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Evaluation of an injectable polymeric delivery system for controlled and localized release of biological factors to promote therapeutic angiogenesis

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Title:
Evaluation of an injectable polymeric delivery system for controlled and localized release of biological factors to promote therapeutic angiogenesis
Creator:
Rocker, Adam John
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Denver, Colo.
Publisher:
University of Colorado Denver
Publication Date:

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Degree:
Master's ( Master of Science)
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University of Colorado Denver
Degree Divisions:
Department of Bioengineering, CU Denver
Degree Disciplines:
Bioengineering

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Auraria Library
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Copyright ADAM JOHN ROCKER. 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
EVALUATION OF AN INJECTABLE POLYMERIC DELIVERY SYSTEM FOR
CONTROLLED AND LOCALIZED RELEASE OF BIOLOGICAL FACTORS TO PROMOTE THERAPEUTIC ANGIOGENESIS
by
ADAM JOHN ROCKER B.A., University of Colorado Boulder, 2012
A thesis submitted to the Faculty of the Graduate School of the University of the Colorado in partial fulfillment of the requirements for the degree of Master of Science Bioengineering Program
2016


This thesis for the Master of Science degree by Adam John Rocker has been approved for the Bioengineering Program by
Daewon Park, Chair Karin Payne Danielle Soranno
Date: December 17, 2016


Rocker, Adam John (M.S., Bioengineering)
Evaluation of an Injectable Polymeric Delivery System for Controlled and Localized Release of Biological Factors to Promote Therapeutic Angiogenesis Thesis directed by Assistant Professor Daewon Park
ABSTRACT
Cardiovascular disease remains as the leading cause of death worldwide and is frequently associated with partial or full occlusion of coronary arteries. Currently, angioplasty and bypass surgery are the standard approaches for treating patients with these ischemic heart conditions. However, a large number of patients cannot undergo these procedures. Therapeutic angiogenesis provides a minimally invasive tool for treating cardiovascular diseases by inducing new blood vessel growth from the existing vasculature. Angiogenic growth factors can be delivered locally through gene, cell, and protein therapy. Natural and synthetic polymer growth factor delivery systems are under extensive investigation due their widespread applications and promising therapeutic potential.
Although biocompatible, natural polymers often suffer from batch-to-batch variability which can cause unpredictable growth factor release rates. Synthetic polymers offer advantages for growth factor delivery as they can be easily modified to control release kinetics. During the angiogenesis process, vascular endothelial growth factor (VEGF) is necessary to initiate neovessel formation while platelet-derived growth factor (PDGF) is needed later to help stabilize and mature new vessels. In the setting of myocardial infarction, additional antiinflammatory cytokines like IL-10 are needed to help optimize cardiac repair and limit the damaging effects of inflammation following infarction. To meet these angiogenic and antiinflammatory needs, an injectable polymer delivery system created from a sulfonated reverse
m


thermal gel encapsulating micelle nanoparticles was designed and evaluated. The sulfonate groups on the thermal gel electrostatically bind to VEGF which controls its release rate, while the micelles are loaded with PDGF and are slowly released as the gel degrades. IL-10 was loaded into the system as well and diffused from the gel over time. An in vitro release study was performed which demonstrated the sequential release capabilities of the polymer system. The ability of the polymer system to induce new blood vessel formation was analyzed in vivo using a subcutaneous injection mouse model. Histological assessment was used to quantify blood vessel formation and an inflammatory response which showed that the polymer delivery system demonstrated a significant increase in functional and mature vessel formation while significantly reducing inflammation.
The form and content of this abstract are approved. I recommend its publication.
Approved: Daewon Park
IV


ACKNOWLEDGEMENTS
I would like to express my appreciation to my professors, mentors, friends, and family for all of their support throughout my research project. Completing this degree would not have been possible without them.
I would like to thank my advisor, Dr. Daewon Park, for his guidance and for allowing me the opportunity to work in the Translational Biomaterials Research Laboratory. Dr. Park has always pushed me to do my best, which has helped lead me through this journey. I would also like to express my gratitude to Dr. Karin Payne and Dr. Danielle Soranno for their insight and teachings throughout my graduate school career.
I truly could not be more thankful for all the assistance and encouragement I received from my fellow labmates, especially David Lee, James Bardill and Melissa Laughter. I will always cherish the times we spent together in and out of lab, and I look forward to seeing what the future holds for us.
Finally, I would like to express my gratitude for my family. My dad has been the greatest role model in my life and I could not have done this without his constant motivation. My mom has always been there for me and she was always available when I needed help with anything throughout my life. To my brother and sister, thank you for providing me with continuous encouragement and challenging me to do my best in life.
Animal model studies were conducted under the University of Colorado at Denver Institutional Animal Care and Use Committee (IACUC) protocol number 102913(12)2D.
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TABLE OF CONTENTS
CHAPTER
I. INTRODUCTION...........................................................1
Overview...............................................................1
Clinical Motivation....................................................2
Anatomy and Physiology.........................................6
Cardiovascular Disease.................................................8
Pathophysiology........................................................9
The Inflammatory Response in Myocardial Infarction............10
Current CAD Treatment.................................................11
Study Objective.......................................................13
II. Background............................................................14
Therapeutic Angiogenesis Using Growth Factors.........................14
Therapeutic Angiogenesis......................................14
Growth Factor Considerations..................................15
Anti-inflammatory Factor Considerations.......................16
Current Biological Factor Delivery Methods....................16
Polymeric Biomaterials for Protein Delivery...........................17
Sustained Growth Factor Release...............................18
Modulating Physical and Chemical Interactions.................18
vi


Thermally Induced Gelling Systems
19
Nanoparticles for Sequential Growth Factor Release..................21
Nanoparticle Considerations.................................21
III. PREVIOUS WORK.......................................................24
SRTG LCST...........................................................24
Cytotoxicity........................................................25
BSA Release Test....................................................26
CD Spectroscopy.....................................................28
IV. HYPOTHESIS AM) SPECIFIC AIMS........................................29
Hypothesis..........................................................29
Specific Aims.......................................................29
V. MATERIALS AND METHODS...............................................30
Materials...........................................................30
Equipment...........................................................31
Polymer Synthesis...................................................32
N-BOC serinol synthesis.....................................32
Poly(serinol hexamethylene urea) or PSHU backbone synthesis.32
PSHU deprotection...........................................33
PNIPAm Synthesis............................................33
PSHU-PNIPAm Conjugation (RTG Synthesis).....................34
vii


Sulfonation of PSHU-PNIPAm (SRTG Synthesis)
34
PEG-PSHU-PEG Micelle Nanoparticle Polymer Synthesis..........35
Micelle Fabrication..........................................35
Polymer Characterization.............................................36
Elemental Analysis...........................................36
Dynamic Light Scattering (DLS)...............................36
Scanning Electron Microscopy (SEM)...........................36
Proton Nuclear Magnetic Resonance (^H NMR)...................37
Fourier Transform Infrared (FTIR) Spectroscopy...............37
In Vitro Biological Factor Release Study.............................37
Enzyme-Linked Immunosorbent Assay (ELISA)....................38
In Vivo Mouse Subcutaneous Biological Factor Injections..............39
Euthanasia and Tissue Harvest........................................40
Histology............................................................40
Quantitative Analysis of Immunofluorescent Staining..................41
Statistical Analysis.................................................41
VI. RESULTS AND DISCUSSION...............................................42
Synthesis of Sulfonated-PSHU-PNIPAm Reaction Sequence................42
Synthesis of PEG-PSHU-PEG Reaction Sequence..........................43
PSHU and dPSHU Characterization 'll NMR..............................44
viii


PSHU-PNIPAm and PEG-PSHU-PEG Characterization Using 'H NMR........45
PSHU-PNIPAm and Sul-PSHU-PNIAPAm Characterization Using FTIR......46
SRTG Sulfonation Detection Using Elemental Analysis................47
PEG-PSHU-PEG Micelle Nanoparticles Size Distribution Using DLS....48
Morphology of SRTG Embedded with NPs Using SEM....................49
In Vitro Multiple Biological Factors Release Study................50
In Vivo Mouse Subcutaneous Biological Factor Injections............51
IHC Analysis of Therapeutic Angiogenesis Response..........52
Functional Vascular Endothelial Cell Analysis..............52
Mature Blood Vessel Formation Analysis.....................56
IHC Analysis of Immune Response............................60
VII. CONCLUSION.........................................................62
VIII. CHAPTER VIII.......................................................64
IX. FUTURE WORK........................................................64
Increase Sample Size...............................................64
Additional Time Points.............................................64
Increase Biological Factor Loading Amount..........................64
Acute Myocardial Infarction Animal Model...........................65
REFERENCES..................................................................66
IX


LIST OF FIGURES
FIGURE
1. Molecular basis of vessel branching [13]...........................................4
2. Structure of Blood Vessels, (a) Arteries and (b) veins share the same general features, but the walls of arteries are much thicker because of the higher pressure at which blood flows through them, (c) A micrograph shows the
relative differences in thickness [25]............................................7
3. (A) Progression of heart failure from a coronary artery occlusion. (B) The
process of therapeutic induced angiogenesis from local GF release [9].............9
4. Simplified schematic of diversity of lesions in human coronary
atherosclerosis [32]..............................................................10
5. Diagram of Angioplasty [40]........................................................12
6. Diagram of CABG surgery [41].......................................................12
7. Sulfonate group exhibiting a negative charge. R denotes an attached carbon
bound to sulfur, as well as the remaining chemical makeup of the polymer..........19
8. Polymer solubility behavior at the LCST. Left hand side shows hydrated polymer below LCST with entropic loss of water and chain collapse above
LCST (right hand side) [69].......................................................20
9. Schematic representation of different NP systems [71]..............................23
10. Temperature-dependent phase transition of sulfonated PSHU-PNIPAm. A:
The sulfonated PSHU-PNIPAm undergoes a sharp, reversible phase transition around 32 C as determined by UV-Visible spectroscopy. B: An aqueous solution of sulfonated PSHU-PNIPAm at room temperature (C)
turns to physical gel at 37 C [22]...............................................24
11. In vitro cytotoxicity of sulfonated PSHU-PNIPAm by MTT assay. Results
demonstrated no cytotoxic effects of sulfonated PSHU-PNIPAm on C2C12 cells after exposure to the polymer extract in medium, while 10% DMSO shows significant cytotoxicity. There is no statistical difference between medium and extracts. Data represent mean SD. indicates p-value < 0.05 (Students t-test) [22]...........................................................26
x


12. BSA release profiles from sulfonated PSHU-PNIPAm and plain PSHU-PNIPAm: sulfonated PSHU-PNIPAm 10%, sulfonated PSHU-PNIPAm 15%, plain PSHU-PNIPAm 10%, plain PSHU-PNIPAm 15%. BSA release was more sustained from sulfonated PSHU-PNIPAm than that from plain PSHU-PNIPAm, with a more sustained profile at higher
concentrations. Data represent mean SD [22].............................27
13. CD spectra of native and released BSA. The similarity of the protein conformation between samples indicates that the secondary structure of the
BSA protein was well preserved by the system [22]..........................28
14. Formed gels in 2 mL vials for in vitro release study.......................38
15. Demonstration of subcutaneous injection in the middle back of a mouse [82].40
16. Synthesis Reaction sequence of sulfonated PSHU-PNIPAm synthesis [22].......42
17. Reaction sequence of PEG-PSHU-PEG synthesis................................43
18. *H NMR spectrum of PSHU confirming the molecular structure [83]............44
19. *H NMR spectrum of PSHU and dPSHU confirming the removal of the BOC
protecting group with the loss of the b peak [83]........................44
20. *HNMR spectrum of PSHU-PNIPAm. Successful conjugation of PNIPAm was confirmed by the presence of methylene and methyl protons at 1.55 and
1.09 ppm, respectively...................................................45
21. *H NMR spectra of PEG-PSHU-PEG. Successful conjugation of PEG was confirmed by the presence of the peak at 3.51 ppm which identifies the
protons on the PEG repeating unit of the polymer backbone [84]...........45
22. FTIR spectra of sulfonated-PSHU-PNIPAm and PSHU-PNIPAm...................46
23. FTIR spectra of sulfonated-PSHU-PNIPAm and PSHU-PNIPAm with an
enlargement of the sulfonate peak........................................47
24. Size distribution of micelles from DLS measurements......................48
25. SEM images of the SRTG and micelle NPs. A: 5% (w/v) of SRTG cross section showing polymer sheets (scale bar =10 pm). B: Micelle nanoparticles confirming the spherical structure (scale bar = 2 pm). C: SRTG encapsulating micelles cross section showing porous configuration (scale bar = 20 pm). D: Enlargement of SRTG encapsulating micelles with
black arrows indicating micelles (scale bar = 1 pm)............................49
26. Cumulative release profile showing the sequential release of all three factors
from the SRTG and plain RTG (n=3 samples)......................................50
xi


27. Representative images of IHC co-staining for ECs (CD31) and blood vessel functionality (vWF) after 7 and 21 days for all 6 groups. Subcutaneous tissue was stained with: CD31 and Alexa Flour 488 and appear green, vWF and Alexa Flour 594 and appear red, and DAPI which appears blue. Scale
bar represents 100 pm..............................................................53
28. Comparison of CD31+ cells between the 6 injected groups quantified from co-staining of CD31 and vWF. Error bars represent standard error of the
mean. indicates p < 0.05.........................................................54
29. Comparison of CD31+ and vWF+ cells between the 6 injected groups quantified from co-staining of CD31 and vWF. Error bars represent standard
error of the mean. indicates p < 0.05............................................55
30. Representative images of IHC co-staining for ECs (CD31) and SMCs (a-SMA) to show blood vessel maturation after 7 and 21 days for all 6 groups.
Subcutaneous tissue was stained with: CD31 and Alexa Flour 488 and appear green, a-SMA and Alexa Flour 594 and appear red, and DAPI which
appears blue. Scale bar represents 100 pm..........................................57
31. Comparison of CD31+ cells between the 6 injected groups quantified from co-staining of CD31 and a-SMA. Error bars represent standard error of the
mean. indicates p < 0.05.........................................................58
32. Comparison of a-SMA+ cells between the 6 injected groups quantified from co-staining of CD31 and a-SMA. Error bars represent standard error of the
mean. indicates p < 0.05.........................................................59
33. Representative images of IHC staining for macrophages (CD68) to show immune response after 7 and 21 days for all 6 groups. Subcutaneous tissue was stained with CD68 and Alexa Flour 594 and appear red and DAPI
which appears blue. Scale bar represents 50 pm.....................................60
34. Comparison of CD68+ cells between the 6 injected groups, bars represent
standard error of the mean. indicates p < 0.05...................................61
xii


LIST OF TABLES
TABLE
1. Elemental analysis of PSHU-PNIPAm (RTG) and sulfonated-PSHU-
PNIPAm (SRTG). Polymers were analyzed for carbon, hydrogen, nitrogen, oxygen and sulfur content and were determined by weight percent for each element................................................................48
xm


LIST OF ABBREVIATIONS
'hnmr proton nuclear magnetic resonance
ABTS 2,2'-Azino-bis(3-ethylbenzothiazoline-6-sulfonic acid)
ACA 4,4-azobis(4-cyanovaleric acid)
ANOVA analysis of variance
bFGF basic fibroblast growth factor
BSA bovine serum albumin
CABG coronary artery bypass graft surgery
CAD coronary Artery Disease
GF growth Factor
CVD cardiovascular disease
DLS dynamic light scattering
DMF N,N-dimethylformamide
DMSO dimethyl sulfoxide
dPSHU deprotected PSHU
EC endothelial cell
ECM extracellular matrix
EDC N-(3-dimethylamino-propyl)-N'-ethylcarbodiimide hydrochloride
ELISA enzyme-linked immunosorbent assay
FTIR fourier transform infrared spectroscopy
GF growth factor
H&E haematoxylin and eosin
HDI hexamethylene diisocyanate
me immunohi stochemi stry
IL-10 interleukin 10
IR infrared
LCST lower critical solution temperature
MI myocardial infarction
MMP matrix metalloproteinase
MTT 3-(4,5-Dimethylthiazol-2-yl)- 2,5-diphenyltetrazolium bromide
MW molecular weight
n2 nitrogen
nh2 amine group
NHS N-Hydroxysuccinimide
NIP Am N-i sopropyl acryl ami de
nm nanometers
NP nanoparticle
OCT optimal cutting temperature
PBS phosphate buffered saline
PDGF platelet-derived growth factor
xiv


PEG poly(ethylene glycol)
PLGA poly(lactic-co-glycolic acid)
PNIPAm poly (N-i sopropyl acryl ami de)
PS 1,3-propane sultone
PSHU Poly(serinol hexamethylene urea)
rpm revolutions per minute
RSD relative standard deviation
RTG reverse thermal gel (used interchangeably with PSHU-PNIPAm)
SEM scanning electron microscopy
SMC smooth muscle cell
SRTG sulfonated reverse thermal gel (used interchangeably with sulfonated-PSHU-PNIPAm)
t-BuOK potassium tert-butoxide
TFA trifluoroacetic acid
TNF-a tumor necrosis factor-alpha
UV ultraviolet
VEGF vascular endothelial growth factor
vWF von Willebrand factor
a-SMA alpha-smooth muscle actin
XV


CHAPTER I
INTRODUCTION
Overview
The accumulation of plaque in the coronary arteries of the heart, accrued from adolescence through adulthood, often leads to a blockage in blood flow which results in a myocardial infarction (MI). This deadly disease, termed coronary artery disease (CAD), is the most common type of heart disease and claimed the lives of over 370,000 individuals in 2010 in the United States [1], Due to the abundant number of people this illness affects, the projected total costs of CAD in 2015 are around $182 billion in the United States with the cost of healthcare services, medications, and lost productivity taken into account [2], This is projected to increase to $322 billion by 2030 [2], Although some non-invasive treatments for CAD exist, such as medications and lifestyle changes, many individuals still must undergo costly surgical interventions. However, there is a subset of patients that are unable to receive these surgical procedures due to the presence of various comorbidities. These patients, along with others seeking minimally invasive treatment options for CAD, have influenced a push in research towards developing growth factor (GF) delivery systems for therapeutic angiogenesis in the pursuit of revascularizing the damaged heart tissue after a MI.
An injectable biological factor polymer delivery system that can revascularize heart tissue could dramatically reduce healthcare costs related to CAD, which would also improve the quality of life for millions of people suffering from this disease. The process of treating CAD can involve many different tests and expensive medications, in addition to the almost inevitable open heart surgery that must be implemented. A standard injectable protein delivery system to stimulate new blood vessel formation around a blocked artery will allow
1


many people to avoid paying these burdensome healthcare bills. Incorporating a supplementary anti-inflammatory protein into this system would also minimize myocardial necrosis and optimize cardiac repair following MI. Additionally, the polymer system may help with reducing the number of heart attacks experienced by patients because the delivery system could be injected before a complete occlusion of a coronary artery occurs, further reducing healthcare costs. Despite advances in polymer delivery systems for CAD treatment, localized and controlled release of proteins to revascularize damaged heart tissue and reduce the detrimental effects of an inflammatory response remain inadequate.
Clinical Motivation
Cardiovascular disease (CVD) is listed as the number one cause of death worldwide by the World Health Organization. This disease represents a myriad of issues pertaining to human health, the medical field, and the economy as this disease causes high morbidity and mortality rates while adding to rising healthcare costs [1], CAD, a type of CVD, is characterized by the accumulation of plaque in the coronary arteries of the heart that often leads to a blockage in blood flow which results in a MI, commonly caused by atherosclerosis [1-3], The insufficient blood supply to a region of the heart during a MI causes cell death and pathological remodeling which often leads to heart failure [1,2,4]. In an attempt to prevent heart failure and resupply blood flow to the damaged heart muscle, much research has been conducted on therapeutic angiogenesis with delivering GFs to the infarct site [3,5-8],
Therapeutic angiogenesis aims to form new blood vessels from the existing vasculature in order to restore blood flow to the affected ischemic heart tissue [3,8,9], This therapy involves the delivery of GFs to the damaged heart tissue through various methods, with certain studies having reached clinical trials. However, many studies have failed to show
2


efficacy for this angiogenic process due to a failure in providing sequential and long term delivery of GFs [5,7,10], Clinical trials using bolus injections of single GFs were unsuccessful due to a loss in the bioactivity of the protein in addition to a lack of other supportive GFs necessary for stable angiogenesis [7,11,12], Further research on understanding the complicated mechanisms behind angiogenesis is needed to develop an effective therapeutic angiogenesis system.
The main mechanism behind the sprouting of new capillaries works through the release of GFs, such as vascular endothelial growth factor (VEGF) and basic fibroblast growth factor (bFGF), which signals for the migration and proliferation of endothelial cells to differentiate into guiding tip cells or proliferating stalk cells [13], Both cell types work together to form directional, elongating, new vessels. Later, a maturation process occurs involving pericytes (mural cells) which are stimulated by other GFs to cover the endothelial cells [14], VEGF and PDGF are well studied GFs and have been tested in human clinical trials [15,16], Although these are key factors in the angiogenesis process, the GFs delivered alone may result in immature blood vessel formation which may regress over time [17], Additionally, it has been demonstrated that when presented simultaneously, early-stage factors can have hindering effects on late-stage GFs, and vice versa [18-20], Therefore, it is essential that the delivery system administers the GFs sequentially, while in their bioactive conformations to mimic their physiological mechanism during angiogenesis. The molecular process of new vessels branching from the existing vasculature is shown in Figure 1. The various GFs involved in this complex mechanism are shown, along with how individual factors affect each step in the process and how they influence the behavior of different cell types.
3


a Selection of tip cell
Quiescent vessel
Flow
Loosening junctions
(VE-cadherin)
Matrix remodelling (MMPs)
Tip-cell formation (VEGFR-2, DLL4, JAGGED1, NRP1, integrins, HIF-la, MT1-MMP, PGC-la)
Angiogenic factors (VEGF, VEGF-C, FGFs, ANG-2, chemokines)
Pericyte detachment
(ANG-2)
Permeability, vasodilation and extravasation (VEGF)
b Stalk elongation and tip guidance
Lumen formation (VE-cadherin, CD34, sialomucins, VEGF)
Pericyte recruitment (PDGF-B, ANG-1, NOTCH, ephrin-B2, FGF)
Tip-cell guidance and adhesion (semaphorins, ephrins, integrins)
Liberation of angiogenic factors from ECM '(VEGF, FGFs)

