrogel technology was developed. Thes e synthetic hydrogels are highly tunable, both mechanically and biochemically, and may be altered to mimic the native ECM while also delivering drugs or growth factors. Furthermore, ECM hydrogels may be administered through a minimally invasive injection 105 196 . This characteristic is extremely beneficial for use in optic neuropathies as these cellular scaffolds are to be transplanted into diseased tissue where further disruption could propel the disease process . However one of the major drawbacks of hydrogels is their extremely rapid degradation profile which may not allow sufficient time for transplanted cells to properly integrate into the retinal tissue 197 . In addition, hydrogels typically have poor mechanical properties incapable of providing the appropriate stability for transplanted cells. On the other hand biohybrid sc affolds possess stronger mechanical properties that are required for these applications. The mechanical properties of a scaffold, particularly the scaffold's stiffness can have large implications on the behavior of encapsulated cells. Furthermore, it is im portant that the stiffness of the scaffold is somewhat proportional to the stiffness of the native ECM in order to limit disruption and encourage integration of the transplanted cells. In addition, these biohybrid scaffolds can be manipulated to possess sp ecific degradation properties as well as chemically modified to possess specific biochemical properties 198 . In this work, we first developed a biomimetic biohybrid polymer ba ckbone with the capacity to attach multiple biochemical moieties. Through this characteristic, we were able to modify this

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! ! ! ! 98 polymer backbone to posses very specific qualities suited for RGC scaffolding. We first conjugated PNIPAAm, a reversible thermal gel that transitions from solution to gel when exposed to body temperature. This property allowed us to administer this polymer scaffold, and the subsequent encapsulated cells, through simple injection. As mentioned previously, this property is crucial as to n ot damage an already damaged retina. In addition this property allowed us to implant a 3D scaffold with RGCs encapsulated and protected from the outside environment. We then conjugated RGD, an integrin binding motif, which provides the encapsulated RGCs wi th a moiety to attach to and aid in axon extension. Next, we analyzed the mechanical properties of this synthetic scaffold through rheological studies. We were able to show that the mechanical properties of this polymer system could be adjusted by simply a ltering the mass percent of polymer in solution. This allowed us to control the stiffness of the polymer so that it is most similar to the soft modulus of native retinal/neural tissue. We then investigated if we could modulate the 3D polymer scaffold to po ssess a similar morphology to native retinal tissue. As discussed previously, native retinal tissue is organized in a sheet like pattern with various cellular types aligned along one 2D sheet. One of the biggest challenges of this work was to synthesize a polymer scaffold that would self assemble into this sheet like structure while still maintaining the ability to be injected through a needle. In order to do this we synthesized a polymer scaffold with a larger proportion of hydrophobic groups that would se lf assemble into a sheet like pattern when injected into tissue. Next, through in vivo studies, we were able to show an increase in RGC survival and axon extension when encapsulated in this synthetic scaffold. However, it is not enough to improve RGC tra nsplantation survival and insulate the cells from diseased tissue through a cellular scaffold. Transplanted RGCs must also show robust axon outgrowth in order to have a chance of reestablishing connections further along the optic nerve.

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! ! ! ! 99 Several growth fact ors have been investigated as to their effects on RGCs. Specifically, it was discovered that CNTF not only exhibits an RGC survival enhancing effect but also promotes axonal outgrowth 176 165 199 . In order t o capitalize on this fact, we developed an additional polymer system capable of releasing long term CNTF into the intravitreal space. The backbone of this polymer was the same as the one used for the cell scaffold, thus maintaining biocompatible properties . However, we further conjugated the polymer with sulfonated groups aimed at sequestering CNTF within the polymer and slowly releasing the CNTF into the surrounding area. Using an ONC model, we were able to show an increase in both RGC survival and axon ex tension following an intravitreal injection of CNTF encapsulated in the SRTG polymer. This indicates that prolonged release of CNTF following optic nerve damage can preserve the RGC population with also encouraging axonal regeneration. Limitations The first limitation of this work stems from critical differences between the synthetic scaffold and the composition of native retinal tissue. Native ECM is typically comprised of glycoproteins, collagens, and hyaluronic acid; however, the synthetic scaffo ld discussed here does not mimic this composition perfectly and instead aims to mimic the characteristics of native ECM. Although it is possible to obtain similar biochemical and mechanical properties to native ECM, the synthetic polymer will likely influe nce encapsulated cells ability to differentiate, proliferate, and migrate to a different degree compared to native ECM. Furthermore, native ECM within the retina and optic nerve provides crucial signaling cues to neurons that control neuronal survival, axo nal growth, and connection formation. Synthetic polymers are unable to provide these signaling cues and change with the changing cell population to meet the needs of that tissue. Secondly, scaffolds intended for 3D culture must have the appropriate porosit y to allow cellular

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! ! ! ! 100 support and guide axonal extension while still permitting adequate diffusion/perfusion of nutrients for cell growth. Through this work we were able to show that this synthetic polymer scaffold possessed the appropriate porosity to permi t axon extension while still maintaining cellular attachments. However, during our analysis of RGC growth within the 3D polymer scaffold we saw time limitations on the number of weeks capable of maintaining this culture. This result indicated that the grow th of the encapsulated cells may have been limited by lack of diffusion of waste products out of the scaffold and diffusion of cell nutrients into the scaffold. This can cause an accumulation of toxic waste products within the cell scaffold causing RGC dea th and nutrient deprivation. One major limitation stems from the inability of CNTF SRTG system to stimulate long distance axonal regeneration. We have successfully shown that CNTF SRTG can have profound effects on RGC neuroprotection as well as effects o n axonal regeneration. However, in order for axonal regeneration to make a functional difference, RGC axons must extend from the retina to beyond the optic chiasm. In order for RGCs to have this amount of directionality during the regeneration process, the y must be propagated by some form of guidance cue. This biochemical signaling cue typically takes the form of NF guiding the RGC axon to the correct synaptic location. Research has shown that without the appropriate signaling molecules RGC axons will take various directions along the appropriate course and may not make it to the their final location 200 . Without this final synapse or connection established, the survival of the RGCs is obsolete as the functional component of these cells stems from their ability to send signals to nuclei in the brain.

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! ! ! ! 101 Future Directions Polymer development To begin, we plan to optimize the porosity of the polymer scaffold. The porosity is a critical characteristic of 3D cultures, allowing the diffusio n of proteins, gases, and nutrients within the scaffold. Along those lines an increase in porosity allows the release of toxic metabolites that will build up from increased cellular growth. Finally the porosity of a polymer scaffold can control the degree of cellular interactions, migration, and axon extension. Various factors can alter the porosity of polymers. In this case, the porosity is controlled by the proportion of co polymer conjugates and biomolecular groups attached to the polymer backbone. Thus, a balance must be found between the molecular and mechanical properties provided by these conjugates and the porosity of the final product. In future work, we plan to increase the porosity of this polymer system but still maintain a sufficient number of i ntregrin binding groups (RGD) attached to the polymer backbone. This will require step down titration of RGD conjugation to find the appropriate balance. We can further increase the porosity of the scaffold through altering the proportion of PSHU to PNIPAA m, while still maintaining the appropriate gelation properties of the final scaffold. There are numerous methods for measuring the porosity of polymers; however, perhaps the simplest is to use SEM images of the polymer scaffold cross sections following ge lation 201 202 . We plan to use this method to determine the porosity of the polymer system while varying the conjugation proportion of RGD, conjugation proportion of PNIPAAm, and concentration of dissolved polymer. Neurotrophic factor optimization Next, we wish to incorporate the use of BDNF along with CNTF into the polymer system. BDNF and its receptor TrkB have been shown to play a pertinent role in the growth and survival of

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! ! ! ! 102 RGCs 203 . Along the same lines as discussed above, disruption of the supply of BDNF following optic nerve injury may pl ay a significant role in the death of RGCs 167 9 . Furthermore, research has shown that the combination of BDNF and CNTF can have an increased effect on RGC survival and axon extension 204 . In future research, we will investigate the use of both CNTF and BDNF and compare these results to previous result using only CNTF. To do this, we will conduct experiments using the ONC model followed by an injection of BDNF and CNTF loaded within the po lymer scaffold. Outcomes will be analyzed as they were before, with an emphasis on RGC survival and axonal extension. Cell transplantation applications Finally, we wish to analyze the use of the polymer scaffold in RGC transplantation studies. As we discu ssed previously, damage to the optic nerve and retina typically ends with death of RGCs. We have shown that certain NFs may be provided to the damaged RGCs in order to protect them from cell death. However, for many patients with poorly controlled disease and those who have lost a large proportion of RGCs, cell replacement therapy remains the ultimate goal. Although cell transplantation research has been conducted in a variety of tissue types, these therapies are hindered by low cell survival and integratio n into the target tissue. Preliminary research studies have shown that RGC survival during transplantation can be aided through the use of polymer scaffolding 205 206 207 . We previously developed a 3D scaffold, providin g both topographical and biochemical cues similar to that of native retinal tissue (RTG RDG) . This 3D polymer system could provide transplanted cells with the appropriate environment and cues to stimulate survival and axon regeneration. Preliminary studies were performed to transplant RGCs encapsulated in the polymer into the subretinal space of a healthy rat. We began by performing a subretinal injection of the polymer system to determine the localization of the polymer upon gelation. If the