Flow Stalk elongation
(VEGFR-1, NOTCH, WNT, NRARP,
PIGF, FGFs, EGFL7)
Myeloid cell recruitment Adjacent vessel (ANG-2, SDF-lct, sprout PIGF)
C Quiescent phalanx resolution
Transendothelial lipid transport (VEGF-B)
Vascular maintenance
Barrier formation
Phalanx cell (PHD2, HIF-2a, VE-cadherin, TIE-2)
Flow deposition (TIMPs, PAI-1)
ephrin-B2, ANG-1,
NOTCH, TGF-pi)
Figure 1 Molecular basis of vessel branching [13],
To implement this spatiotemporal aspect while protecting the bioactivity of VEGF and PDGF, a controlled polymeric delivery system composed of nanoparticles encapsulated
4


within a sulfonated reverse thermal gel (SRTG) is proposed. The SRTG is sulfonated to mimic heparin sulfate function, which is a proteoglycan with an intrinsic negative charge that stores, protects, and stabilizes positively charged heparin-binding proteins such as VEGF [21,22], This will allow for hydrogen and ionic bonding interactions between VEGF and the sulfonate groups to reduce the burst-release of the GF while also protecting the GF from degradation. The thermal gelling properties will be utilized to encapsulate the nanoparticles which provides the sequential release of PDGF. The SRTG is liquid at room temperature, where it will be mixed with the nanoparticles containing PDGF, then once it is injected at body temperature, the polymer will rapidly gel causing the nanoparticles to be entrapped.
As this system is intended to revascularize myocardium following a MI, a third protein IL-10 will be released from the polymer scaffold. IL-10 will help minimize the deleterious effects of myocardial necrosis and promote optimal cardiac repair by reducing the inflammatory response after MI. It has been shown that this cytokine may have a role in suppressing the acute inflammatory response and in modulating extracellular matrix metabolism [23], Harmful cardiac remodeling is the leading cause of heart failure and death, and this delivery system may promote more effective tissue repair which could help reduce compensatory remodeling [24], This interleukin also provides added benefits in that it can help lessen the immune response induced by the injection of the foreign polymer delivery system. Due to its biodegradable properties, the thermal gel will degrade over time, releasing the proteins, while also releasing the nanoparticles, to provide the sequential delivery of PDGF.
5


Anatomy and Physiology
Systemic arteries provide oxygen rich blood to the bodys tissues, while the deoxygenated blood is returned to the heart through systemic veins. These different types of vessels vary slightly in their structures but generally have the same features. Arteries and arterioles have thicker walls than veins and venules as they are closer to the heart and must receive blood from a large aortic flow at far greater pressure (Figure 2A-B). Blood flows through the lumen of these vessels and arteries have smaller lumens than veins to maintain the pressure of blood moving through the system. These combined characteristics give arterial lumens a more rounded appearance in cross sections than the lumens of veins (Figure 2C) [25-27],
Both arteries and veins have the same three distinctive tissue layers which are called tunics. The tunica layers, from the most outer layer to the interior, are the tunica externa, the tunica media, and the tunica intima. The tunica intima consists of an endothelium layer which appears smooth in veins and wavy in arteries due to the constriction of smooth muscle cells (SMCs). The tunica media generally consists of SMCs and elastic fibers in arteries and this layer tends to be the thickest in these vessels. For veins, the tunica media also consists of SMCs but is mostly filled with collagenous fibers and is normally thinner than the tunica externa. The tunica externa is the thickest layer in veins, and contains some SMCs with predominately collagenous and smooth fibers. Arteries tend to have a thinner tunica externa compared to the tunica media, except for in the largest arteries, and this layer contains collagenous and elastic fibers. Small blood vessels consist only of endothelial cells (ECs), while larger vessels are surrounded by mural cells. These mural cells consist of pericytes in medium-sized vessels and SMCs in large vessels [13,25,28],
6


Artery
Vein
Tunica externa Tunica media Tunica intima
Smooth muscle
Internal elastic membrane
Vasa vasorum
External elastic membrane Nervi vasorum Endothelium Elastic fiber
Tunica externa
Tunica media Tunica intima
Vasa vasorum Smooth muscle
Endothelium
(a)
(b)
(c)
Figure 2 Structure of Blood Vessels, (a) Arteries and (b) veins share the same general features, but the walls of arteries are much thicker because of the higher pressure at which blood flows through them, (c) A micrograph shows the relative differences in thickness [25],
7


Cardiovascular Disease
CVD represents a myriad of issues pertaining to human health, the medical field, and the economy as this disease causes high morbidity and mortality rates while adding to rising healthcare costs [7], In the United States, the prevalence of this disease is stunning with an estimated one in three adults having one or more types of CVD [3], Coronary artery disease (CAD) is one of several classifications of CVD where blood flow in the coronary arteries becomes blocked by excess plaque build-up [7], This blockage prevents blood flow into the vascular heart tissue, leading to symptoms that can range from mild angina and shortness of breath to more severe issues such as a myocardial infarction if the CAD goes unnoticed for many years [7],
Once the local oxygen supply decreases significantly from a coronary artery occlusion, the tissue will respond to hypoxia by increasing the transcription of proangiogenic factors, cytokines, and matrix metalloproteinases (MMPs). The myocardium tries to restore oxygen supply and replace the damaged tissue, but frequently these adaptive responses are not effective and myocardium hypertrophy occurs. This process causes a permanent injury that would lead to heart failure if left untreated (Figure 3 A) [9], The extent of blood vessel damage and the progression of the disease impacts the treatment plan for the patient. If a local controlled release of angiogenic factors is implemented following the heart injury, the natural process of remodeling and angiogenesis would be enhanced. Myocardial functional recovery and effective revascularization could ultimately be achieved over time using this factor-based treatment approach (Figure 3B) [9],
8


rtiigralicn
(sprouting)
Figure 3 (A) Progression of heart failure from a coronary artery occlusion. (B) The process of therapeutic induced angiogenesis from local GF release [9],
Pathophysiology
Lack of blood flow to the myocardium is characterized as myocardial ischemia. Occlusion of the coronary vessels in CAD is almost always a result of atheromatous plaque narrowing the arteries [29], Beginning in childhood, early atheroma starts to develop and eventually leads to mature, cholesterol-rich, subintimal plaque formation over time [30], Atheroma mainly affects the intima layer of the vessel and can be attributed to certain risk factors such as smoking, hypertension, obesity, diabetes, hypercholesterolaemia, sedentary lifestyle, excessive alcohol intake, genetic predisposition for CVD, and others [31,32], The mature plaque is made of two components produced by two different cell types. Necrotic foam cells which migrate into the intima and ingest lipids forms the lipid core. While the connective tissue matrix is resultant from SMCs, which have migrated from the media to the intima, where they expand and change their phenotype to form a fibrous envelope around the lipid core [32], If the atheroma is only partially blocking the lumen of the artery, then the downstream physiological effects are generally mild and can range from asymptomatic ischemia to angina and a decrease in exercise tolerance [29], In more life-threatening
9


situations, acute coronary events can arise when thrombus development follows disruption of a plaque (Figure 4).
Type of Lesion
Stenotic
Fm
nbrotfc Thick Cap
Less Compensatory Enlargement
Non-Stenotic Infarction
Many
Lipid-itch
Clinical
Manifestation
Ischemia
Krqmi Pctom
Povttv* E*reu* Tm
Pvrfinion Ovfact
Thin Cap
Compensatory Enlargement
Management
Local Therapy/ Revascularization
PICA St*nt
CABG
Systemic Therapy
Lifestyle Modification
Drug Therapy
Figure 4 Simplified schematic of diversity of lesions in human coronary atherosclerosis [32], Intimal injury leads to removal of the thrombogenic matrix and generates thrombus formation [29], This thrombus formation induces a MI which prompts acute cardiac failure. Cardiomyocytes not receiving adequate blood flow in the infarct region can become necrotic within hours of the event [30], In an attempt to reduce the number of apoptotic cells from the MI, or as a way to prevent the MI altogether, the potential of delivering GFs in a local and controlled manner to promote therapeutic angiogenesis to increase blood flow to proximate cells was investigated.
The Inflammatory Response in Myocardial Infarction
Determining ideal strategies for the minimization of myocardial necrosis and optimization of cardiac repair following an MI is one of the most important therapeutic targets of modem cardiology [23,24], Cardiac pathophysiological conditions including MI and ischemia reperfusion injury leading to heart failure have been linked with stimulation of
10


inflammatory mediators in the heart [33-35], Myocardial necrosis after MI generates complement activation and free radical formation, which triggers a cytokine cascade initiated by the release of tumor necrosis factor-alpha [23], Ischemia reperfusion injury leading to heart failure is associated with intense inflammation as chemokines, cytokines, and the complement system recruit neutrophils to the ischemic and reperfused area [33-35],
Expression of pro-inflammatory cytokines such as tumor necrosis factor-a, interleukin-1, and interleukin-6, and anti-inflammatory cytokine IL-10, all facilitate homeostasis within the heart in response to injury [33], However, long-term expression of these inflammatory mediators at abundantly high levels could lead to an adverse consequence in the failing heart [36,37], Sustained inflammatory response connected with increased MMPs production may lead to excessive ECM degradation in the early phase of MI, which can impair infarct healing and aggravate early remodeling which in turn causes cardiac rupture [38,39], Moreover, the increased cytokine gene expression from inflammation causes a secondary, self-sustaining autocrine and paracrine GF and cytokine expression [33], Therefore, reducing the sustained inflammatory response from MI and ischemia reperfusion injury will be beneficial for promoting optimal cardiac repair.
Current CAD Treatment
Treatment options for CAD involve lifestyle changes, pharmaceutical drugs, and surgical interventions [9], Healthy lifestyle changes can provide the simplest means to reversing this fatal disease. These changes include regular exercise, eating foods low in fat and cholesterol, and reducing stress [3], Various pharmaceutical drugs can be used to treat CAD by reducing blood pressure and slowing the heart rate [32], Cholesterol-modifying medications reduce the amount of low-density lipoprotein cholesterol in the blood which can
11


reduce the plaque build-up in the arteries [32], Invasive surgical treatment options consist of angioplasty (Figure 5) and coronary artery bypass graft surgery (CABG) [32],
Detlated Inflated balloon balloon compresses plaque in artery against artery walls
Figure 5 Diagram of Angioplasty [40],
These surgical options aim to revascularize the ischemic tissue by compressing the plaque in the blocked artery with a catheter/balloon system (angioplasty) or by grafting a new coronary vessel to bypass the blocked artery through the CABG procedure. The latter treatment requires open heart surgery and is generally only implemented in patients with
multiple narrowed coronary arteries (Figure 6) [32],
Figure 6 Diagram of CABG surgery [41],
Aorta
Bypass vein graft
Bypass vein graft
Right
coronary
artery
Left
coronary
arteries
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However, there are a subset of patients that are unable to receive these surgical procedures due to various comorbidities [10,42], As a result, there has been a push towards alternative treatments that will revascularize the damaged heart tissue [3,10,43],
Study Objective
The overall clinical motivation for this research pertains to the treatment of CAD, after a MI, through the route of therapeutic angiogenesis by long-term and sequential delivery of multiple GFs. To achieve this type of GF delivery, a biocompatible polymer gel with encapsulated nanoparticles was designed. The RTG is sulfonated to protect VEGF from degradation and to slow its release over time. The nanoparticles will be loaded with PDGF and embedded within the SRTG, which will slowly release the particles over an extended period of time as the gel degrades. As this treatment is intended to help revascularize the heart after a MI, an anti-inflammatory cytokine, IL-10, will be incorporated into the SRTG as well. IL-10 will be mixed with the polymer system last, and should release first from the system to help reduce inflammation and reduce the injury effect of cardiac remodeling.
13


CHAPTER II
BACKGROUND
Therapeutic Angiogenesis Using Growth Factors
A potential alternative treatment to revascularizing ischemic tissue is therapeutic angiogenesis through GF delivery. Angiogenesis is the growth of the nascent blood vessels from existing vasculature networks. The main mechanism behind the sprouting of new capillaries works through the release of GFs, such as VEGF and FGF-2, which signals for the migration and proliferation of endothelial cells to differentiate into guiding tip cells or proliferating stalk cells [17], Both of these cell types work together to form directional, elongating, new vessels. Later, a maturation process occurs involving pericytes which are stimulated by other GFs to cover the endothelial cells [44], The synergy of these cells through this process enables fully functional vessels to deliver blood flow to new or ischemic areas. Therapeutic angiogenesis aims to induce and control this angiogenic response in order to revascularize ischemic tissues.
Therapeutic Angiogenesis
Therapeutic angiogenesis has been extensively studied for the treatment of many human diseases. The generation of functional blood vessels from single GF administration was pioneered by JM Inser when he injected VEGF 165 in a rabbit hindlimb ischemia model [45], Since then, several other angiogenic GFs have been used separately or together to promote angiogenesis. These GFs are not only utilized for diseases, but for other processes as well, such as wound healing, and organ repair and regeneration. As a result of dedicated research over the last two decades, many different types and combinations of GFs have been elucidated for the purpose of inducing angiogenesis for MI treatment.
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Growth Factor Considerations
Many different GFs have been extensively studied and applied for therapeutic angiogenesis purposes. These angiogenic GFs consist of VEGF, bFGF, hepatocyte growth factor, PDGF, angiopoietin-1, transforming growth factor-P and insulin-like growth factor. VEGF has shown to be the most significant controller of physiological angiogenesis during growth, healing and in response to hypoxia [13,46], VEGF provides angiogenic effects by binding to specific VEGF receptors on ECs, which leads to receptor dimerization and downstream signal transduction [12], When VEGF binding to ECs occurs, it induces their migration, proliferation, and formation of nascent or large vessels. Hypoxia, oncogenes, tumor suppressor genes, inflammatory cytokines, and other GFs can all activate the production of VEGF and the ensuing angiogenesis response [47], However, when presented alone, VEGF can cause leaky and immature vessel formation with poor function and these vessels may regress quickly [17,48,49], Therefore, a second GF should be delivered to help stabilize and mature the neovessels.
PDGF has been shown to promote vascular maturation by recruiting SMCs and pericytes to newly formed vessels [12], This GF is produced by various cell types in response to external cues such as hypoxia or GF stimulation [50], PDGF also induces expression of VEGF which could provide a positive feedback loop to produce an increased angiogenesis response [51], Moreover, the PDGF signal regulates over 80 genes that are involved with matrix and cytoskeleton proteins, GFs, growth inhibitors, transcription factors involved in cell cycle regulation, and others [9], In addition to stimulating pericytes to help mature newly formed vessels, it has also been shown that PDGF promotes vessel stabilization through stimulation of proliferation and migration of vascular ECs [52], Sequentially delivering
15


VEGF and PDGF should provide an external stimulus to promote stable and mature vascularization for therapeutic angiogenesis applications.
Anti-inflammatory Factor Considerations
To limit the damaging effects of the immune response, delivery of anti-inflammatory cytokines has been shown to optimize cardiac repair following an MF IL-10 is a potent antiinflammatory cytokine which acts on monocytes by deactivating them and suppressing many pro-inflammatory mediators [33], This cytokine has powerful inhibitory effects on monocytic cell lines and macrophages by inhibiting a panel of pro-inflammatory cytokine mRNA expressions in both cell lines [53], It has been shown that IL-10 treatment in an acute MI mouse model significantly improved LV functions, reduced infarct size and attenuated infarct wall thinning [33], Furthermore, a study performed using a rat post-MI heat failure model showed that therapy with IL-10 significantly improved LV function post-MI and resulted in a reduced myocardial infiltration of macrophages [54], Limiting the detrimental effects of pro-inflammatory cytokines is critical for improving heart function after MI.
Current Biological Factor Delivery Methods
Therapeutic angiogenesis involves the delivery of GFs, through proteins or genes encoding target proteins, or using stem cells to initiate the response [7], Because some approaches to therapeutic angiogenesis involve the use of viral vectors or cells, GF, protein delivery is the most simple and direct way [8], VEGF and bFGF are well-studied GFs and have been tested in human clinical trials. The clinical studies used direct injection of free GFs but failed to demonstrate efficacy; this is likely due to the GFs being limited by their rapid diffusion rate, poor biostability, and short half-lives in vivo [15], Therefore, controlled delivery systems for GFs is highly desirable and currently under extensive research [3,7,8],
16


Conversely, there remains significant barriers to GF delivery systems related to loading capacity and long-term efficacy [5],
Polymeric Biomaterials for Protein Delivery Polymeric biomaterials may be a solution to providing sustained, local delivery of GFs for therapeutic angiogenesis [5], Certain biomaterials can overcome the short half-lives and rapid diffusion of free GFs by releasing the proteins over the course of several weeks while stabilizing them in the polymeric network [4], Initial systems for therapeutic angiogenesis demonstrated the use of poly(lactic-co-glycolic acid) (PLGA) as a base polymer for controlled release of VEGF [55], Due to its biodegradability and ability to control GF release, PLGA remains a favorable polymer even though it can cause undesirable inflammation from the increased acidic environment created after the polymer is degraded [8], Poly(ethylene glycol) (PEG) is a biocompatible polymer found in nature which has been utilized to create cell-responsive hydrogels for VEGF release [56], Heparin, a naturally sulfated polysaccharide, has been shown to have binding affinity to many different proteins such as GFs, thus heparin plays an essential role in regulating various biological signaling processes [57], However, naturally derived materials generally experience batch-to-batch variance and have less tunable chemical properties in comparison to synthetic polymers [3], In order to achieve effective therapeutic angiogenesis, the proper combination of biocompatibility and GF encapsulation properties need to be incorporated into the polymer backbone to provide a sustained release of these proteins.
17


Sustained Growth Factor Release
The long-term release of biological signals from a polymeric delivery system involves many complex physical and chemical interactions which need to be optimized to develop a consistent drug delivery vehicle. Most delivery systems have an initial burst-release where the majority of loaded GFs are released in the first hours after injection [3], Although not fully understood, the mechanism behind the burst-release process relates to processing conditions, surface characteristics, and geometry of the biomaterials used [58], Different strategies have been implemented to prevent this burst release by modulating the physical, biochemical affinity, and covalent binding interactions between the loaded GFs and polymeric material. Physical entrapment of the protein in polymeric scaffolds allows for controlled release of the biological signal based on the pore size of the polymer [5], In comparison to free VEGF injections in a myocardial ischemia or hindlimb ischemia model, controlled release of VEGF from PLGA scaffolds significantly improved angiogenesis [3], This improved angiogenesis, with controlled VEGF release, has also been demonstrated using PLGA microspheres containing alginate hydrogel [59], While physical entrapment of the GF in the polymer displays controlled release properties and improves angiogenesis, biochemical affinity interactions with heparin can be applied to further increase new blood vessel formation.
Modulating Physical and Chemical Interactions
Hydrogen and ionic bonding interactions between GFs and heparin can be utilized to slow down protein release [59], bFGF preferentially binds to heparin sulfate and loading bFGF into PLGA nanospheres conjugated with heparin led to a linear release of the GF over
18


a 15 day period [21], The positive charge on GF residues allow for preferential binding to the sulfonate groups on heparin which are inherently negative (Figure 7).
o
R------S------Cf
O
Figure 7 Sulfonate group exhibiting a negative charge. R denotes an attached carbon bound to sulfur, as well as the remaining chemical makeup of the polymer.
Further prolonging of this release can be achieved by encapsulating these PLGA nanospheres into fibrin hydrogels [12], Once encapsulated in the hydrogel, sustained release of bFGF led to significant increases in capillary density and cell proliferation in a murine hindlimb ischemia model [12], Studies have also shown that the release of bFGF can be adjusted by varying the heparin concentration in the polymer [60], Similarly to bFGF, VEGF has positively charged domains of basic amino acid residues that bind to the sulfonate groups on heparin [61,62], Furthermore, some proteins can be covalently linked to polymeric scaffolds through a hydrolytically degradable or MMP-degradable linker. It has been shown that the formation of a more controlled and stable vasculature can be obtained using scaffolds that incorporate covalently attached VEGF [56], On the other hand, the chemical modifications utilized for covalent linkages can affect the biological activity of the attached GF [63], Developing a temperature-responsive hydrogel, incorporated with heparin and encapsulating nanoparticles containing GFs, should provide controlled and sustained protein release while improving angiogenesis.
Thermally Induced Gelling Systems
The clinical feasibility of therapeutic angiogenesis methods can be improved by using polymeric materials that offer spatial and temporal control of GF release [64], Temperature-
19


responsive hydrogels are one type of biomaterial that can provide this controlled release process. The incorporation of poly(N-isopropylacrylamide) (PNIPAm) into hydrogels has been heavily researched due to the polymers reverse thermal gelling properties at different temperatures [65-68], This polymer exhibits a transition from a soluble liquid solution to a solid gel once the lower critical solution temperature (LCST) threshold has been crossed (Figure 8) [65],
Hydrophilic Hydrophobic
hydrated polymer chains collapse of polymer
Figure 8 Polymer solubility behavior at the LCST. Left hand side shows hydrated polymer below LCST with entropic loss of water and chain collapse above LCST (right hand side) [69],
PNIPAm-based hydrogels have a sharp phase transition where the polymers remain a liquid at room temperature and rapidly become a solid gel when injected at body temperature (37 Q, trapping any nanoparticles or proteins in the gel during this transition [64,70], In comparison with other cationic polymers, PNIPAm moieties for targeted delivery offers the advantages of good sensitivity, reversible transitions, and low cytotoxicity [30], Additionally, protein release is easily tunable by manipulating the degradation properties of these polymers. By adjusting the hydrophilicity interactions of the polymer, and through altering the chemical structure of PNIPAm, GF release can be varied through a change in the biomaterials degradation rate [70], In order to implement a further controlled and sequential
20


release of factors, nanoparticles containing GFs can be entrapped in the hydrogel to provide additional angiogenic proteins after the gel has degraded.
Nanoparticles for Sequential Growth Factor Release
Several studies have demonstrated the use of combining nanoparticle/hydrogel delivery systems for local and sustained release of GFs for therapeutic angiogenesis [3,17,59], The study performed by Awada and colleagues (2015) demonstrated the efficacy and importance of this dual delivery system to provide sequential release of GFs to improve revascularization and heart function after MF As discussed earlier, it is essential that the delivery system administers the GFs sequentially, while in their bioactive conformations to mimic their physiological mechanism during angiogenesis [3,17,18], If delivered simultaneously, early-stage angiogenic factors can interfere with late-stage factors which can cause immature blood vessel formation [18], These studies illustrate the importance of releasing early-stage GFs from the hydrogel first, and then using another delivery vehicle (PLGA nanoparticles for example) to sequentially supply late-stage GFs several days after the release of the first factor. The optimal delivery system to improve revascularization and heart function after a MI will need to provide sequential, sustained, and local release of GFs to the damaged heart tissue.
Nanoparticle Considerations
Many different compositions of nanoparticle (NP) delivery systems for GFs have been designed and fabricated from diverse types of synthetic and natural polymers. GFs can be loaded in the system before or after the fabrication of the particles by covalent or non-covalent methods. As for non-covalent means, GFs can be incorporated by adsorption, electrostatic interactions, or complexation [71], NPs are distinguished from other particulate
21


systems by their smaller size, with the range of 1-1000 nm compared to > 1 pm for microparticles [72], NPs offer advantages over microparticles because they can penetrate deeper into tissues through small capillaries and epithelial lining, allowing more efficient delivery of GFs to the target sites [73], When designing NPs for GF delivery, several issues must be considered: environmental factors that may denature or deactivate the NPs, physiochemical properties such as size distribution and surface charge, targeting ability of NPs for site-specific delivery of GFs, and controllable GF release profiles to meet temporal and spatial demands [71], Successfully meeting these specific combinations of factors while designing NPs are likely to lead to more desirable therapeutic outcomes.
There are several different types of NP systems that have been researched and elucidated for GF delivery such as lipid, polymer-based, and other miscellaneous systems (Figure 9). Polymer-based NPs in the form of nanospheres, nanocapsules, micelles, and dendritic particles have all been broadly examined for GF delivery. Polymeric micelles are created from amphiphilic polymers that typically self-assemble in an aqueous environment to form colloidal NPs [74], These micelles have a unique core-shell structure which consists of an inner hydrophobic core and a hydrophilic exterior shell. Polymeric micelles offer advantages of smaller size and a uniform size distribution, in addition to high extravasating and tissue penetrating ability with reduced cytotoxicity [71], Studies have demonstrated high GF loading efficiency and sustained release of bFGF over a 2-month period without an initial burst-release from polymeric micelles [75,76], Polymeric micelles present significant advantages over other NP delivery systems which makes them suitable for in vivo GF delivery to promote therapeutic angiogenesis.
22


Polymeric nanocapsules Polymeric atmospheres
Growth factor Polymers
nine block: hydrophilic md Yellow block: hydrophobic end
Blue circle: hydrophilic end Black line: lipophilic end O Hydrophobic core group
Hydrophilic surface groups
Figure 9 Schematic representation of different NP systems [71],
23


CHAPTER III
PREVIOUS WORK
Sulfonated-PSHU-PNIPAm (sulfonated-PSHU-PNIPAm will be used interchangeably with SRTG) has previously been designed and studied as a potential delivery vehicle for positively charged proteins. A synthesis method was established and the polymer has been characterized. Additionally, in vitro cytotoxicity testing, BSA release tests, and CD spectroscopy were completed for the SRTG [22],
SRTG LCST
LCST was used to determine the sol-gel phase transition temperature of sulfonated-PSHU-PNIPAm (Figure 10).
u
e
m
c
re
Temperature (C)
Figure 10 Temperature-dependent phase transition of sulfonated PSHU-PNIPAm. A: The sulfonated PSHU-PNIPAm undergoes a sharp, reversible phase transition around 32 Cas determined by UV-Visible spectroscopy. B: An aqueous solution of sulfonated PSHU-PNIPAm at room temperature (C) turns to physical gel at 37 C [22],
24


An aqueous solution of sulfonated-PSHU-PNIPAm was heated slowly from 20 Cto 44 C and percent transmittance was recorded during the process. The transmittance of the solution decreased slowly upon initial heating from 20 Cto 31 C and then rapidly approached zero transmittance at 32 C. Upon further heating to over 33 C, the solution turned to an opaque solid indicating that the aqueous solution turns into a physical gel as the temperature increases. Thermodynamic competition between hydration of the polymer chains and the hydrophobic interactions between the polymer molecules causes this phase transition to occur [77], At low temperatures, or below the LCST which is 32 Cin this case, hydration of the polymer is thermodynamically favorable and the polymer molecules are maintained in a solution state (Figure 10B). When the temperature is above the LCST, hydrophobic interactions between the polymer and chains are favored and the polymer molecules interact to form a stand-alone physical gel (Figure 10) [22],
Cytotoxicity
An MTT assay with C2C12 myoblast cells was implemented to investigate in vitro cytotoxicity effects (Figure 11). This method is well documented for measuring cell viability and for supplying a general indication of cell health [55-57], Figure 11 shows no statistical difference between the absorbance of pure medium and the polymer extract, while the addition of 10% DMSO significantly reduced the absorbance level. In this experiment, pure medium was used as a positive control and the DMSO was used as a negative control. The differences in absorbance are directly related to differences in metabolic activity of the cells, which indicates the change in viable cells. Thus, the similar absorbance levels of the polymer extract and the pure medium samples is evidence that the sulfonated-PSHU-PNIPAm is non-
25


cytotoxic [22], Plain PSHU-PNIPAm was not included in this study as a previous study has proven that this polymer has good biocompatibility [78],
Extract Medium 10% DMSO
Figure 11 In vitro cytotoxicity of sulfonated PSHU-PNIPAm by MTT assay. Results demonstrated no cytotoxic effects of sulfonated PSHU-PNIPAm on C2C12 cells after exposure to the polymer extract in medium, while 10% DMSO shows significant cytotoxicity. There is no statistical difference between medium and extracts. Data represent mean SD. indicates p-value < 0.05 (Students t-test) [22],
BSA Release Test
A BSA release test was implemented to examine the protein release rate properties of PSHU-PNIPAm and sulfonated-PSHU-PNIPAm at varying wt. % concentrations. As this polymer system is intended to be a protein delivery vehicle for positively charged molecules, BSA was used for this study due to its inherent positive charge [22], The sulfonated-PSHU-PNIPAm showed an increased sustained release profile in comparison to plain PSHU-PNIPAm (Figure 12).
26


100
Release time (day)
Figure 12 BSA release profiles from sulfonated PSHU-PNIPAm and plain PSHU-PNIPAm: sulfonated PSHU-PNIPAm 10%, sulfonated PSHU-PNIPAm 15%, plain PSHU-PNIPAm 10%, plain PSHU-PNIPAm 15%. BSA release was more sustained from sulfonated PSHU-PNIPAm than that from plain PSHU-PNIPAm, with a more sustained profile at higher concentrations. Data represent mean SD [22],
This showed that the negatively charged sulfonated groups on the sulfonated-PSHU-PNIPAm may effectively hold BSA in the polymer matrix. Furthermore, the BSA release profile showed more sustained release at higher polymer concentrations which supports that the release rate can be readily modified by changing the sulfonated PSHU-PNIPAm concentration [22], A possible explanation is that higher concentrations of the sulfonated PSHU-PNIPAm leads to an increased sulfonate group density in a designated volume of the polymer gel, which results in more negative overall charge. This causes an enhanced electrostatic interaction between sulfonated PSHU-PNIPAm and BSA.
27


CD Spectroscopy
If the sulfonated PSHU-PNIPAm will serve as a protein delivery vehicle, then it should not affect protein structure nor lead to the denaturation of the proteins. To demonstrate this, the secondary structure of the released BSA from the final day of the release study was analyzed by CD spectroscopy and compared with natural BSA solution. No significant difference was observed in all samples (Figure 13) with a typical a-helix conformation [79-81], which confirms that the protein structure is well-preserved [22],
Wavelength (nm)
Figure 13 CD spectra of native and released BSA. The similarity of the protein conformation between samples indicates that the secondary structure of the BSA protein was well preserved by the system [22],
28


CHAPTER IV
HYPOTHESIS AND SPECIFIC AIMS Hypothesis
Based on preliminary data, it was hypothesized that the sulfonated thermal gel encapsulating micelle nanoparticles, while sequentially releasing biological factors, will induce vascularization and reduce an inflammatory response.
Specific Aims
The first specific aim was to synthesize and characterize PEG-PSHU-PEG micelle NPs. As the micelles will be utilized to provide sequential delivery of PDGF from the polymer scaffold, it is important to confirm the overall molecular structure of the micelles and that they are uniformly sized on the nanometer scale. The second specific aim was to ensure that the biological factors were being delivered sequentially and for an extended amount of time from the delivery system in vitro. To assess this, a release study was performed with the polymer delivery system and samples were taken daily. The amount of protein released from the samples was quantified using specific ELISAs for each factor. The third specific aim was to demonstrate a substantial angiogenic response with a reduced immune response in vivo. This was evaluated by injecting the polymer system into the subcutaneous region of the lower backs of mice. Angiogenesis and immune response were observed using immunohistochemistry (IHC) with stains specific for: endothelial cells as a marker for early blood vessel development, von Willebrand factor (vWF) to show functional vessels, a-smooth muscle actin (a-SMA) to identity the recruitment of smooth muscle cells which demonstrates mature blood vessel formation and macrophages to determine the reduction in the inflammatory response.
29


CHAPTER V
MATERIALS AND METHODS Materials
N-isopropylacrylamide (NIPAm) was purchased from Tokyo Chemical Industry (Chuo-ku, Tokyo, Japan). Anhydrous N,N-dimethylformamide (DMF) was purchased from EMD Millipore (Billerica, MA, USA). 4,4'-azobis(4-cyanovaleric acid) (ACA), N-hydroxysuccinimide (NHS), sodium bicarbonate, 1-bromohexane, TWEEN 20, Triton X-100, y-globulins from bovine blood, urea, hexamethylene diisocyante (HDI), acetic acid, sodium acetate, 1,3-propane sultone (PS), trifluoroacetic acid (TFA), triethylamine, dimethyl sulfoxide (DMSO), 2,2'-Azino-bis(3-ethylbenzothiazoline-6-sulfonic acid) (ABTS) Liquid Substrate System, dichloromethane (DCM) and bovine serum albumin (BSA) were purchased from Sigma Aldrich (St. Louis, MO, USA). Di-tert-butyl dicarbonate, ethyl acetate, N-hydroxysuccinimide (NHS), N-(3-dimethylamino-propyl)-N-ethylcarbodiimide hydrochloride (EDC), 2-Amino-1,2-propanediol, 98% (Serinol), anhydrous methanol, dimethyl sulfoxide-d6 (DMSO-d6), polyethylene glycol 1000, ethanol, hexane and anhydrous diethyl ether were purchased from Fisher Scientific (Pittsburgh, PA, USA). Goat anti-Rabbit IgG (H+L) Secondary Antibody Alexa Fluor 594, Goat anti-Rat IgG (H+L) Secondary Antibody Alexa Fluor 488, Rabbit anti-Goat IgG (H+L) Secondary Antibody Alexa Fluor 594, CD31 Antibody (Rat IgG2a) and Smooth Muscle Actin Antibody (a-SMA, rabbit IgG) were purchased form Thermo Fisher Scientific (Waltham, MA, USA). Anti-Von Willebrand Factor antibody (vWF, sheep IgG) and Anti-CD68 antibody (Rabbit IgG) were purchased from Abeam (Cambridge, Ma). Human PDGF-BB Standard ABTS ELISA Development Kit, Murine IL-10 Standard ABTS ELISA Development Kit, Murine VEGF Standard ABTS ELISA Development Kit, Recombinant Human PDGF-BB, Recombinant Murine IL-10 and
30


Recombinant Murine VEGF165 were purchased from Peprotech (Rocky Hill, NJ, USA). Saline, and isoflurane were purchased from MWI Veterinary Supply (Boise, ID, USA). 10 % formalin was purchased from JT Baker (Phillipsburg, NJ, USA). Sucrose (RNASE &
DNASE free) was purchased from VWR Life Science (Radnor, PA, USA). Optimal cutting temperature (OCT) compound was purchased from Sakura (Torrance, CA, USA). Phosphate buffered saline (PBS) was purchased from HyClone Laboratories, Inc. (South Logan, Utah, USA). Alexa Fluor 594 (goat anti-rabbit IgG) was purchased from Life Technologies (Carlsbad, CA, USA). Dapi flouromount-G was purchased from Electron Microscope Sciences (Hartfield, PA, USA). Spectra/Por dialysis membranes (MWCO: 3500-5000 and 12,000-14,000 Da) were purchased from Spectrum Laboratories (Rancho Dominguez, CA).
Equipment
Proton nuclear magnetic resonance (*H NMR) was performed on a Varian Inova 500 NMR Spectrometer and samples were run in DMSO-d6 at room temperature. Fourier transform infrared spectroscopy (FTIR) was performed on aNicolet 6700 FTIR Spectrometer and samples were run on polyethylene infrared (IR) sample cards. Polymer morphology was imaged using a JEOL (Peabody, MA) JSAM-60101a analytical scanning electron microscope. The low critical solution temperature (LCST) was determined using a Cary 100 Bio UV-Visible spectrophotometer with a temperature-controlled cell holder. Nanoparticle size was measured using a Zetasizer Nano ZS (Malvern Instruments Ltd, Worcestershire, UK). Elemental analysis was performed by MicroAnalysis, Inc.(Wilmington, DE). ELISA color development was monitored with an ELISA plate reader (BioTek Synergy 2 Multi-Mode Reader) at 405 nm with wavelength correction set at 650 nm. Tissue was sectioned using a
31


CryoStar NX70 Cryostat. Confocal images were taken using a Zeiss LSM 780 spectral microscope. ImageJ was used to quantify variables for image analysis.
Polymer Synthesis
N-BOC serinol synthesis
N-Boc serinol was synthesized through Boc group protection of serinol. Serinol (1.959 g) was dissolved in 20mL of absolute ethanol and stirred at 4 C. Di-tert-butyl dicarbonate (5.973 mL) was dissolved in 20 mL of absolute ethanol and added dropwise to the serinol solution over a period of one hour, while maintaining 4 C and constant stirring. The solution was warmed to 37 C with vigorous stirring and reacted for one hour. The ethanol was removed by rotary evaporation at 45 C and 10 mbar vacuum and the solid was re-dissolved in a 1:1 mixture of hexane and ethyl acetate by gentle heating. Additional hexane was added until precipitation was observed and the resulting suspension was stored at 4 C overnight to allow recrystallization. Subsequent vacuum filtration yielded a white flaky product.
Polvfserinol hexamethylene urea) or PSHU backbone synthesis
PSHU is the backbone of the polymer system used in this study. N-BOC-serinol (1.149 g, 6 mmol) and urea (0.36 g, 6 mmol) were lyophilized for 24 hours and then dissolved in 6 mL of anhydrous DMF in a 25 mL round bottom flask under a nitrogen atmosphere. HDI (1.928 mL, 12 mmol) was added dropwise to the flask and the polymerization was performed for seven days at 90 C under a nitrogen atmosphere. After cooling down to ambient temperature, the mixture was precipitated twice into excess cool anhydrous diethyl ether and twice in distilled water for purification. The product was subsequently lyophilized at -45 C for 24 h and recovered for further conjugation.
32


PSHU deprotection
In order to conjugate other polymers/chemicals to PSHU, it was deprotected by removing the tert-Butyloxycarbonyl (BOC) groups and exposing the primary amine groups through the following process. PSHU (1.00 g) was dissolved in 10 mL of a TFA/DCM (1:1, v/v) in a round bottom flask. The solution was left gently stirring with no heat at room temperature for 30 minutes. Rotary evaporation was used to remove solvents at 45 C and 10 mbar. The remaining contents were re-dissolved in DMF (1 mL), and diethyl ether was added to the flask to precipitate the polymer out of solution. Excess ether was poured off while taking care to keep polymer within the flask. Rotary evaporation was then used to remove residual ether and the deprotected PSHU (dPSHU) was purified by two more precipitations into diethyl ether. Finally, the polymer was decanted in water and lyophilized at -45 C for 24 h. At the end of this process, the polymer should look dry and white.
PNIPAm Synthesis
PNIPAm was conjugated using radical polymerization with an azobis initiator. NIP Am (5.0 g, 44.2 mmol) and AC A (0.062 g) were dissolved in anhydrous methanol (25 mL), and the mixture was purged with nitrogen gas for 30 minutes at room temperature. A reflux condenser apparatus was set up, and the solution was stirred for three hours at 68 C under a nitrogen atmosphere. Next, the solution was precipitated into milliQ water at 60 C in a dropwise manner. Following precipitation, the warm water was discarded and 40 mL of cold Milli-Q water was added to the precipitate. The precipitate was allowed to dissolve into the water in a cold room at 4 C. The resulting solution was dialyzed using 3500 kDa MWCO dialysis tubing against 1 L of milliQ water for 48 h. The product was lyophilized at -45 C for 48 h.
33