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! ! ! ! 103 gelation time is too rapid, the polymer will not have sufficient time to spread evenly along the subretinal space before forming a solid gel. To do this we mixed FITC with the polymer solution at room temperature and performed subretinal injections within a rat eye. Fo llowing euthanization we processed, sectioned, and imaged the retinal cross section to determine the localization of the polymer. You can see from figure 5.1 that the polymer was able to spread evenly along the subretinal space creating a nice layer before forming a solid gel. Figure 5.1. Subretinal Injections. GFP was mixed with the polymer system and injected into the subretinal space to determine the localization of the polymer. We were able to show even distribution o f the polymer along the entire subretinal space. Next we isolated primary RGCs as described in the previous chapter s . Once this cellular population was purified, we labeled them with a fluorescence live cell stain to keep track of the transplanted cells fo llowing transplantation. Next we encapsulated the labeled primary RGCs

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! ! ! ! 104 within the polymer scaffold (RTG RGD) ready for injection. For preliminary studies we performed a subretinal injection of the isolated cells within the polymer scaffold in a three healt hy rat eye s (using the contralateral eye as a control). During this study, our aim was to determine the location and survival of the transplanted cells. After two weeks , we euthani zed the rats and either sectioned and stained (brn3a, b tubulin, dapi) or fl atmounted and stained. Flatmounted retinas from this experiment are shown in Figure 5.2. As you can see, the ganglion cell layer is shown through the brn3 stain. To locate the transplanted cells, we looked for cells that were staining pink (live cell dye p rior to transplantation) as well as brn3a (to signify the cells are RGCs). You can see numerous transplanted cells survived the encapsulation and injection process. We were also able to show that some of the cells were beginning to sprout axons as shown th rough the b tubulin stain. However, we wanted to investigate whether or not the transplanted cells were localizing to the ganglion cell layer. To determine the location of the transplanted cells we used cross sections of the retina. From Figure 5.3 we were able to show incorporation of the transplanted cells into the ganglion cell layer. Future studies will include the injection of the transplanted cells within the polymer scaffo ld following optic nerve crush injury . The control for this study will be a bol us injection of free cells (without the polymer scaffold). Next we will analyze the proportion of surviving cells with and without the polymer scaffold. Furthermore we will analyze the proportion of surviving cells able to localize to the appropriate retin al layer and those cells that extend axons. A power study will be performed to determine the appropriate study size to show efficacy of the transplantation . ! In this preliminary study we were able to show that cells transplanted within the polymer system were present two weeks post injection and migrated into the appropriate retinal layer. These results also show that this injectable scaffold could be a promising delivery tool for use in future cellular replacement studies aimed at treating optic neurodege nerative diseases.

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! ! ! ! 1 05 Figure 5.2 . Flatmount retinas imaged using 3D confocal imaging. Following the appropriate injections, retinas were dissected, stained (Brn3, Beta Tubulin, and Dapi), flatmounted, and imaged. Rats that received a saline injection show no transplanted cells (A). Following injection of labeled RGCs (pink) encapsulated in the polymer, flatmount retinas show presence of transplanted cells 2 weeks post injection.

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! ! ! ! 106 Figure 5.3 . Subretinal injection can be used to deliver replacement RGCs intermixed with polymer to the retinal cell layer. Initial results show integration of transplanted cells into the ganglion cell layer two weeks post injection.