PSHU-PNIPAm Conjugation fRTG Synthesis')
PNIPAm was conjugated to 25% of the free amine groups on dPSHU. EDC/NHS chemistry was used to achieve this conjugation. PNIPAm (0.75 g) was dissolved in 5 mL of anhydrous DMF with three molar excess of EDC and NHS in a 25 mL round bottom flask at room temperature under a nitrogen atmosphere, and the mixture was stirred for 24 h. One milliliter of dPSHU solution (0.75 g/mL) prepared in anhydrous DMF was added slowly to the PNIPAm solution flask and the reaction was performed for 24 h at room temperature under a nitrogen atmosphere. The resulting solution was capped (no vent) and left stirring for 24 hours to complete the reaction. The polymer was precipitated three times in diethyl ether, and the solvent was removed each time by rotary evaporation. The dried polymer was dissolved in milliQ water at 4C and dialyzed (MWCO: 12-14 KDa) against 1 L milliQ water for 48 h at room temp. The product was lyophilized at -45 C for 48 hours.
Sulfonation of PSHU-PNIPAm (SRTG Synthesis)
As PNIPAm, was only conjugated to 25% of the available primary amine groups, the remaining amine groups were left for the attachment of sulfonate groups. For this sulfonation reaction, PS (0.034 g, 5 mmol) and t-BuOK (0.032 g, 5 mmol) were dissolved in 3 mL of anhydrous DMF in a 25 mL round bottom flask at 50 C under a nitrogen atmosphere. A solution of 3 mL PSHU-PNIPAm (0.1 g/mL) in anhydrous DMF was added slowly to the flask and the sulfonation reaction was performed for 3 days at 60 C under a nitrogen atmosphere. After cooling down to ambient temperature, the mixture was precipitated into excess diethyl ether three times. Finally, the polymer was dissolved in milliQ water and dialyzed (MWCO: 12-14 KDa) against 1 L of milliQ water for 48 h at room temperature and lyophilized at -45 C for 48 h.
34


PEG-PSHU-PEG Micelle Nanoparticle Polymer Synthesis
N-Boc-serinol (0.2873 g, E5 mmol) and urea (0.09 g, 1.5 mmol) were weighed out and lyophilized at -45 C for 48 h. The reactants were dissolved in 1.5 mL of anhydrous DMF in a 25 mL round bottom flask at 90C under gentle stirring and a nitrogen atmosphere. Hexamethylene diisocyanate (HDI) (0.482 mL, 3 mmol) was added drop-wise, and the polymerization was carried out for 5 days. After the specified time, an excess of polyethylene glycol (PEG 1000, 4 mmol) was dehydrated and added to the reaction. The PEGylation reaction was carried out for 24 h at 90C. The resulting product, poly(ethylene glycol)-block-poly(serinol hexamethylene urea)-block-poly(ethylene glycol) (PEG-PSHU-PEG), was purified by three precipitations in diethyl ether and then dried completely by extended rotary evaporation at 50C and 10 mbar vacuum. Subsequently, the polymer was dissolved in milliQ water and dialyzed (MWCO: 3.5 kDa) against 1 L of milliQ water for 72 h at room temperature. Then the product was lyophilized -45 C for 48 h to yield a white flaky material. Micelle Fabrication
GF-loaded micelles were fabricated by a traditional emulsification-sonication procedure. The PEG-PSHU-PEG polymer and GF were dissolved in 1 mL DMSO at 1 wt% (polymer/DMSO). This solution was then added drop-wise to a beaker containing 20 mL of milliQ water partially submerged in an ultrasonic bath. The resulting emulsion was sonicated for 10 min. Removal of DMSO was carried out by centrifugation at 11,000 revolutions per minute (rpm) for 5 min, pouring off the supernatant and then re-suspending the micelles in milliQ water. This DMSO extraction procedure was carried out 3 times. The resulting micelles were either used immediately or stored at -20C for later use.
35


Polymer Characterization
Elemental Analysis
To determine the amount of sulfonation on sulfonated-PSHU-PNIPAm, the polymer was sent to Micro-Analysis, Inc. (Wilmington, DE) for elemental analysis testing. Carbon, hydrogen and nitrogen (CHN) analysis was used to determine the percent, by weight, of carbon, hydrogen, and nitrogen contained in the polymer. The company has a separate testing method for sulfur content using an Antek analysis machine. The Antek provides the determination of total nitrogen and total sulfur in organic materials, both solids and liquids at trace levels. Sample material is combusted in combination with oxygen, converting chemically bound nitrogen to NO and sulfur to S02. The combustion gases are routed through a membrane drying system to remove all water and then to the detector modules for quantification. Nitrogen is detected by way of chemiluminescence and sulfur by ultraviolet fluorescence.
Dynamic Light Scattering (DLS)
DLS is a technique that can be used to determine the size distribution profile of different particles, or polymers, in solution. The Zetasizer Nano ZS uses this technique to measure the diffusion of particles moving under Brownian motion, and converts this to size and a size distribution using the Stoke-Einstein relationship. Non-Invasive Back Scatter technology (NIBS) is incorporated into this instrument to give the highest sensitivity simultaneously with the highest size and concentration range.
Scanning Electron Microscopy (SEMI
SEM was used to examine polymer structure and morphology on the nanometer scale. This technique was used show the morphology of the NPs alone and encapsulated within the
36


SRTG. For particle morphology assessment, dry micelles were sputter coated with 5 nm Au and examined by SEM.
Proton Nuclear Magnetic Resonance (1H NMR)
Proton nuclear magnetic resonance (*H NMR) was completed with a Varian Inova 500 NMR Spectrometer. *H NMR was used to confirm the structure of PSHU, the structure of PEG-PSHU-PEG, the removal of BOC protecting groups on dPSHU, and the successful conjugation of PNIPAm to dPSHU.
The PSHU, dPSHU, PSHU-PNIPAm and PEG-PSHU-PEG samples (3-5 mg) to be analyzed, were dissolved in 600 pL of DMSO-d6. Spectra were processed and analyzed using ACD ID NMR Processor software (Advanced Chemistry Development, Inc.).
Fourier Transform Infrared tFTIR) Spectroscopy
Sulfonated-PSHU-PNIPAm and PSHU-PNIPAm samples were evaluated using FTIR spectroscopy. Samples were dissolved in tetrahydrofuran (THF) and placed on polyethylene windowed cards which were then analyzed by aNicolet 6700 (Thermo Fisher Scientific).
In Vitro Biological Factor Release Study
5 wt% polymeric (RTG + VEGF + NPs + PDGF + IL-10 and SRTG + VEGF + NPs + PDGF + IL-10) solutions were created using PBS with 0.2% BSA. Solutions were mixed and left to dissolve overnight at 4 C. 500 ng of each factor was added to the gels at different times. VEGF was added directly to the polymer solutions 2 hours before the gelling was to occur. The micelle NPs loaded with PDGF were added 30 minutes before gelling, and IL-10 was added 10 minutes before gelation. 50 pL gels were formed in 2 mL vials (Figure 14) in a 37 C incubator for 2 minutes and then 1 mL of warm (37 C) release solution (lx PBS with 0.2% BSA) was added to the vials. After 2 minutes, 1 mL of release solution was removed from each sample for the day 0-timepoint, and 1 mL of fresh warm release solution was
37


added. Samples were then taken every 24 hours and immediately placed in a -80 C freezer for further analysis.

Figure 14 Formed gels in 2 mL vials for in vitro release study.
Enzyme-Linked Immunosorbent Assay (ELISA)
ELISA is a laboratory technique for quantifying the concentration of specific antigens or antibodies in various solutions. A sandwich ELISA format was used to quantify the concentration of factors in the release solution samples from the in vitro polymer delivery systems. The protocol from the manufacture was followed for each ELISA kit. First, the primary antibodies are attached to the wells. Once the antibodies have attached overnight, the samples are added to the wells and the protein of interest attaches to the primary antibody. A secondary antibody is then added, which will attach to the protein of interest. Next, an enzyme that attaches to the secondary antibody is added. Subsequently, a liquid substrate is added that interacts with the enzyme which produces a color change. The color development is detected using a spectrophotometer set to 405 nm with wavelength correction set at 650 nm. The plate is monitored at 5-minute intervals for approximately 20 minutes. These measurements correspond to the concentration of the protein of interest in the initial sample
38


solution. To quantify the amount of protein from the absorbance readings for each well, a standard curve is created with each ELISA plate to compare values to the unknown concentrations. Additionally, the plate is washed several times between each step of the process to remove any residual materials not attached to the bottom of the wells.
In Vivo Mouse Subcutaneous Biological Factor Injections
The immune response and angiogenesis analysis procedure was approved by the Institutional Animal Care and Use Committee (IACUC). A total of 36 adult male C57BL/6 mice were used for the study, 3 mice per each injection group (saline, SRTG, SRTG +
VEGF, SRTG + VEGF + PDGF, SRTG + VEGF + NPs + PDGF, or SRTG + VEGF + NPs + PDGF + IL-10) for 2 time points (7, 21 days). The mice were allowed 7 d to acclimate prior to injections and maintained on a 14/10-hour light/dark cycle with access to food and water ad libitum.
The mice were anaesthetized using continuous isoflurane and oxygen inhalation. Initial induction was at 5 % isoflurane in oxygen and then maintained at 2 % isoflurane in oxygen. Under anesthesia, 60 pLof saline, SRTG, SRTG + VEGF, SRTG + VEGF + PDGF, SRTG + VEGF + NPs + PDGF, or SRTG + VEGF + NPs + PDGF + IL-10 was subcutaneously injected in the middle back of the mice through a 25G insulin needle (Figure 15). The amount of protein injected for each factor was 250 ng. The injection bolus was clearly noticeable under the subcutaneous tissue.
39


Figure 15 Demonstration of subcutaneous injection in the middle back of a mouse [82],
Euthanasia and Tissue Harvest
7 or 21 days after injection, the mice were euthanized by carbon dioxide and cervical dislocation. After removing the hair around the injection site with depilatory cream, the subcutaneous tissue (2 cm x 2 cm) was harvested at the site of injection. The site was clearly visible as a result of shaving the backs of the mice at the time of the initial injection.
Histology
The subcutaneous tissue was fixed using 4% Paraformaldehyde in PBS for 24 h, cryoprotected with 30% sucrose in PBS for 24 h, embedded in OCT compound, and frozen at -80 C. The tissue was sectioned transversely with a thickness of 5 pm and placed on glass slides.
The sections were fixed in acetone for 10 min and washed 3 times in Wash Buffer (IX PBS, 0.1% Tween 20) for 5 min each. The sections were permeabilized in Permeabilizing Buffer (IX PBS, 0.5% Triton X-100) for 10 min and washed 3 times in Wash Buffer for 5 min each. The sections were blocked in Blocking/Dilution Buffer (IX PBS, 0.25% Triton X-100, 2% BSA, 4% Bovine Gamma Globulins) for 60 min and washed 3 times in Wash Buffer for 5 min each. All antibodies were diluted in Blocking/Dilution Buffer. The sections were
40


stained with primary antibodies (single or double stains) overnight, washed 3 times in Wash Buffer for 5 min each, stained with the designated secondary antibody for 60 min, washed 3 times in Wash Buffer for 5 min each and washed 3 times in milliQ water for 5 min each. Dapi flouromount-G was used to mount the slides. The sections were double immunostained with either CD31 (1:50) and a-SMA (1:250), or CD31 (1:50) and vWF (1:50). The sections were single immunostained with CD68 (1:500). For the secondary antibodies, CD31 sections were stained with Alexa Fluor 488 (1:500), a-SMA sections were stained with Alexa Flour 594 (1:500), vWF sections were stained with Alexa Flour 594 (1:500) and CD68 sections were stained with Alexa Flour 594 (1:500). Images were taken using confocal microscopy.
Quantitative Analysis of Immunofluorescent Staining 5 low magnification (200x) fields containing the highest number of CD31-, a-SMA-, vWF- or CD68-positive cells were selected for each group following previously published criteria [8], The number of CD31-, a-SMA-, vWF- or CD68-positive cells in the field was counted and confirmed by DAPI-positive nuclei.
Statistical Analysis
All results are reported as means standard error of the mean. Analysis of variance (ANOVA) was utilized to determine significant differences between groups and followed by a two-tailed t-test test when applicable. Statistical significance (*) was indicated when p < 0.05.
41


CHAPTER VI
RESULTS AND DISCUSSION
Synthesis of Sulfonated-PSHU-PNIPAm Reaction Sequence
The synthesized sulfonated-PSHU-PNIPAm was designed to mimic heparin and hold similar biofunction to natural heparin, while having the advantage of being a reverse thermal gel with consistent reproducibility. Moreover, it contains amide and ester groups to provide sites for degradation and the hydrophobic alkyl chains allow for decreased degradation rates and allows for improved cell attachment. The sulfonated-PSHU-PNIPAm was synthesized using PSHU-PNIPAm copolymer (Figure 16).
Hexamethylene diisocyanate
>^o
o
HjN^NH,
Urea
HO, HO
| O .
'n o
H
N-BOC serinol
o;nh
Step 1
90C, 7 days
Step 2 deprotection
------------^
TFA/DCM
H A H
N 6 6 k
S 5
nh2
H jA H
N O O N,
S 5
Step 3
PNIPAAm-COOH
----------^
EDC/NHS
Step 4
1.3 propane sultone t-BuOK
O O NH
NOON
<5 6
V'
O NH
t W
o o o
N^N^NJN' H H H H
'Y'
O^ NH
O
NH

o o o N^iCn^N'
H H H H
O O O
N^N^N^N' H H H H
NH2
N 6" 6 N
<5 6
o o
S
O )
HN
O O O
N^N^N^N" H H H H
Jy
H A H V V o o 0 H A H O O 0
H H H H N 6 0 N 5 6 -^"N "N 'N "N H H H H
X
Figure 16 Synthesis Reaction sequence of sulfonated PSHU-PNIPAm synthesis [22],
The primary amine in PSHU is normally protected by a BOC group but it can easily be deprotected to a free amine group in the mixture of TFA/DCM. Then these deprotected
42


groups can be utilized for further conjugations. PNIPAm was conjugated to 25% of the available primary amine groups and the remaining primary amines were then available for the attachment of sulfonate groups.
Synthesis of PEG-PSHU-PEG Reaction Sequence The novel synthesized PEG-PSHU-PEG was based on the same PSHU backbone as sulfonated-PSHU-PNIPAm, which makes the polymer biocompatible and biodegradable. As this polymer is used to fabricate micelle NPs, it must have differing sections of hydrophobic and hydrophilic chains to form these particles. The PSHU backbone provides the hydrophobic core of the micelles, while the terminal PEG chains provide the hydrophilic interactions on the exterior shell which can ionically bind to GFs (Figure 17). PEG adds an additional benefit as this provides a stealth effect to the NPs for an extended half-life in circulation. When the micelles are injected together with sulfonated-PSHU-PNIPAm, the micelles provide long-term and sequential release of PDGF while the SRTG provides a scaffold for their retention at the injection site.
Hexamethylene Diisocyanate
O
Ji
nh2 nh2
Urea
N-BOC Serinol
A
Step 1
90 C, 5 d
O
X
Cr NH
NH XX X NH
Y
o
o o o
. 1 JL 1
NH NH NH
Poly(serinol hexamethylene urea) (PSHU)
90 C, 1 d H
X -
CT NH
o
O^NH
NH O A O NH
r dr
o o
o o o
111 NH NH NH NH
v
x O
PEG-PSHU-PEG
Figure 17 Reaction sequence of PEG-PSHU-PEG synthesis.
43


PSHU and dPSHU Characterization JH NMR
The synthesis of PSHU from N-BOC serinol, urea, and HDI was confirmed using 'H NMR (Figure 18).
Chemical Shift [ppm]
Figure 18 *H NMR spectrum of PSHU confirming the molecular structure [83],
*H NMR was used to confirm the removal of BOC protecting groups. The resulting free amines were used to conjugate PNIPAm and sulfonate groups. The loss of the b peak confirms the removal of the BOC groups (Figure 19).
1.7 1.6 1.5 1.4 1.3 1.2 1.1 1
Chemical Shift [ppm]
Figure 19 *H NMR spectrum of PSHU and dPSHU confirming the removal of the BOC protecting group with the loss of the b peak [83],
44


PSHU-PNIPAm and PEG-PSHU-PEG Characterization Using JH NMR
The synthesis of PSHU-PNIPAm from N-BOC serinol, urea, HDI, and PNIPAm was confirmed using *H NMR (Figure 20).
Chemical Shift [ppm]
Figure 20 ^NMR spectrum of PSHU-PNIPAm. Successful conjugation of PNIPAm was confirmed by the presence of methylene and methyl protons at 1.55 and 1.09 ppm, respectively.
The synthesis of PEG-PSHU-PEG from N-BOC serinol, urea, HDI, and PEG was confirmed using 'HNMR (Figure 21).
Chemical Shift [ppm]
Figure 21 *H NMR spectra of PEG-PSHU-PEG. Successful conjugation of PEG was confirmed by the presence of the peak at 3.51 ppm which identifies the protons on the PEG repeating unit of the polymer backbone [84],
45


PSHU-PNIPAm and Sulfonated-PSHU-PNIAPAm Characterization Using FTIR
In addition to JH NMR, FTIR was used to further characterize the chemical structure of PSHU-PNIPAm and sulfonated-PSHU-PNIAPAm. This method was used to confirm the conjugation of the sulfonate groups to sulfonated-PSHU-PNIAPAm (Figure 22).
3850.00 3350.00 2850.00 2350.00 1850.00 1350.00 850.00
Wavenumber [1/cm]
Figure 22 FTIR spectra of sulfonated-PSHU-PNIPAm and PSHU-PNIPAm.
FTIR can identify the bond between a single sulfur molecule and oxygen on the sulfonate groups. PSHU-PNIPAm was used as a baseline spectrum to compare with the sulfonated-PSHU-PNIPAm to identify the sulfonate peak. Thus, peaks appearing from the sulfonated polymer spectrum represent changes in bonds due to the sulfonation process. Figure 23 highlights a peak area between 1050 1025 cm"1 that has been identified as a sulfonate peak, which confirms that the SRTG has been sulfonated as this shift is not present in the plain RTG.
46


1200.00 1150.00 1100.00 1050.00 1000.00
Wavenumber [1/cm]
Figure 23 FTIR spectra of sulfonated-PSHU-PNIPAm and PSHU-PNIPAm with an enlargement of the sulfonate peak.
SRTG Sulfonation Detection Using Elemental Analysis
After confirmation of sulfonation with FTIR, elemental analysis was performed to further verify the sulfonation of the SRTG, while also determining the amount of sulfonation. Elemental analysis is a process where a sample of a compound is analyzed to determine its elemental composition for carbon, hydrogen, nitrogen, oxygen and sulfur (Micro-Analysis Inc.). The sample material is combusted and the gases produced are routed through a membrane drying system to remove all water and then the detector modules quantify the chemicals in the compound. Table 1 shows the results of the elemental analysis for the RTG and SRTG. The RTG was used as a baseline for how much sulfur is in the polymer without being sulfonated. The results show that 0.11% wt. of the SRTG contains sulfur groups, which further verifies that the polymer is sulfonated when compared to the plain RTG with 0.02% wt. sulfur content.
47


Table 1 Elemental analysis of PSHU-PNIPAm (RTG) and sulfonated-PSHU-PNIPAm (SRTG). Polymers were analyzed for carbon, hydrogen, nitrogen, oxygen and sulfur content
and were determinec by weight percent for each element.
Polymer C [%] H [%] N [%] 0 [%] S [%]
SRTG 57.09 10.07 11.86 16.65 0.11
RTG 56.44 9.95 12.09 16.15 0.02
PEG-PSHU-PEG Micelle Nanoparticles Size Distribution Using DLS
DLS is a technique that can be used to determine the size distribution profile of different particles, or polymers, in solution. The micelle NPs size was determined using a Zetasizer Nano ZS. Three separate batches of micelles were made using the same protocol and the particle size was measured (Figure 24). The DLS data are plotted using an intensity-weighted distribution. Micelles had a mean diameter of 216.5 nm with a relative standard deviation (RSD) of 8.62 %. As shown with a low RSD value, the micelles consist of a monodispersed population of particles and batch to batch variability appears to be low.
Size Distribution by Intensity
Figure 24 Size distribution of micelles from DLS measurements.
48


Morphology of SRTG Embedded with NPs Using SEM
Scanning electron microscopy (SEM) was used to characterize the morphologies of the SRTG, micelles and the combined delivery system. SEM revealed that the SRTG alone consists of polymer sheets (Figure 25 A), while the micelle NPs were uniformly spherical with the expected core-shell structure (Figure 25B).
Figure 25 SEM images of the SRTG and micelle NPs. A: 5% (w/v) of SRTG cross section showing polymer sheets (scale bar =10 pm). B: Micelle nanoparticles confirming the spherical structure (scale bar = 2 pm). C: SRTG encapsulating micelles cross section showing porous configuration (scale bar = 20 pm). D: Enlargement of SRTG encapsulating micelles with black arrows indicating micelles (scale bar = 1 pm).
Figure 25C shows the morphology of the SRTG encapsulating the micelle NPs. The
morphology of the polymeric gel embedded with micelles is substantially altered from the
original polymer sheets, as the cross-sectional image shows the gel in a highly porous
configuration. Figure 25D is an enlargement of the porous area of the same polymeric gel
49


showing that the micelles are embedded within the gel and their uniform spherical shape has been maintained. The black arrows indicate the micelles in the gel.
In Vitro Multiple Biological Factors Release Study An in vitro release test study was performed to compare the cumulative release rates of the RTG to the SRTG with all three biological factors and the micelles loaded into the gels. 5 wt% polymeric (RTG + VEGF + NPs + PDGF + IL-10 and SRTG + VEGF + NPs + PDGF + IL-10) solutions were created using PBS with 0.2% BSA. 500 ng of each factor was added to the gels and the samples were incubated at 37 C with samples taken daily. The samples were then analyzed using ELISAs to quantify the cumulative release for each protein (Figure 26).
-RTG VEGF -SRTG -VEGF -- RTG IL-10 -SRTG IL-10 -RTG PDGF -SRTG PDGF
Figure 26 Cumulative release profile showing the sequential release of all three factors from the SRTG and plain RTG (n=3 samples).
The release profiles for the RTG and SRTG show that the sulfonate groups are
reducing the amount of protein released for each factor. The SRTG shows a 12.5 % reduction
50


in the burst release of VEGF compared to the plain RTG for the day 0 timepoint. This confirms that the sulfonate groups on the SRTG are mimicking heparin sulfate function and electrostatically bind to VEGF to reduce its release rate. Furthermore, the SRTG slowed the release rate for IL-10 and PDGF slightly. Heparin sulfate binds to areas rich in basic amino acid residues such as arginine and lysine, and it has been shown that IL-10 and PDGF contain these regions and bind to the sulfonate groups on heparin [85,86],
Additionally, both polymer systems displayed sequential release of all three factors, VEGF released first, followed by IL-10 and then PDGF released last. This demonstration of sequentially releasing factors from the polymer scaffold is essential because if PDGF released first, it could interfere with the therapeutic angiogenesis process during the later stage of vessel maturation. If the blood vessels do not mature properly, then this could cause leaky and non-functional vessels. The factors also showed sustained release from the delivery system as the samples showed a consistent release rate of about 0.25 0.5 ng per day through 17 days, after the initial burst release. The GFs need to have this sustained release to help stabilize the neo-vessels during their initial formation through their final development into functional blood vessels. Demonstrating sustained release of IL-10 is also important to reduce an inflammatory response from the polymer system itself, and to limit the amount of inflammation from an MI to minimize the damaging effects of cardiac remodeling.
In Vivo Mouse Subcutaneous Biological Factor Injections In order to assess the angiogenesis and immune response efficacy in vivo, the SRTG polymer delivery system was subcutaneously injected in the middle back of mice. The amount of protein injected for each factor was 250 ng and the following 6 groups were injected: saline, SRTG, SRTG + VEGF, SRTG + VEGF + PDGF, SRTG + VEGF + NPs +
51


PDGF, or SRTG + VEGF + NPs + PDGF + IL-10) for 2 time points (7 days, 21 days). Immunohistochemistry (IHC) was performed to directly identify functional and mature blood vessels, and to identify infiltrating macrophages to quantify a change in the immune response.
IHC Analysis of Therapeutic Angiogenesis Response
IHC was used to identify and quantify the different cell types involved in the process of blood vessel formation. Vascular endothelial cells (ECs) were stained with CD31 and Alexa Flour 488, vascular smooth muscle cells were stained a-SMA and Alexa Flour 594, and ECs associated with the functionality of blood vessels were stained with vWF and Alexa Flour 594. Saline was used as a negative control to compare how many blood vessels normally exist in the subcutaneous tissue and the number of cells (SMCs, ECs, vWF+ cells) that constitute those vessels. Positively stained cells were counted and quantified when DAPI was identified with the stains.
Functional Vascular Endothelial Cell Analysis
Co-staining for CD31 and vWF, a factor involved in hemostasis, was examined to identify the amount of functional vascular endothelial cells within blood vessels for 7 and 21 days after injection (Figure 27). In all experimental injection groups, an increase in functional blood vessels was observed qualitatively by comparing organized rings of ECs (CD31+ and vWF+ cells) to saline injections (Figure 27).
52


Figure 27
SRTG + VEGF
SRTG
Saline
7 day 21 day 7 day 21 day


Figure 27 Representative images of IHC co-staining for ECs (CD31) and blood vessel functionality (vWF) after 7 and 21 days for all 6 groups. Subcutaneous tissue was stained with: CD31 and Alexa Flour 488 and appear green, vWF and Alexa Flour 594 and appear red, and DAPI which appears blue. Scale bar represents 100 pm.
After 21 days, SRTG + VEGF + PDGF, SRTG + VEGF + NPs + PDGF, and SRTG + VEGF + NPs + PDGF + IL-10 groups revealed a higher quantity of functional ECs compared to the day 7 time point demonstrating that the cells are maturing over time (Figure 27)
Staining for CD31 was first examined to compare the proliferation of ECs associated with blood vessel lumen formation. After 7 days, all 5 experimental SRTG groups showed significantly more ECs (CD31+ cells) than saline (Figure 28). After 21 days, all 4 SRTG groups loaded with biological factors showed significantly more ECs than saline (Figure 28). The SRTG group, without any factors, showing significantly more ECs at 7 days, may be attributed to an immune response induced by the polymer injection. It has been shown that tumor necrosis factor-alpha (TNF-a) is secreted by activated macrophages during inflammation and immune response [87], TNF-a then acts on ECs by inducing cell differentiation and by stimulating the production of angiogenic factors from other cells [88],
Saline --SRTG
ASRTG + VEGF SRTG + VEGF + PDGF
*SRTG + VEGF + NP + PDGF -#-SRTG + IL-10 + VEGF + NP + PDGF
Figure 28 Comparison of CD31+ cells between the 6 injected groups quantified from costaining of CD31 and vWF. Error bars represent standard error of the mean. indicates p <
0.05.
54


Additionally, the SRTG + VEGF + NPs + PDGF + IL-10 group displayed significantly more ECs than saline, SRTG, and SRTG + VEGF + PDGF after 7 days and it showed significantly more ECs than saline and SRTG after 21 days. This demonstrates that the biological factors being released from the polymer delivery systems stimulates the proliferation of ECs.
Co-staining for CD31 and vWF, a factor involved in hemostasis, was examined to quantify the amount of functional vascular endothelial cells within blood vessels. After 7 and 21 days, the SRTG + VEGF + NPs + PDGF + IL-10 and SRTG + VEGF + NPs + PDGF groups showed significantly more functional ECs (CD31+ and vWF+ cells) than saline, SRTG and SRTG + VEGF (Figure 29). The SRTG + VEGF group did not show a statistically significant difference of functional ECs compared to the SRTG and saline groups which shows that the delivery of PDGF is important to increase the proliferation of vascular ECs (Figure 29).
Saline SRTG
A SRTG + VEGF * SRTG + VEGF + PDGF
* SRTG + VEGF + NP + PDGF SRTG + IL-10 + VEGF + NP + PDGF
Time [day]
Figure 29 Comparison of CD31+ and vWF+ cells between the 6 injected groups quantified from co-staining of CD31 and vWF. Error bars represent standard error of the mean. indicates p < 0.05.
55


Mature Blood Vessel Formation Analysis
Co-staining for CD31 and a-SMA, a marker for vascular SMCs (mural cells), was examined to identify mature and stable blood vessels for 7 and 21 days after injection (Figure 30). More mature vessels should show an organized circle of ECs surrounded by mural cells. After 21 days, an increase in blood vessels surrounded by mural cells was observed qualitatively for the SRTG + VEGF + NPs + PDGF + IL-10 and SRTG + VEGF + NPs + PDGF groups compared to saline and SRTG injections (Figure 30). Additionally, it appeared that all biological factor injection groups stimulated the proliferation of ECs compared to saline and SRTG, showing that the GFs are inducing a potent angiogenic response.
Staining for CD31 was first examined again to compare the proliferation of ECs associated with blood vessels and similar results were found compared to the CD31 and vWF co-staining results. After 7 days, the SRTG + VEGF, SRTG + VEGF + NPs + PDGF, and SRTG + VEGF + NPs + PDGF + IL-10 groups showed a statistically significant increase in EC proliferation compared to the saline and SRTG groups (Figure 31). After 21 days, all 4 biological factor SRTG groups showed significantly more ECs than saline and SRTG groups (Figure 31). Although the SRTG group did show an increase in ECs after 7 and 21 days compared to saline, it was not significantly different which shows that the TNF-a secreted by macrophages may not be having as much of an angiogenic response as previously thought. Furthermore, SRTG + VEGF + NPs + PDGF group had significantly more ECs after 21 days compared to 7 days (Figure 31).
56


Figure 30
SRTG + VEGF
SRTG
Saline
7 day 21 day 7 day 21 day


Figure 30 Representative images of IHC co-staining for ECs (CD31) and SMCs (a-SMA) to show blood vessel maturation after 7 and 21 days for all 6 groups. Subcutaneous tissue was stained with: CD31 and Alexa Flour 488 and appear green, a-SMA and Alexa Flour 594 and appear red, and DAPI which appears blue. Scale bar represents 100 pm.
Saline SRTG
SRTG + VEGF SRTG + VEGF + PDGF
SRTG + VEGF + NP + PDGF SRTG + IL-10 + VEGF + NP + PDGF
Figure 31 Comparison of CD31+ cells between the 6 injected groups quantified from costaining of CD31 and a-SMA. Error bars represent standard error of the mean. indicates p < 0.05.
Staining for a a-SMA was examined to quantify the formation of mature blood vessels induced by the polymer delivery systems. After 7 days, the SRTG + VEGF + NPs + PDGF + IL-10 group induced significantly more mural cells than the saline, SRTG and SRTG + VEGF groups (Figure 32). After 21 days, the SRTG + VEGF + NPs + PDGF + IL-10 and SRTG + VEGF + NPs + PDGF groups both displayed significantly more mural cells compared with the saline, SRTG and SRTG + VEGF + PDGF groups (Figure 32). Furthermore, the SRTG + VEGF, SRTG + VEGF + NPs + PDGF and SRTG + VEGF + NPs + PDGF + IL-10 groups all induced significantly more mural cells over time when comparing the 7 and 21 day time points (Figure 32). This demonstrates that the GFs are
58


inducing more mature neovasculature from the early to late stages of the angiogenesis process.
--Saline --SRTG
-A-SRTG + VEGF HSRTG + VEGF + PDGF
* SRTG + VEGF + NP + PDGF SRTG + IL-10 + VEGF + NP + PDGF
Figure 32 Comparison of a-SMA+ cells between the 6 injected groups quantified from costaining of CD31 and a-SMA. Error bars represent standard error of the mean. indicates p <
0.05.
The SRTG +VEGF + PDGF group did not show any increase in mural cells between days 7 and 21, unlike the other delivery systems loaded with GFs (Figure 32). This group does not have PDGF encapsulated within NPs so both VEGF and PDGF are most likely burst releasing simultaneously. As stated earlier, it has been shown that early stage angiogenic factors can have inhibitory effects on late stage GFs and vice versa, when presented concurrently [5,17,18,89], This data for the SRTG +VEGF + PDGF group further confirms the importance of sequentially delivering angiogenic GFs to develop stable and mature blood vessels.
59


IHC Analysis of Immune Response
IHC was used to identify and quantify the number of macrophages present in the subcutaneous tissue and whether there was a reduced immune response from the polymer delivery system loaded with IL-10. Macrophages were stained with CD68 and Alexa Flour 594. Saline was used as a negative control to compare how many macrophages infiltrate the tissue from the needle injection. Positively stained cells were counted and quantified when DAPI was identified with the stains. The IHC stains qualitatively show a decrease in the number of macrophages for the SRTG + VEGF + NPs + PDGF + IL-10 group compared to all other polymer delivery groups at 7 and 21 days (Figure 33). Moreover, the inflammatory marker showed a large number of CD68+ cells for all polymer groups (Figure 33).
7 day 21 day f day 21 day
Figure 33 Representative images of IHC staining for macrophages (CD68) to show immune response after 7 and 21 days for all 6 groups. Subcutaneous tissue was stained with CD68 and Alexa Flour 594 and appear red and DAPI which appears blue. Scale bar represents 50 pm.
60


The number of macrophages were additionally quantified by counting CD68+ and DAPI+ cells at the boundary area of the polymer injection site. The saline injection group showed significantly less macrophages than all polymer delivery groups at 7 and 21 days (Figure 34). After 21 days, the SRTG + VEGF + NPs + PDGF + IL-10 group showed significantly less CD68+ cells than the 4 other polymer biological factor delivery groups (Figure 34). This result demonstrates that IL-10 is significantly reducing the immune response caused by the polymer gel. Additionally, the SRTG + VEGF, SRTG + VEGF + PDGF and SRTG + VEGF + NPs + PDGF + IL-10 all showed a significant reduction is macrophages from 7 to 21 days (Figure 34).
Saline SRTG
SRTG + VEGF SRTG + VEGF + PDGF
SRTG + VEGF + NP + PDGF SRTG + IL-10 + VEGF + NP + PDGF
Figure 34 Comparison of CD68+ cells between the 6 injected groups, bars represent standard error of the mean. indicates p < 0.05.
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CHAPTER VII
CONCLUSION
In this study, a sulfonated reverse thermal gel encapsulating novel micelle NPs was investigated for the application of localized and controlled protein delivery to induce angiogenesis. The polymer delivery system was synthesized, characterized and evaluated for its therapeutic potential with an in vitro release test and subsequent in vivo animal study.
Sulfonated-PSHU-PNIPAm was successfully synthesized and the molecular structure was verified through NMR, FTIR and elemental analysis. Polymeric PEG-PSHU-PEG micelles were fabricated and characterized by NMR, DLS and SEM to validate the chemical structure and to ensure uniform spherical size distribution. Additionally, SEM was further used to demonstrate micelle encapsulation within the sulfonated thermal gel.
The entire polymer delivery system was then evaluated for controlled and sequential release of three biological factors using an in vitro release study. Release test samples were taken daily and protein release was then quantified using ELISAs. The results showed that the sulfonated-PSHU-PNIPAm system significantly reduced the burst-release of VEGF from the polymer scaffold compared to plain PSHU-PNIPAm, demonstrating the effectiveness of sulfonate groups controlling VEGF release. Furthermore, the release profile showed that the encapsulated micelles were being slowly released from the SRTG which provided the sequential release of PDGF. This spatiotemporal release of GFs from the polymer scaffold is critical to promote stable and mature angiogenic vessel formation.
The ability of the polymer system to induce new blood vessel formation was analyzed in vivo using a subcutaneous injection mouse model. Subcutaneous injections of the SRTG polymer system with 4 different combinations of factors was evaluated for angiogenesis
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activity against saline and SRTG injections. Histological assessment was used to directly observe blood vessel formation. Functional vascular ECs and vascular SMCs were quantified using three different stains and were subsequently counted as positive cells when identified with DAPI. After 21 days, the SRTG + VEGF + NPs + PDGF + IL-10 group showed significantly more functional ECs and SMCs than every other group, except for the SRTG + VEGF + NPs + PDGF group. These results show that the polymer delivery system with all three biological factors is inducing a substantial angiogenic response with stable and more mature vessels.
The polymer system was analyzed for a reduced immune response as well. IHC was used to examine the effects of IL-10 on suppressing macrophage infiltration at the polymer injection site. A stain for macrophages was quantified by counting the number of CD68+ cells identified with DAPI. After 21 days, the SRTG + VEGF + NPs + PDGF + IL-10 group showed significantly less macrophages present in the subcutaneous tissue compared to all other polymer groups. These results are encouraging as limiting an immune response is important for optimizing cardiac repair following an MI.
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CHAPTER VIII
FUTURE WORK Increase Sample Size
For several of the groups, statistically significant differences in CD31+, vWF+ and a-SMA+ cells were not seen between the groups themselves and between the time points. This is due to the high variability from differing amounts of blood vessels while counting positive cells from the selected images. Increasing the sample size would provide more tissue to assess histologically and this could reduce variability which may allow for observation of significant statistical differences between the groups time points.
Additional Time Points
Although statistical significant differences between some of the groups were not observed, different trends between the groups were seen from 7 to 21 days that could be further elucidated with a longer time point at 4 or 6 weeks. Many mature blood vessels were observed from counting vWF+ and a-SMA+ cells after 21 days, but an additional time point to allow the vessels to mature further could provide a more apparent difference between the groups.
Increase Biological Factor Loading Amount
To further increase the statistical significant differences between the groups, increasing the amount of proteins loaded into the polymer delivery system could enhance the amount of functional and mature blood vessels. GFs have a short half-life in vivo and are quickly degraded in circulation so increasing the local amount of GFs to the tissue could promote higher EC proliferation rates while also recruiting more SMCs to the vessels. Additionally,
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increasing the amount of IL-10 could help reduce the amount of inflammation caused by the polymer system.
Acute Myocardial Infarction Animal Model
As the target of this polymer delivery system is for treatment of an MI, an acute MI mouse model should be implemented to examine the effects of revascularizing ischemic myocardium and reducing detrimental cardiac remodeling. Validating the efficacy of this biological factor delivery system for treatment of MI is essential to determine if this system can produce therapeutic angiogenesis effects in the heart. IL-10 has been shown to attenuate post-MI left ventricular dysfunction, reduce infarct size and attenuate Mi-induced cardiac cell death [33], so it is important to demonstrate that this polymer system can deliver IL-10 and provide those same benefits. Echocardiography, IHC, TUNEL staining and Western blot analysis will all be utilized to examine the therapeutic effects of this protein delivery on treatment for MI.
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EVALUATION OF AN INJECTABLE POLYMERIC DELIVERY SYSTEM FOR CONTROLLED AND LOCALIZED RELEASE OF BIOLOGICAL FACTORS TO PROMOTE THERAPEUTIC ANGIOGENESIS by ADAM JOHN ROCKER B.A., University of Colorado Boulder, 2012 A thesis submitted to the Faculty of the Graduate School of the University of the Colorado in partial fulfillment of the requirements for the degree of Master of Science Bioengineering Program 2016

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ii This thesis for the Master of Science degree by Adam John Rocker has been approved for the Bioengineering Program by Daewon Park, Chair Karin Payne Danielle Soranno Date: December 17, 2016

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iii Rocker Adam John (M.S. Bioengineering) Evaluation of an Injectable Polymeric Delivery System for Controlled and Localized Release of Biological Factors to Promote Therapeutic Angiogenesis Thesis directed by Assistant Professor Dae w on Park ABSTRACT Cardiovascular disease remains as the leading cause of death worldwide and is frequently associated with partial or full occlusion of coronary arteries. Currently, angioplasty and bypass surgery are the standard approaches for treating patients with these ischemic heart conditions. However, a large number of patients cannot undergo these procedures. Therapeutic angiogenesis provides a minimally invasive tool for treating cardiovascular diseases by inducing new blood vessel growth from the existing vasculatu re. Angiogenic growth factors can be delivered locally through gene, cell, and protein therapy. Natural and synthetic polymer growth factor delivery systems are under extensive investigation due their widespread applications and promising therapeutic poten tial. Although biocompatible, natural polymers often suffer from batch to batch variability which can cause unpredictable growth factor release rates. Synthetic polymers offer advantages for growth factor delivery as they can be easily modified to control release kinetics During the angiogenesis process, vascular endothelial growth factor (VEGF) is necessary to initiate neo vessel formation while platelet derived growth factor (PDGF) is needed later to help stabilize and mature new vessels. In the setting o f myocardial infarction, additional anti inflammatory cytokines like IL 10 are needed to help optimize cardiac repair and limit the damaging effects of inflammation following infarction. To meet these angiogenic and anti inflammatory needs, an injectable p olymer delivery system created from a sulfonated reverse