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! ! ! ! 119 205. Tomita, M. et al. Biodegradable Polymer Composite Grafts Promote the Survival and Differentiation of Retinal Progenitor Cells. Tissue Eng. 1579 Ð 1588 (2005). doi:10.1634/stem cells.2005 0111 206. Yao, Q. et al. Electrospun collagen/poly(L lactic acid co ) caprolactone) scaf folds for conjunctival tissue engineering. Exp. Ther. Med. 14, 4141 Ð 4147 (2017). 207. Dasgupta, A. et al. Functional recovery following traumatic spinal cord injury mediated by a unique polymer scaffold. 99, (2002). 208. Pennadam, S. S., Firman, K., Alexan der, C. & G—recki, D. C. Protein polymer nano machines. Towards synthetic control of biological processes. J. Nanobiotechnology 2, 8 (2004). ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! !

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! ! ! ! 120 APPENDIX A ! GPC analysis of pshu Figure A.1 . GPC refractive index (blue) and light scattering (red) curves for PSHU. Vertical axis represents signal strength (mV). Horizontal axis represents retention volume (mL) of solvent passing through the column. ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! ! !

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! ! ! ! 121 ! !"#$%& ' () ' (* ' ()+(* ' ,!-. ' ""#$ ! %&%' ! ()'*" ! ' Figure A.2 . GPC refractive index (red) and log Mw (black). The straight log Mw line is an indication of excellent polymer size separation (resolution) within the GPC column. Vertical axis represents signal strength (mV). Horizontal axis represents retention volume (mL) of solvent passing through the column.

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! ! ! ! 122 APPENDIX B HPLC standard curve and quantification of RGD conjugation ' Figure B.1 . HPLC Curve of molar ratio of RGD/free amine groups (A). Using the calibration curve for RGD (B) the conjugation of RGD to the polymer backbone shows close to 100% conjugation of RGD to the free amine groups on the polymer backbone. / ' 0 '

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! ! ! ! 123 APPENDIX C Loss modulus of pshu pnipaam rgd and pshu pnipaam Figure C.1 . Rheological properties of polymer scaffold. The loss modulus was plotted vs. temperature for various concentrations of each polymer system (PSHU PNIPAAm RGD and PSHU PNIPAAm). At physiological temperatures, the storage modulus remained above the loss modulus indicatin g dominant elastic behavior.

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! ! ! ! 124 APPENDIX D Figure D.1 . Maximum intensity projections of representative 3D fluorescent images of RGCs cultured inside PSHU PNIPAAm. 3D images of RGCs cultures inside of 5 wt% PSHU PNIPAAm were taken using a confocal microscope after 3, 5, and 7 days in culture. RGCs show no axon extension (green) at both day 3 and day 7 and minimal axon extension at day 5.

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! ! ! ! 125 APPENDIX E S ynthetic route of sul pshu nipaam Fig E. 1. Schematic of synthesis process for Sul PSHU PNIPAAm (SRTG). Step 1) synthesis of the poly(serinol hexamethylene urea) PSHU; Step 2) conjugation of PNIPAAm COOH; and Step 3) s ulfonation of free amine groups

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! ! ! %(& ! APPENDIX F C haracterization of srtg and rtg Fig F.1 . Elemental analysis of SRTG and RTG. SRTG showed a 0.09 mass% of sulfur present in the polymer compared to 0.00 mass% detected in the RTG polymer.

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! ! ! %(+ ! Fig F.2 . Temperature dependent phase transition of SRTG (blue) and RTG (red). Both polymers display similar gelling temperature while the incorporation of the sulfonate groups to the SRTG slightly increased the gelling time.

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! ! ! %(* ! Fig F.3 . In vitro release profile of CNTF released from 1% SRTG and 1% RTG. Samples were taken over a month long release test.

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! ! ! %(, ! APPENDIX G Intravitreal injection control Fig. G.1 . Expression of Brn3a in retinal cross sections following ONC and Sul PSHU PNIPAAm intravitreal treatment after two weeks (B) when compared to a eye receiving ONC and saline injection (B). There was no significant difference observed between these two groups (C). Blue staining indicates DAPI. Red staining indicates Brn3a express ion. Scale bar = 50µm. !

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! ! ! % ! DECLARATION OF ORIGINAL WORK I affirm that all work in this document is my original work. Further, I confirm that all writing is my own writing. Work from others has been cited appropriately. ___________________ ________________ Student Signature Date !"#$"#%