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iv thermal gel encapsulating micelle nanoparticles was designed and evaluated. The sulfonate groups on the thermal gel electrostatically bind to VEGF which controls its release rate, while the micelles are loaded with PDGF and are slowly released as the gel degrades. IL 10 was loaded into the system as well and diffused from the gel over time. An in vitro release study was performed which demonstrated the sequential release capabilities of the polymer sy stem. The ability of the polymer system to induce new blood vessel formation was analyzed in vivo using a subcutaneous injection mouse model. Histological assessment was used to quantify blood vessel formation and an inflammatory response which showed that the polymer delivery system demonstrated a significant increase in functional and mature vessel formation while significantly reducing inflammation. 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 express my appreciation to my professors, mentors, friends, and family for all of their supp ort throughout my research p roject. Completing this degree would not have been possible without them. I would like to thank my advisor, Dr. Daewon Park, for his guidance and for allowing me the opportunity to work in the Translational Biomaterials Research Laboratory. Dr. Park has always pushed me to do my best, which has helped le a d me through this journe y. I would also like to express my gratitude to Dr. Karin Payne and Dr. Danielle Soranno fo r their insi ght and teachings throughout my graduate school career I truly could not be more thankful for all the assistance and encouragement I received from my fellow labmates, especially David Lee, James Bardill and Melissa Laughter. I will always cherish the times we spent together in and out of lab, and I look forward to seeing what the future holds for us. Finally, I would like to express my gratitude for my family. My dad has been the greatest role model in my life and I could not have done this without his constant motivation. My mom has always been there for me and she was always available when I needed help with anything throughout my life. To my brother and sister, thank you for providing me with continuous encouragement and challenging me to do my best in life. Animal model studies were conducted under the University of Colorado at Denver Institutional Animal Care and Use Committee ( IACUC ) protocol number 102913(12)2 D

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vi TABLE OF CONTENTS CHAPTER I. IN TRODUCTION ................................ ................................ ................................ .. 1 Overview ................................ ................................ ................................ ................. 1 Clinical Motivation ................................ ................................ ................................ 2 Anatomy and Physiology ................................ ................................ ......... 6 Cardiovascular Disease ................................ ................................ ........................... 8 Pathophysiology ................................ ................................ ................................ ...... 9 The Inflammatory Response in Myocardial Infarction .......................... 10 Current CAD Treatment ................................ ................................ ....................... 11 Study Objective ................................ ................................ ................................ ..... 13 II. Background ................................ ................................ ................................ ........... 14 Therapeutic Angiogenesis Using Growth Factors ................................ ................ 14 Therapeutic Angiogenesis ................................ ................................ ...... 14 Growth Factor Considerations ................................ ............................... 15 Anti inflammatory Factor Considerations ................................ ............. 16 Current Biological Factor Delivery Methods ................................ ........ 16 Polymeric Biomaterials for Protein Delivery ................................ ....................... 17 Sustained Growth Factor Release ................................ .......................... 18 Modulating Physical and Chemical Interactions ................................ ... 18

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vii Thermally Induced Gelling Systems ................................ ...................... 19 Nanoparticles for Sequential Growth Factor Release ................................ ........... 21 Nanoparticle Considerations ................................ ................................ .. 21 III. PREVIOUS WORK ................................ ................................ .............................. 24 SRTG LCST ................................ ................................ ................................ .......... 24 Cytotoxicity ................................ ................................ ................................ ........... 25 BSA Release Test ................................ ................................ ................................ 26 CD Spectroscopy ................................ ................................ ................................ .. 28 IV. HYPOTHESIS AND SPECIFIC AIMS ................................ ............................... 29 Hypothesis ................................ ................................ ................................ ............. 29 Specific Aims ................................ ................................ ................................ ........ 29 V. MATERIALS AND METHODS ................................ ................................ .......... 30 Materials ................................ ................................ ................................ ............... 30 Equipment ................................ ................................ ................................ ............. 31 Polymer Synthesis ................................ ................................ ................................ 32 N BOC seri nol synthesis ................................ ................................ ........ 32 Poly(serinol hexamethylene urea) or PSHU backbone synthesis .......... 32 PSHU deprotection ................................ ................................ ................ 33 PNIPAm Synthesis ................................ ................................ ................. 33 PSHU PNIPAm Conjugation (RTG Synthesis) ................................ .... 34

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viii Sulfonation of PSHU PNIPAm (SRTG Synthesis) ............................... 34 PEG PSHU PEG Micelle Nanoparticle Polymer Synthesis .................. 35 Micelle Fabrication ................................ ................................ ................ 35 Polymer Characterization ................................ ................................ ...................... 36 Elemental Analysis ................................ ................................ ................ 36 Dynamic Light Scattering (DLS) ................................ ........................... 36 Scanning Electron Microscopy (SEM) ................................ .................. 36 Proton Nuclear Magnetic Resonance ( 1 H NMR) ................................ ... 37 Fourier Transform Infrared (FTIR) Spectroscopy ................................ 37 In Vitro Biological Factor Release Study ................................ ............................. 37 Enzyme Linked Immunosorbent Assay (ELISA) ................................ .. 38 In Vivo Mouse Subcutaneous Biological Factor Injections ................................ 39 Euthanasia and Tissue Harvest ................................ ................................ ............. 40 Histology ................................ ................................ ................................ ............... 40 Quantitative Analysis of Immunofluorescent Staining ................................ ......... 41 Statistical Analysis ................................ ................................ ................................ 41 VI. RESULTS AND DISCUSSION ................................ ................................ ........... 42 Synthesis of Sulfonated PSHU PNIPAm Reaction Sequence ............................. 42 Synthesis of PEG PSHU PEG Reaction Sequence ................................ .............. 43 PSHU and dPSHU Characterization 1 H NMR ................................ ...................... 44

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ix PSHU PNIPAm and PEG PSHU PEG Characterization Using 1 H NMR ........... 45 PSHU PNIPAm and Sul PSHU PNIAPAm Characterization Us ing FTIR ......... 46 SRTG Sulfonation Detection Using Elemental Analysis ................................ ..... 47 PEG PSHU PEG Micelle Nanoparticles Size Distribution Using DLS ............... 48 Morphology of SRTG Embedded with NPs Using SEM ................................ ..... 49 In Vitro Multiple Biological Factors Release S tudy ................................ ............. 50 In Vivo Mouse Subcutaneous Biological Factor Injections ................................ 51 IHC Analysis of Therapeutic Angiogenesis Response .......................... 52 Functional Vascular Endothelial Cell Analysis ................................ ..... 52 Mature Blood Vessel Formation Analysis ................................ ............. 56 IHC Analysis of Immune Response ................................ ....................... 60 VII. CONCLUSION ................................ ................................ ................................ ..... 62 VIII. CHAPTER VIII ................................ ................................ ................................ .... 64 IX. FUTURE WORK ................................ ................................ ................................ .. 64 Increase Sample Size ................................ ................................ ............................ 64 Additional Time Points ................................ ................................ ......................... 64 Increase Biological Factor Loading Amount ................................ ........................ 64 Acute Myocardial Infarction Animal Model ................................ ........................ 65 REFERENCES ................................ ................................ ................................ ....................... 66

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x LIST OF FIGURES FIGURE 1 Molecular basis of vessel branching [13]. ................................ ................................ ......... 4 2 Structure of Blood Vessels. (a) Arteries and (b) veins share the same general features, but the walls of arteries are much thicker because of the higher pressure at which blood flows through them. (c) A micrograph shows the relative differences in thickness [25]. ................................ ................................ ................ 7 3 (A) Progression of heart failure from a coronary artery occlusion. (B) The process of therapeutic induced angiogenesis from local GF release [9]. .......................... 9 4 Simplified schematic of diversity of lesions in human coronary atherosclerosis [32]. ................................ ................................ ................................ ......... 10 5 Diagram of Angioplasty [40]. ................................ ................................ .......................... 12 6 Diagram of CABG surgery [41]. ................................ ................................ ..................... 12 7 bound to sulfur, as well as the remai ning chemical makeup of the polymer. ................. 19 8 Polymer solubility behavior at the LCST. Left hand side shows hydrated polymer below LCST with entropic loss of water and chain collapse above LCST (right hand side) [69]. ................................ ................................ ........................... 20 9 Schematic representation of diff erent NP systems [71]. ................................ .................. 23 10 Temperature dependent phase transition of sulfonated PSHU PNIPAm. A: The sulfonated PSHU PNIPAm und ergoes a sharp, reversible phase transition around 32 as determined by UV Visible spectroscopy. B: An aqueous solution of sulfonated PSHU PNIPAm at room temperature (C) turns to physical gel at 37 [22]. ................................ ................................ ................... 24 11 In vitro cytotoxicity of sulfonated PSHU PNIPAm by MTT assay. Results demonstrated no cytotoxic effects of sulfonated PSHU PNIPAm on C2C12 cells after exposure to the poly mer extract in medium, while 10% DMSO shows significant cytotoxicity. There is no statistical difference between medium and extracts. Data represent mean SD. indicates p value < 0.05 test) [22]. ................................ ................................ ................................ ...... 26

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xi 12 BSA release profiles from sulfonated PSHU PNIPAm and plain PSHU PNIPAm PNIPAm 15%. BSA release was more sustained from sulfonated PSHU PNIPAm than that from plain PSHU PNIPAm, with a more sustained profile at higher concentrations. Data represent mean SD [22]. ................................ .............................. 27 13 CD spectra of n ative and released BSA. The similarity of the protein conformation between samples indicates that the secondary structure of the BSA protein was well preserved by the system [22]. ................................ ...................... 28 14 Formed gels in 2 mL vials for in vitro release study. ................................ ...................... 38 15 Demonstra tion of subcutaneous injection in the middle back of a mouse [82]. .............. 40 16 Synthesis Reaction sequence of sulfonated PSHU PNIPAm s ynthesis [22]. .................. 42 17 Reaction sequence of PEG PSHU PEG synthesis. ................................ .......................... 43 18 1 H NMR spectrum of PSHU confirming the molecular structure [83]. ........................... 44 19 1 H NMR spectrum of PSHU and dPSHU confirming the removal of the BOC protecting group with the loss of the b peak [83]. ................................ ........................... 44 20 1 H NMR spectrum of PSHU PNIPAm. Successful conjugation of PNIPAm was confirmed by the presence of methylene and methyl protons at 1.55 and 1.09 ppm, respectively. ................................ ................................ ................................ .... 45 21 1 H NMR spectra of PEG PSHU PEG. Successful conjugation of PEG was confirmed by the presence of the peak at 3.51 ppm which identifies the protons on the P EG repeating unit of the polymer backbone [84]. ................................ 45 22 FTIR spectra of sulfonated PSHU PNIPAm and PSHU PNIPAm. ................................ 46 23 FTIR spectra of sulfonated PSHU PNIPAm and PSHU PNIPAm with an enlargement of the sulfonate peak. ................................ ................................ .................. 47 24 Size distribution of micelles from DLS measurements. ................................ .................. 48 25 SEM images of the SRTG and micelle NPs. A: 5% (w/v) of SRTG cross section showing polymer sheets (scale bar = 10 m). B: Micelle nanoparticles confirming the spherical structure (scale bar = 2 m). C: SRTG encapsulating micelles cross section showing poro us configuration (scale bar = 20 m). D: Enlargement of SRTG encapsulating micelles with black arrows indicating micelles (scale bar = 1 m). ................................ ...................... 49 26 Cumulative release profile showing the sequential release of all three factors from the SRTG and plain RTG (n=3 samples). ................................ ............................... 50

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xii 27 Representative images of IHC co staining for ECs (CD31) and blood vessel functionality (vWF) after 7 and 21 days for all 6 groups. Subcutaneous tissue was stained with: CD31 and Alexa Flour 488 and appear green, vWF and Alexa Flour 594 and appear red, and DAPI which appears blue. Scale bar represents 100 m. ................................ ................................ ................................ .... 53 28 Comparison of CD31+ cells between the 6 injecte d groups quantified from co staining of CD31 and vWF. Error bars represent standard error of the mean. indicates p < 0.05. ................................ ................................ .............................. 54 29 Comparison of CD31+ and vWF+ cells between the 6 injected groups quantified from co staining of CD31 and vWF. Error bars represent standard error of the mean. indicates p < 0.05. ................................ ................................ ........... 55 30 Representative images of IHC co SMA) to show blood vessel maturation after 7 and 21 days for all 6 groups. Subcutaneous tissue was stained with: CD31 and Alexa Flour 488 and SMA and Alexa Flour 594 and a ppear red, and DAPI which appears blue. Scale bar represents 100 m. ................................ ................................ .... 57 31 Comparison of CD31+ cells between the 6 injected grou ps quantified from co SMA. Error bars represent standard error of the mean. indicates p < 0.05. ................................ ................................ .............................. 58 32 C SMA+ cells between the 6 injected groups quantified from co SMA. Error bars represent standard error of the mean. indicates p < 0.05. ................................ ................................ .............................. 59 33 Representative images of IHC staining for macrophages (CD68) to show immune response after 7 and 21 days for all 6 groups. Subcutaneous tissue was stained with CD68 and Alexa Flour 594 and appear red and DAPI which appears blue. Scale bar represents 50 m. ................................ ............................ 60 34 Comparison of CD68+ cells between the 6 injected groups. bars repre sent standard error of the mean. indicates p < 0.05. ................................ ............................ 61

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xiii LIST OF TABLES TABLE 1 Elemental analysis of PSHU PNIPAm (RTG) and sulfonated PSHU PNIPAm (SRTG). Polymers were analyzed for carbon, hydrogen, nitrogen, oxygen and sulfur content and were determined by weight percent for each element. ................................ ................................ ................................ ........................... 48

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xiv LIST OF ABBREVI ATIONS 1 H NMR proton nuclear magnetic resonance ABTS Azino bis(3 ethylbenzothiazoline 6 sulfonic acid) ACA azobis(4 cyanovaleric acid) ANOVA analysis of variance bFGF basic fibroblast growth factor BSA bovine serum albumin CABG coronary artery bypass graft surgery CAD GF coronary Artery Disease growth Factor CVD cardiovascular disease DLS dynamic light scattering DMF N,N dimethylformamide DMSO dimethyl sulfoxide dPSHU deprotected PSHU EC endothelial cell ECM extracellular matrix EDC N (3 dimethylamino propyl) ethylcarbodiimide hydrochloride ELISA enzyme linked immunosorbent assay FTIR fourier transform infrared spectroscopy GF growth factor H&E haematoxylin and eosin HDI hexamethylene diisocyanate IHC immunohistochemistry IL 10 interleukin 10 IR infrared LCST lower critical solution temperature MI myocardial infarction MMP matrix metalloproteinase MTT 3 (4,5 Dimethylthiazol 2 yl) 2,5 diphenyltetrazolium bromide MW molecular weight N 2 nitrogen NH 2 amine group NHS N Hydroxysuccinimide NIPAm N isopropylacrylamide nm nanometers NP nanoparticle OCT optimal cutting temperature PBS phosphate buffered saline PDGF platelet derived growth factor

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xv PEG poly(ethylene glycol) PLGA poly(lactic co glycolic acid) PNIPAm poly(N isopropylacrylamide) PS 1,3 propane sultone PSHU Poly(serinol hexamethylene urea) rpm revolutions per minute RSD relative standard deviation RTG reverse thermal gel (used interchangeably with PSHU PNIPAm) SEM scanning electron microscopy SMC smooth muscle cell SRTG sulfonated reverse thermal gel (used interchangeably with sulfonated PSHU PNIPAm) t BuOK potassium tert butoxide TFA trifluoroacetic acid TNF tumor necrosis factor alpha UV ultraviolet VEGF vascular endothelial growth factor vWF von Willebrand factor SMA alpha smooth muscle actin

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1 CHAPTER I INTRODUCTION Overview The accumulation of plaque in the coronary arteries of the heart, accrued from ado lescence through adulthood often leads to a blockage in blood flow which results in a myocardial infarction (MI) This deadly disease, termed coronary artery disease (CAD), is the most common type of heart disease a nd claimed the lives of over 370,000 individuals in 2010 in the United States [1] Due to the abundant number of people this illness affects, the projected total costs of CAD in 2015 are around $182 billion in the United States with the cost of healthcare services, medications, and lost productivity taken into account [2] This is projected to increase to $322 billion by 2030 [2]. Although some non invasive treatments for CAD exist, such as medications and lifestyle changes, many individuals still must undergo costly surgical i nterventions. However, there is a su bset of patients that are unable to receive these surgical procedures due to the presence of various comorbidities. These patients, along with others seeking minimally invasive treatment options for CAD, have influenced a push in research towards developin g growth factor (GF) delivery system s for therapeutic angiogenesis in the pursuit of revasculari zing the damaged heart tissue after a MI. An injectable biological factor polymer delivery system that can revascularize heart tissue could dramatically reduce healthcare costs related to CAD, which would also improve the quality of life for millions of people suffering from this disease. The process of treating CAD can involve many different tests and expensive medications, in addition to the almost inevitable o pen heart surgery that must be implemented. A standard injectable protein delivery system to stimulate new blood vessel formation around a blocked artery will allow

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2 many people to avoid paying these burdensome healthcare bills. Incorporating a supplementar y anti inflammatory protein into this system would also minimize myocardial necrosis and optimize cardiac repair following MI. Additionally, the polymer system may help with reducing the number of heart attacks experien ced by patients because the delivery system could be injected before a complete occlusion of a coronary artery occurs further reducing h ealthcare costs. Despite advances in polymer delivery system s for CAD treatment localized and controlled release of proteins to revascularize damaged heart tissue and reduce the detrimental effects of an inflammatory response remain inadequate. Clinical Motivation Cardiovascular disease (CVD) is listed as the number one cause of death worldwide by the World Health Organization. This disease represents a myr iad of issues pertaining to human health, the medical field, and the economy as this disease causes high morbidity and mortality rates while adding to rising healthcare costs [1] CAD a type of CVD, is characterized by the accumulation of plaque in the coronary arteries of the heart that often leads to a blockage in bloo d flow which results in a MI commonly caused by atherosclerosis [1 3] The insufficient blood supply to a region of the heart during a MI causes cell death and pathological remodeling which often leads to heart failure [1,2,4] In an attempt to prevent heart failure and resupply blood flow to the damaged heart muscle, much research has been c onducted on therapeutic angiogenesi s with delivering GFs to the infarct site [3,5 8] Therapeutic ang iogenesis aims to form new blood vessels from the existing vasculature in order to restore blood flow to the affected ischemic heart tissue [3,8,9] This therapy involves the delivery of GFs to the damaged heart tissue through various methods, with certain studie s having reached clinical trials However, many studies have failed to show

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3 efficacy for this angiogenic process due to a failure in providing sequential and long term delivery of GFs [5,7,10] Clinical trials using bolus injections of single GFs were unsuccessful due to a loss in the bioactivity of the protein in addition to a lack of other supportive GFs necessary for stable angiogenesis [7,11,12] Further research on understanding the complic ated mechanisms behind angiogenesis is needed to develop an effective therapeutic angiogenesis system The main mechanism behind the sprouting of new capillaries works through the release of GFs, such as vascular endothelial growth factor ( VEGF ) and basic fibroblast growth factor (b FGF ) which signals for the migration and proliferation of endothelial cells to differentiate into guiding tip cells or proliferating stalk cells [13] Both cell types work together to form direc tional, elongating, new vessels. Later, a maturation process occurs involving pericytes (mural cells) which are stimulated by other GFs to cover the endothelial cells [14] VEGF and PDGF are well studied GFs and have been tested in human clinical trials [15,16] Although these are key factors in the angiogenesis process, the GFs delivered alone may result in immature blood vessel formation which may regress over time [17] Additionally, it has been demonstrated that when pre sented simultaneously, early stage factors can have hindering effects on late stage GFs, and vice versa [18 20] Therefore, it is essential that the delivery system administers the GFs sequentially, while in their bioactive conformations to mimic their physiological mechanism during angiogenesis. The molecular process of new vessels branching from the existing vasculature is shown in Figure 1 The various GFs involved in this complex mechanism are shown, along with how individual factor s affect each step in the process and how they influence the behavior of different cell types.

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4 Figure 1 Molecular basis of vessel branching [13] To implement this spatiotemporal aspect while protectin g the bioactivity of VEGF and PD GF, a controlled polymeric delivery system composed of nanopartic les encapsulated

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5 within a sulfonated reverse thermal gel ( S RTG) is proposed The S RTG is sulfonated to mimic heparin sulfate function, which is a pr oteoglycan with an intrinsic negative charge that stores, protects and stabilizes positively charged hepari n binding proteins such as VEGF [21,22] This will allow for hydrogen and ionic bonding interactions between VEGF and the sulfonate grou ps to reduce the burst release of the GF while also protecting the GF from degradation The thermal ge l l ing properties will be utilized to encapsulate the nanoparticles which provide s the sequential release of PDGF The S RTG is liquid at room temperature, where it will be mixed with the nanoparticles containing PDGF then once it is injected at body temperature, the pol ymer will rapidly gel causing the nanoparticles to be entrapped. As this system is intended to revasculariz e myocardium following a MI, a third protein IL 10 will be released from the polymer scaffold IL 10 will help minimize the deleterious effects of myocardial necrosis and promote optimal cardiac repair by reducing the inflammatory response after MI It has been shown that this cytokine may have a role in suppressing the acute inflammatory response and in modulating extracellular matrix metabolism [23] Harmful cardiac remodeling is the leading cause of heart failure and death and this delivery system may promote more effective tissue repair which could help reduce compensatory remodeling [24] This interleukin also provides added benefits in that it can help lessen the immune response induced by the injection of the foreign polymer delivery system. Due to its biodegradable properties, the thermal gel wi ll degrade over time, releasing the proteins while also releasing the nanoparticles, to provide the sequential delivery of PD GF.

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6 Anatomy and Physiology deoxygenated blood is r eturned to the heart through systemic veins. These different types of vessels vary slightly in their structures but generally have the same features. Arteries and arterioles have thicker walls than veins and venules as they are closer to the heart and must receive blo od from a large aortic flow at far greater pressure ( Fi gure 2 A B). B lood flows through the lumen of these vessels and arteries have smaller lumens than veins to maintain the pressure of blood moving through the system. These combined characteristics giv e arterial lumens a more rounded appearance in cross sections than the lumens of veins ( Fi gure 2 C) [25 27] Both arteries and veins have the same three distinctive tissue layers which are called tunics. The tunica layers, from the most outer layer to the interior, are the tunica externa, the tunica media, and the tunica intima. The tunica intima consists of an endothelium layer which appears smooth in veins and wavy in arteries due to the constriction of smooth muscle cells (SMCs). The tunica media generally consists of SMCs and elastic fibers in arteries and this layer tends to be the thickest in these vessels. For veins, the tunica media also consists of SMCs but is mostly filled with c ollagenous fibers and is normally thinner than the tunica externa. The tunica externa is the thickest layer in veins, and contains some SMCs with predominately collagenous and smooth fibers. Arteries tend to have a thinner tunica externa compared to the tunica media, except for in the largest arteries, and this layer contains collagenous and elastic fibers. Smal l blood vessels consist only of endothelial cells (ECs), while larger vessels are surrounded by mural cells. These mural cells consist of pericytes in medium sized vessels and SMCs in large vessels [13,25,28]

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7 Fi gure 2 Structure of Blood Vessels. (a) Arteries and (b) veins share the same general features, but the walls of arteries are much thicker because of the higher pressure at which blood flows through them. (c) A micrograph shows the relative differences in thickness [25]

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8 Cardiovascular Disease CVD represents a myriad of issues pertaining to human health, the medical field, and the economy as this disease causes high morbidity and mortality rates while adding to rising healthcare costs [7] In the United States, the prevalence of this disease is stunning with an estimated one in three adults having one or more types of CVD [3] Coronary artery disease (CAD) is one of severa l classifications of CVD where blood flow in the coronary arteries becomes blocked by excess plaque build up [7] This blockage prevents blood flow into the vascular heart tissue, lead ing to symptoms that can range from mild angina and shortness of breath to more severe issues such as a myocardial infarction if the CAD goes unnoticed for many years [7] Once the local oxygen supply decreases significantly from a coronary artery occlusion, the tissue will respon d to hypoxia by increasing the transcription of proangiogenic factors, cytokines and matrix metallo proteinases (MMPs) The myocardium tries to restore oxygen supply and replace the damaged tissue, but frequently these adaptive responses are not effective and myocardium hypertrophy occurs. This process causes a permanent injury that would lead to heart failure if left untreated ( Figure 3 A) [9] The extent o f blood vessel damage and the progression of the disease impact s the treatment plan for the patient. If a local controlled release of angiogenic factors is implemented following the heart injury, the natural process of remodeling and angiogenesis would be enhanced. Myocardial functional recovery and effective revascularization could ultimately be achie ved over time using this factor based treatment approach ( Figure 3 B) [9]

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9 Figure 3 (A) Progression of heart failure from a coronary artery occlusion (B) T he process of therapeutic induced angiogenesis from local GF release [9] Pathophysiology Lack of blood flow to the myocardi um is characterized as myocardial ischemia. Occlusion of the coronary vessels in CAD is almost always a result of atheromatous plaque narrowing the arteries [29] Beginning i n childhood, early atheroma starts to develop and eventually leads to mature cholesterol rich, subintimal plaque formation over time [30] Atheroma mainly affects the intima layer of the vessel and can be attributed to certa in risk factors such as smoking, hypertension, obesity, diabetes, hypercholesterolaemia, sedentary lifestyle, excessive alcohol intake, genetic predisposition for CVD and others [31,32] The mature plaq ue is made of two components produced by two different cell types. Necrotic foam cells which migrate into the intima and ingest lipids forms the lipid core. While the connective tissue matrix is resultant from SMCs, which have migrated from the media to the intima, where they expand and change th eir phenotype to form a fibrous envelope around the lipid core [32] If the atheroma is only partially blocking the lumen of the artery, then the downstream physiological effects are generally mild and can range from asymptomatic ischemia to angina and a decrease in exercise tolerance [29] In more life threatening

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10 situations, acute coronary events can arise when thrombus development foll ows disruption of a plaque ( Figure 4 ) Figure 4 Simplified schematic of diversity of lesions in human coronary atherosclerosis [32] Intimal injury leads to removal of the thrombogenic matrix and generates thrombus formation [29] This thrombus formation induces a MI which prompts acute cardiac failure. C ardiomyocytes not receiving adequate blood flow in the infarct region can become necrotic within hours of the event [30] In an attempt to reduce the number of apoptotic cells from the M I, or as a way t o prevent the MI altogether, the potential of delivering GFs in a local and controlled manner to promote therapeutic angiogenesis to increase blood flow to proximate cells was investigated The Inflammatory Response in Myocardial Infarctio n Determining ideal strategies for the minimization of myocardial necrosis and optimization of cardiac repair following an MI is one of the most important therapeutic targets of modern cardiology [23,24] Cardiac pathophysiological conditions including MI and ischemia reperfusion injury leading to heart failure have been linked with stim ulation of

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11 inflammatory mediators in the heart [33 35] Myocardial necrosis after M I generates complement activation and free radical formation, which triggers a cytokine cascade initiated by the release of tumor necrosis factor alpha [23] Ischemia reperfusion injur y leading to heart failure is associated with intense inflammation as chemokines, cytokines and the complement system recruit neutrophils to th e ischemic and reperfused area [33 35] Expression of pro inflammatory cytokines such as tumor necrosis factor 1 and interleukin 6, and anti inflammatory cytokine IL 10, all faci litate homeostasis within the heart in response to injury [33] However, long term expression of these inflammatory mediators at abundantly high levels could lead to an adv erse consequence in the failing heart [36,37] Sustained inflammatory response connected with increased MMPs production may lead to excessive ECM degradation in the early phase of MI, which can impair infarct healing and aggravate early remodeling which in turn causes cardiac rupture [38,39] Moreover, the increased cytokine gene expression from inflammation causes a secondary, self sustaining autocrine and paracrine GF an d cytokine expression [33] Therefore, reducing the sustained inflammatory response from MI and ischemia reperfusion injury will be beneficial for promoting optimal cardiac repair. Current CAD Treatment Treatment options for CAD involve lifestyle changes, pharmaceutical drugs, and surgical interventions [9] Healthy lifestyle changes can provide the simplest means to reversing this fatal disease. These changes include regular exercise, eating foods low in fat and cholesterol, and reducing stress [3] Various pharmaceutical drugs can be used to treat CAD by reducing blood pressure and slowing the heart rate [32] Cholesterol modifying medications reduce the amount of low density lipoprotein cholesterol in the blood which can

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12 reduce the plaque build up in the arteries [32] Invasive surgical treatment opti ons consist of angioplasty ( Figure 5 ) a nd coronary artery bypass graft surgery (CABG) [32] Figure 5 Diagram of Angioplasty [40] These surgical options aim to revascularize the ischemic tissue by compressing the plaque in the blocked artery with a catheter/balloon system (angioplasty) or by grafting a new coronary vessel to bypass the blocked artery through the CABG procedure The latter treatment requires open heart surgery and is generally only i mplemented in patients with multiple narrowed coronary arteries ( Figure 6 ) [32] Figure 6 Diagram of CABG surgery [41]

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13 However, there are a subset of patients that are unable to receive these surgical procedures due to various comorbidities [10,42] As a result, there has been a push towar ds alternative treatments that will revascularize the damaged heart tissue [3,10,43] Study Objective The overall clinical motivation for this research pertains to the treatme nt of CAD after a MI, through the route of therapeutic angiogenesis by long term and sequential delivery of multiple GFs T o achieve this type of GF delivery a biocompatible polymer gel with encapsulated nanoparticles was designed The RTG is sulfonated to protect VEGF from degradation and to slow its release over time. The nanoparticles will be loaded with PDGF and embedded within the SRTG, which will slowly release the particles over an extended period of time as the gel degrades. As this treatment is i ntended to help revascularize the heart after a MI, an anti inflammatory cytokine IL 10, will be incorporated into the SRTG as well. IL 10 will be mixed with the polymer system last, and should release first from the system to help reduce inflammation and reduce the injury effect of cardiac remodeling.

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14 CHAPTER II BACKGROUND Therapeutic Angiogenesis Using Growth Factors A potential alternative treatment to revascularizing is chemic tissue is therapeutic angiogenesis through GF delivery Angiogenesis is t he growth of the nascent blood vessels fr om existing vasculature networks The main mechanism behind the sprouting of new capillaries works throug h the release of G Fs such as VEGF and FGF 2 which signals for the migration and proliferation of endothelial cells to differentiate into guiding tip cells or proliferating stalk cells [17] Both of these cell types work together to form directional, elongating, new vessels. Later, a maturation process occurs involving pericytes which are stimulated by other GFs to cover the endothelial cells [44] The synergy of these cells through this process enables fully functional vessels to deliver blood flow to new or ischemic areas. Therapeutic angiogenesis aims to induce and control this angiogenic response in order to revascularize ischemic tissues. Therapeutic Angiogenesis Therapeutic angiogenesis has been extensively studied for the treatment of many human diseases. The generation of functional blood vessels from single GF administrat ion was pioneered by JM Inser when he injected VEGF165 in a rabbit hindlimb ischemia model [45] Since then, several other angiogenic G Fs have been used separately or together to promote angiogenesis. These GFs are not only utilized for diseases, but for other processes as well, such as wound healing an d organ repair and regeneration. As a result of dedicated research over the last two decades, many different types an d combinations of GFs have been elucidated for the purpose of inducing angiogenesis for MI treatment.

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15 Growth Factor Considerations Many different GFs have been extensively studied and applied for therapeutic angiogenesis purposes. These angiogenic GFs con sist of VEGF, bFGF, hepatocyte growth factor, PDGF, angiopoietin 1, transforming growth factor like growth factor. VEGF has shown to be the most significant controller of physiological angiogenesis during growth, healing and in response to hy poxia [13,46] VEGF provides angiogenic effects by binding to specific VEGF receptors on ECs, w hich leads to receptor dimerization and downstream signal transduction [12] When VEGF binding to ECs occurs, it induces their m igration, proliferation and formation of nascent or large vessels. Hypoxia, oncogenes, tumor suppressor genes, inflammatory cytokines and other GFs can all activate the production of VEGF and the ensuing angiogenesis response [47] However, when presented alone, VEGF can cause leaky and immature vessel formation with poor function and these vessels may regress quickly [1 7,48,49] Therefore, a second GF should be delivered to help stabilize and mature the neo vessel s PDGF has been shown to promote vascular maturation by recruiting SMCs and pericytes to newly formed vessels [12] This GF is produced by various cell types in response to external cues such as hypoxia or GF stimulation [50] PDGF also induces expression of VEGF which could provide a positive feedback loop to produce an increased angiogenesis response [51] Moreover, the PDGF signal regulates over 80 genes that are involved with matrix and cytoskeleton proteins, GFs, growth inhibitors, transcription factors involved in cell cycle regulation and others [9] In addition to stimulatin g pericytes to help mature newly formed vessels, it has also been shown that PDGF promotes vessel stabilization through stimulation of proliferation and migration of vascular ECs [52] Sequentially delivering

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16 VEGF and PDGF should provide an external stimulus to promote stable and mature vascularization for therapeutic angiogenesis applications. Anti inflammatory Factor Considerations To limi t the damaging effects of the immune response, delivery of anti inflammatory cytokines has been shown to optimize cardiac repair following an MI. IL 10 is a potent anti inflammatory cytokine which acts on monocytes by deactivating them and suppressing many pro inflammatory mediators [33] This cytokine has powerful inhibitory effects on monocytic cell lines and macrophages by inhibiting a panel of pro inflammatory cytokine m RNA expressions in both cell lines [53] It has been shown that IL 10 treatment in an acute MI mouse model significantly improved LV functions, reduced infarct size and attenuated infarct wall thinning [33] Furthermore, a study performed using a rat post MI heat failure model showed that therapy with IL 10 significantly improved LV function post MI and resulted in a reduced myocardial infiltration of m acrophages [54] Limiting the detrimental effects of pro inflammatory cytokines is critical for improving heart function after MI. Current Biological Factor Delivery Methods Therapeutic angiogenesis involves the delivery of GFs through proteins or genes encoding target proteins, or using stem cells to initiate the response [7] Because some approaches to therapeutic angiogenesis involve the use of viral vectors or cells, GF, protein delivery is the most simple and direct way [8] VEGF and bFGF are well studied GFs and have been tested in human clinical trials. The clinical studies used direct injection of free GF s but failed to demonstrate efficacy; this is likely due to the GF s being limited by their rapid diffusion rate, poor biostability, and short half lives in vivo [15] Therefore, controlled delivery systems for GF s is highly desirable and currently under extensive research [ 3,7,8]

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17 Conversely, there remains significant barriers to GF delivery systems related to loading capacity and long term efficacy [5] Polymeric Biomaterials for Protein Delivery Polymeric biomaterials may be a solution to providing sustained, local delivery of GFs for therapeutic angiogenesis [5] Certain biomaterials can overcome the short half lives and rapid diffusion of f ree GFs by releasing the proteins over the course of several weeks while stabilizing them in the polymeric network [4] Initial systems for therapeutic angiogenesis demonstrated the use of poly(lactic co glycolic acid) (PLGA) as a base polymer for controll ed release of VEGF [55] Due to its biodegradability and abili ty to control GF release, PLGA remains a favorable polymer even though it can cause undesirable inflammation from the increased acidic environment created after the polymer is degraded [8] Poly(ethylene glycol) (PEG) is a biocompatible polymer found in nature which has been utilized to create cell responsive hydrogels for VEGF release [56] Heparin, a naturally sulfated polysaccharide, has been shown to have binding affinity to many differen t proteins such as GF s, thus heparin plays an essential role in regulating various biological signaling processes [57] However, naturally derived material s generally experience batch to batch variance and have less tunable chemical properties in comparison to synthetic polymers [3] In order to achieve effective therapeutic angiogenesis, the proper combination of biocompatibility and GF encapsulation properties need to be incorporated into the polymer backbone to provide a susta ined release of these proteins.

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18 Sustained Growth Fact or Release The long term release of biological signals from a polymeric delivery system involves many complex physical and chemical interactions which need to be optimized to develop a consistent drug delivery vehicle. Most delivery systems have an initi the majority of loaded GF s are released in the first hours after injection [3] Although not fully unders tood, the mechanism behind the burst rel ease process relates to processing conditions surface characteristics, and geometry of the biomaterials used [58] Different strategies have been implemented to prevent this burst release by modulating the physical, biochemical affinity, and covalent binding interactions between the loaded GF s and polymeric material. Physical entrapment of the protein in polymeric sc affolds allows for controlled release of the biological signal based on the pore size of the polymer [5] In comparison to free VEGF injections in a myocardial ischemia or hindlimb ischemia model, controlled release of VEGF from PLGA scaffolds significantly improved angiogenesis [ 3] This improved angiogenesis, with controlled VEGF release, has also been demonstrated using PLGA microspheres containing alginate hydrogel [59] While physical entrapment of the GF in the polymer displays controlled release properties and improves angiogenesis, biochemical affinity interactions with heparin can be applied to further increase new blood vess el formation. Modulating Physical and Chemical Interactions Hydrogen and ionic bonding in teractions between GF s and heparin can be utilized to slow down protein release [59] bFGF preferentially binds to heparin sulfate and loading bFGF into PLGA nanospheres conjugated with heparin led to a linear release of the GF over

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19 a 15 day period [21] The positive charge on GF residues allow for preferential binding to the sulfon ate groups on heparin which are inherently negative ( Figure 7 ). Figure 7 ed carbon bound to sulfur, as well as the remaining chemical makeup of the polymer. Further prolonging of this release can be achieved by encapsulating these PLGA nanospheres into fibrin hydrogels [12] Once encapsulated in the hydrogel, sustained release of bFGF led to significant increases in capillary density and cell proliferation in a murine hindlimb ischemia model [12] Studies have also shown that the release of bFGF can be adjusted by varying the heparin concentration in the polymer [60] Similarly to bFGF, VEGF has positively charged domains of basic amino acid residues that bind to the sulfonate groups on heparin [61,62] Furthermore, some proteins can be covalently linked to polymeric scaffolds through a hydrolytically degradable or MMP degradable linker. It has been shown that the formation of a more controlled and stable vascula ture can be obtained using scaffolds that incorporate covalently attached VEGF [56] On the other hand, the chemical modifications utilized for covalent linkages can affect the biological activity of the attached GF [ 63] Developing a temperature responsive hydrogel, incorporated with heparin and encapsulating nanopa rticles containing GF s, should provide controlled and sustained protein releas e while improving angiogenesis. Thermally Induced Gelling Systems The cli nical feasibility of therapeutic angiogenesis methods can be improved by using polymeric materials that offer spatial and t emporal control of GF release [64] Temperature

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20 responsive hydrogels are one type of biomater ial that can provide this controlled release process. The incorporation o f poly(N isopropylacrylamide) (PNI PA m) into hydrogels has been heavily researched due to the polymer s reverse thermal gelling properties at different temperatures [65 68] This polymer exhibits a transition from a soluble liqui d solution to a solid gel once the lower critical solution temperature (LCST) threshold has been crossed ( Figure 8 ) [65] Figure 8 Polymer solubility behavior at the LCST. Left hand side shows hydrated polymer below LCST with entropic loss of water and chain collapse above LCST (right hand side) [69] P NIPAm based hydrogels have a sharp phase transition where the polymers remain a liquid at room temperature and rapidly become a solid gel when injected at body temperature (37 ), trapping any nanoparticles or proteins in the gel during this transition [64,70] In comparison with other cationic polymers, P NIPAm moieties for targete d delivery offers the advantages of good sensitivity, reversible transitions, and low cytotoxicity [30]. Additionally, protein release is easily tunable by manipulating the degradation properties of these polymers. By adjusting the hydrophilicity interacti ons of the polymer, and through alte ring the chemical structure of P NIPAm GF release can be varied through a change in the [70] In order to implement a further controlled and sequential

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21 release of factors, nanoparticles containing GF s can be entr apped in the hydrogel to provide additional angiogenic proteins a fter the gel has degraded Nanoparticles for Sequential Growth Factor Release Several studies have demonstrated the use of combining nanoparticle/hydrogel delivery systems for local and susta ined release of GF s for therapeutic angiogenesis [3,17,59] The study performed by Awada and colleagues (2015) demonstrated the efficacy and impo rtance of this dual delivery system to provide sequential release of GF s to improve revascularization and heart func tion after MI As discussed earlier, it is essential that the delivery system administers the GF s sequentially, while in their bioactive con formations to mimic their physiological mechanism during angiogenesis [3,17,18] If delivered simultaneously, early stage angiogenic factors can interfere with late stage factors which can cause immature blood vessel formation [18] These studies illustrate the importance of releasing early stage GF s from the hydrogel first, and then using another delivery vehicle (PLGA nanoparticles for example) to sequentially supply late stage GF s several days after the release of the first factor The optimal delivery system to improve revascularization and heart functi on after a MI will need to provide sequential, sustained, an d local release of G F s to the damaged heart tissue. Nanoparticle Considerations Many different compositions of nanoparticle (NP) delivery systems for GF s have been designed and fabricated from diverse types of synthetic and natural polymers. GFs can be loaded in the system before or after the fabrication of the particles by covalent or non covalent methods. As for non covalent means, GFs can be incorporated by adsorption, electrostatic interactions, or complexation [71] NP s are distinguished from other particulate

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22 systems by their smaller size, with the range of 1 1000 nm compared to > 1 m for microparticles [72] NPs offer advantages over microparticles because they can penetrate deeper into tissues through small capillaries and epithelial lining, allowing more efficient delivery of GFs to the target sites [73] When designing NPs for GF delivery, several issues must be considered: environmental factors that may denature or deactivate the NPs, physiochemical properties such as size distribution and surface charge, targeting ability of NPs for site specific deliv ery of GFs, and controllabl e GF release profiles to meet temporal and spatial demands [71] Successfully meeting these specific combinations of factors while designing NPs are likely to lead to more desirable therapeutic outcomes. There are several different types of NP systems that have been researched and elucidated for GF delivery such as lipid, polymer based and other miscellaneous systems ( Figure 9 ) Polymer based NPs in the form of nanospheres, nanocapsules, micelles and dendritic particles have all been broadly examined for GF delivery. Polymeric micelles are created from amphiphilic pol ymers that typically self assemble in an aqueous environment to form colloidal NPs [74] These micelles have a unique core shell structure which consists of an inner hydrophobic core and a hydrophilic exterior shell. Polymeric micelles offer advantages of smaller size and a uniform size distribution, in addition to high extravasating and tissue penetrating ability with reduced cytotoxicity [71] Studies have demonstrated high GF loading efficiency and sustained release of bFGF over a 2 mont h period without an initial burst release from polymeric micelles [75,76] Polymeric mice lles present significant advantages over other NP delivery systems which makes them suitable for in vivo GF delivery to promote therapeutic angiogenesis.

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23 Figure 9 Schematic representation of different NP systems [71]

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24 CHAPTER III PREVIOUS WORK Sulfonated PSHU PNIPA m (sulfonated PSHU PNIPAm will be used interchangeably with SRTG) has previously been designed and studied as a potential delivery vehicle for positively charged proteins. A synthesis method was established and the polymer has been characterized. Additionally, in vitro cytotoxicity testing, BSA release tests and CD spectroscopy were completed for the SRTG [22] SRTG LCST LCST was used to determ ine the sol gel phase transition temperature of sulfonated PSHU PNIPAm ( Figure 10 ) Figure 10 Temperature dependent phase transition of sulfonated PSHU PNIPAm A: The sulfonated PSHU PNIPAm undergoes a sharp, reversible phase transition around 32 as determined by UV Visible spectroscopy. B: An aqueous solution of sulfonated PSHU PNIPAm at room te mperature (C) turns to physical gel at 37 [22]

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25 An aqueous solution of sulfonated PSHU PNIPAm was heated slowly from 20 to 44 and percent transmittance was recorded during the process. The transmittance of the solution decreased slowly upon initial heating from 20 to 31 and then r apidly approached zero transmittance at 32 Upon further heating to over 33 the solution turned to an opaque solid indicating that the aqueous solution turns into a physical gel as the temperature increases. Thermodynamic competition between hydration of the polymer chains and the hydrophobic interactions between the polymer molecules causes this phase transition to occur [77] At low temperatures, or below the LCST which is 32 in this case, hydration of the polymer is thermodynamically favorable and the polymer molecules are maintained in a solution state ( Figure 10 B). When the temperature is above the LCST, hydrophobic interactions between the polymer and chains are favored and the polymer molecules interact to form a stand alone physical gel ( Figure 10 ) [22] Cytotoxicity An MTT assay with C2C12 myoblast cells was implemented to investigate in vitro cytotoxicity effects ( Figure 11 ). This method is well documented for measu ring cell viability and for supplying a general indication of cell health [55 57]. Figure 11 shows no statistical difference between the absorbance of pure medium and the polymer extract, while the addition of 10% DMSO significantly reduced the absorbance level. In this experiment, pure medium was used as a positive control and the DMSO was used a s a negative control. The differences in absorbance are directly related to differences in metabolic activity of the cel ls, which indicates the change in viable cells. Thus, the similar absorbance levels of the polymer extract and the pure medium samples i s evidence that the sulfonated PSHU PNIPAm is non

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26 cytotoxic [22] Plain PSHU PNIPAm was not included in this study as a previous study has proven that this polymer has good biocompatibility [78] Figure 11 In vitro cytotoxicity of sulfonated PSHU PNIPAm by MTT assay. Results demonstrated no cytoto xic effects of sulfonated PSHU PNIPAm on C2C12 cells after exposure to the polymer extract in medium, while 10% DMSO shows significant cytotoxicity. There is no statistical difference between medium and extracts. Data represent mean SD. indicates p value < t test) [22] B SA Release Test A BSA release test was implemented to examine the protein release rate properties of PSHU PNIPAm and sulfonated PSHU PNIPAm at varying wt. % concentrations. As this polymer system is intended to be a protein delivery vehicle for positively charged molecule s, BSA was used for this study due to its inherent positive charge [22] The sulfonated PSHU PNIPAm showed an increased sustained release profile in comparison to plain PSHU PNIPAm ( Figur e 12 ).

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27 Figur e 12 BSA release profiles from sulfonated PSHU PNIPAm and plain PSHU PNIPAm : sulfonated PSHU PNIPAm 10%, PSHU PNIPAm PSHU PNIPAm plain PSHU PNIPAm 15%. BSA release was more sustained from sulfonated PSHU PNIPAm than that from plai n PSHU PNIPAm with a more sus tained profile at higher conc entrations. Data represent mean SD [22] This showed that the negatively charged sulfonated groups on the sulfonated PSHU PNIPAm may effectively hol d BSA in the polymer matrix. Furthermore, the BSA release profile showed more sustained release at higher polymer concentrations which supports that the release rate can be readily modified by changing the sulfonated PSHU PNIPAm concentration [22] A possible explanation is that higher concen trations of the sulfonated PSH U PNIPAm leads to an increased sulfonate group density in a designated volume of the polymer gel, which results in more negative overall charge This causes an enhanced electrostatic interaction between sulfonated PSHU PNIPAm and BSA.

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28 CD Spectroscopy If the sulfonated PSHU PNIPAm will serve as a protein delivery vehicle, then it should not affect protein structure n or lead to the denaturation of the proteins. To demonstrate this, the secondary structure of the released BSA f rom the final day of the release study was analyzed by CD spectroscopy and compared with natural BSA solution. No significant difference was observed in all samples ( Figure 13 helix conformation [79 81] which confirms that the protein structure is well preserved [22] Figure 13 CD spectra of native and released BSA. The similarity of the protein conformation betw een samples indicates that the secondary structure of the BSA protein was well preserved by the system [22]

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29 CHAPTER IV HYPOTHESIS AND SPECI FIC AIMS Hypothesis Based on preliminary data it was hypothesized that the sulfonated thermal gel encapsula ting micelle nanoparticles, while sequentially releas ing biological factors, will induce vascularization and reduce an inflammatory response. Specific Aims The first specific aim was to synthesize and characterize PE G PSHU PEG micelle NPs As the micelles will be utilized to provide sequ ential delivery of PDGF from the polymer scaffold, it is important to confirm the overall molecular structure of the micelles and that they are uniformly sized on the nanometer scale. The second specific aim was to ensure that the biological factors were b eing delivered sequentially and for an extended amount of time from the delivery system in vitro To assess this, a release study was performed with the polymer delivery system and samples were taken daily. The amount of p rotein released from the samples w as quantified using specific ELISAs for each factor. The third specific aim was to demonstrate a substantial angiogenic response with a reduced immune response in vivo This was evaluated by injecting the polymer system into the subcutaneous region of the lower backs of mice. Angiogenesis and immune response were observed using immunohistochemistry (IHC) with stains specific for : endothelial cells as a marker for early blood vessel development, von W illebrand factor (vWF) to show functional smoot h muscle actin SMA) to identity the recruitment of smooth muscle cells which demonstrates mature blood vessel formation and macrophages to determine the reduction in the inflammatory response.

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30 CHAPTER V MATERIALS AND METHOD S Materials N isopropylacrylamide ( NIPAm ) was purchased from Tokyo Chemical Industry (Chuo ku, Tokyo, Japan). Anhydrous N ,N dimethylformamide (DMF) was purchased from EMD Millipore (Billerica MA, USA ). 4,4' a zobis(4 cyanovaleric acid) (ACA) N hydroxysuccinimide (NHS ), sodium bicarbonate, 1 bromohexane, TWEEN 20 Triton TM X 100 globulins from bovine blood urea, he xamethylene diisocyante (HDI), acetic acid, sodium acetate, 1,3 p ropane sultone (PS) trifluoroacetic acid (TFA), triethylamine, dimethyl sulfoxide (DMS O) Azino bis(3 ethylbenzothiazoline 6 sulfonic acid) (A BTS) Liquid Substrate System, dichloromethane (DCM) and bovine serum albumin (BSA) were purchased from Sigma Aldrich (St. Louis, MO, USA) Di tert butyl dicarbonate, ethyl acetate, N hydroxysuccinimide (NHS), N (3 dimethylamino propyl) ethylcarbodiimide hydrochloride ( EDC), 2 Amino 1,2 propanediol, 98% (Serinol) anhydrous methanol dimethyl sulfoxide d6 (DMSO d6), polyethylene glycol 1000, ethanol, hexane and a nhydrous diethyl eth er were purchased from Fisher Scientific (Pittsburgh, PA, USA). Goat anti Rabbi t IgG (H+L) Secondary Antibody Alexa Fluor 594 Goat anti Rat IgG (H+L) Secondary Antibody Alexa Fluor 488 Rabbit anti Goat IgG (H+L) Secondary Anti body Alexa Fluor 594, CD31 Antibody (Rat IgG2a) and Smooth Muscle Actin Antibody SMA, rabbit IgG) were purchased form Thermo Fisher Scientific (Waltham, MA, USA). Anti Von Willebrand Factor antibody (vWF, sheep IgG) and Anti CD68 antibody (Rabbit IgG) were purchased from Abc am (Cambridge, Ma). Human PDGF BB Standard ABTS ELISA Development Kit Murine IL 10 Standard ABTS ELISA Development Kit Murine VEGF Standard ABTS ELISA Development Kit Recombinant Human PDGF BB Recombinant Murine IL 10 and

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31 Recombinant Murine VEGF165 wer e purchased from Peprotech (Rocky Hill, NJ, USA). S aline, and isoflurane were purchased from MWI Veterinary Supply (Boise, ID, USA). 10 % formalin was purchased from JT Baker (Phillipsburg, NJ, USA). Sucrose (RNASE & DNASE free) was purchased from VWR Life Science (Radnor, PA, USA). Optimal cutting temperature (OCT) compound was purchased from Sakura (Torrance, CA, USA). Phosphate buffered s aline (PBS) was purchased from HyC lone Laboratories, Inc. ( South Logan Utah, USA ) Alexa Fluor 594 (goat anti rabbit IgG) was purchased from Life Technologies (Carlsbad, CA, USA). Dapi flouro mount G was purchased from Electron Microscope Sciences (Hartfield, PA USA ). Spectra/Por dialysis membranes (MWCO: 3500 5000 and 12,000 14,000 Da) were purchased from Spectrum Laboratories (Rancho Dominguez, CA). Equipment Proton nuclear magnetic resonance ( 1 H NMR ) was performed on a Varian Inova 500 NMR Spectrometer and samples were run in DMSO d6 at room temperature. Fourier tran sform infrared spectroscopy (FT IR) was performe d on a Nicolet 6700 FTIR Spectrometer and samples were run on polyethy lene infrared (IR) sample cards Polymer morphology was imaged using a JEOL (Peabody, MA) JSAM 6010la analytical scanning electron microscope. The low critical solution temperature (LCST ) was determined using a Cary 100 Bio UV Visible spectrophotome ter with a temperature controlled cell holder. Nanoparticle size was measured using a Zetasizer Nano ZS ( Malvern Instruments Ltd Worcestershire, UK ). Elemental analysis was per formed by MicroA nalysis, Inc.(Wilmington, DE) ELISA color development was monitored with an ELISA plate reader (BioTek Synergy 2 Multi Mode Reader) at 405 nm with wavelength correction set at 650 nm. Tissue was sectioned using a

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32 CryoStar NX70 Cryostat. Con focal images we re taken using a Zeiss LSM 780 spectral microscope ImageJ was used to quantify variables for image analysis. Polymer Synthesis N BOC serinol synthesis N Boc serinol was synthesized through Boc group protecti on of serinol. Serinol (1.959 g ) was dissolved in 20mL of absolute ethanol and stirred at 4 Di t ert butyl dicarbonate (5.973 mL ) was dissolved in 20 mL of absolute ethanol and added dropwise to the serinol solution over a period of one hour, while maintaining 4 and constant stirring The solution was warmed to 37 with vigorous stirring and reacted for one hour. The ethanol was removed by rotary evaporation at 45 and 10 mbar vacuum and the solid was re dissolved in a 1:1 mixture of hexane and ethyl acetate by gentle heating. Additional hexane was added until precipitation was observed and the resul ting suspension was stored at 4 overnight to allow recrystallization. Subsequent vacuum filtration yielded a white flaky product Poly (serinol hexamethylene urea) or PSHU backbone synthesis PSHU is the backbone of the polymer system used in this study. N BOC serinol (1.149 g, 6 mmol) and urea (0.36 g, 6 mmol) were lyophilized for 24 hours and then dissolved in 6 mL of anhydrous DMF in a 25 mL round bottom flask under a n itrogen atmosphere. HDI (1.928 mL 12 mmol) was added dropwise to the flask and the polymerization was performed for seven days at 90 under a nitrogen atmosphere After cooling down to ambient temperature, the mixture was precipitated twice into excess cool anhydrous diethyl ether an d twice in distilled water for purification T he product was subsequently lyophilized at 45 for 24 h and recovered for further conjugation.

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33 PSHU deprotection In order to conjugate other polymers/chemicals to PSHU, it was deprotected by removing the te rt Butyloxycarbonyl (BOC) groups and exposing the primary amine groups through the following process. PSHU (1.00 g) was dissolved in 10 mL of a TFA/DCM (1:1, v/v) in a round bottom flask. The solution was left gently stirring with no heat at room temperatu re for 30 minutes. Rotary evaporation was used to remove solvents at 45 C and 10 mbar. The remaining contents were re dissolved in DMF (1 mL), and diethyl ether was added to the flask to precipitate the polymer out of solution. Excess ether was poured off while taking care to keep polymer within the flask. R otary evaporation was then used to remove residual ether and the deprotected PSHU (dPSHU) was purified by two more precipitations into diethyl ether Finally, the polymer was decanted in water and lyoph ilized at 45 C for 24 h At the end of this process, the polymer should look dry and white. P NIPAm Synthesis P NIPAm was conjugated using radical polymerization with an azobis initiator NIPAm (5.0 g, 44.2 mmol) and ACA (0.062 g ) were dissolved in anhydrous methanol (25 mL), and the mixture was purged with nitrogen gas for 30 minutes at room temperature. A reflux condenser apparatus was set up, and the solution was stirred for three hours at 68 C under a nitrogen atmosphere Nex t, the solution was precipitated into milli Q water at 60 C in a dropwise manner. Following precipitation, the warm water was discarded and 40 mL of cold Milli Q water was added to the precipitate. The precipitate was allowed to dissolve into the water in a cold room at 4 C. The resulting solution was dialyzed using 3500 kDa MWCO dialysis tubing against 1 L of milli Q water for 48 h The product was lyophilized at 45 C for 48 h.

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34 PSHU PNIP Am Conjugation (RTG Synthesis) P NIPAm was conjugated to 25% of the free amine groups on dPSHU. EDC/NHS chemistry was used to achieve this conjugation P NIPAm (0.75 g) was dis solved in 5 mL of anhydrous DMF wi th three molar excess of EDC and NHS in a 25 mL round bottom flask at room temperature under a nitrogen atmosphere and the mixture was stirred for 24 h. One milliliter of dPSHU solution (0.75 g/mL) prepared in anhydrous DMF was added slowly to the P NIPAm solution flask and the reaction was performed for 24 h at room temperature under a nitrogen atmosphere. The result ing solution was capped (no vent) and left stirring for 24 hours to complete the reaction. The polymer was precipitated three times in diethyl ether, and the solvent was removed each time by rotary evaporation. The dried polymer was dissolved in milliQ wat er at 4C and dialyzed (MWCO: 12 14 K Da) against 1 L milliQ water for 48 h at room temp The product was lyophilized at 45 C for 48 hours. Sulfonation of PSHU P NIPAm (SRTG Synthesis) As P NIPAm was only conjugated to 25% of the available primary amine gr oups, the remaining amine groups were left for the attachment of sulfonate groups. For this sulfonation reaction, PS (0.034 g, 5 mmol) and t BuOK (0.032 g, 5 mmol) were dissolved in 3 mL of anhy drous DMF in a 25 mL round bottom flask at 50 C under a nitro gen atmosphere. A solution of 3 mL PSHU PNIPAm (0.1 g/mL) in anhydrous DMF was added slowly to the flask and the sulfonation reaction was performed for 3 days at 60 C under a nitrogen atmosphere. After cooling down to ambient temperature, the mixture was precipitated into excess diethyl ether three times Finally, the polymer was dissolved in milliQ water and dialyzed (MWCO: 12 14 KDa) against 1 L of milliQ water for 48 h at room temperature and lyophilized at 45 C for 48 h.

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35 PEG PSHU PEG Micelle Nanoparticle Polymer Synthesis N Boc serinol (0.2873 g, 1.5 mmol) and urea (0.09 g, 1.5 mmol) were weighed out and lyophilized at 45 C for 48 h The reactants were dissolved in 1.5 mL of anhydrous DMF in a 25 mL round bottom flask at 90C under gentle st irring and a nitrogen atmosphere. Hexamethylene diisocyanate (HDI) (0.482 mL, 3 mmol) was added drop wise, and the polymerization was carried out for 5 days. After the specified time, an excess of polyethylene glycol (PEG 1000, 4 mmol) was dehydrated and a dded to the reaction. The PEGylation reaction was carried out for 24 h at 90C. The resulting product, poly(ethylene glycol) block poly(serinol hexamethylene urea) block poly(ethylene glycol) (PEG PSHU PEG), was purified by three precipitations in diethyl ether and then dried completely by extended rotary evaporation at 50C and 10 mbar vacuum. Subsequently, t he polymer was dissolved in mil liQ water and dialyzed (MWCO: 3.5 k Da) against 1 L of milliQ water for 72 h at room temperature. Then the product was lyophilized 45 C for 48 h to yield a white flaky material. Micelle Fabrication GF loaded micelles were fabricated by a traditional emulsification sonication procedure. The PEG PSHU PEG polymer and GF were dissolved in 1 mL DMSO at 1 wt% (polymer/DMSO). This solution was then added drop wise to a beaker containing 20 mL of milliQ water partially submerged in an ultrasonic bath. The resulting emulsion was sonicated for 10 min. Removal of DMSO was carrie d out by centrifugation at 11,000 revolutions per minute (rpm) for 5 min, pouring off the supernatant and then re suspending the micelles in milliQ water. This DMSO extraction procedure was carried out 3 times. The resulting micelles were either used immed iately or stored at 20C for later use.

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36 Polymer Characterization Elemental Analysis To determine the amount of sulfonation on sulfonated PSHU PNIPAm the polymer was sent to Micro Analysis, Inc. (Wilmington, DE) for elemental analysis testing. Carbon, hydrogen and nitrogen (CHN) a nalysis was used to determine the percent, by weight, of carbon, hydrogen, and n itrogen contained in the polymer The company has a separate testing method for sulfur content using an Antek analysis machine. The Antek provides the determination of total nitrogen and total sulfur in organic materials, both solids and liquids at trace levels. Sample material is combusted in combination with oxygen, converting chemically bound nitrogen to NO and sulfur to SO2. The combustion gases are routed through a membrane drying system to remove all water and then to the detector modules for quantification. Nitrogen is detected by way of chemiluminescence and sulfur by ultraviolet f luorescence. Dynamic Light Scattering (DLS) DLS is a technique that can be used to determine the size distribution profile of different particles, or polymers, in solution. The Zetasizer Nano ZS uses this technique to measure the diffusion of particles moving under Brownian motion, and converts this to size and a size distribution using the Stoke Einstein relationship. Non Invasive Back Scatter technology (NIBS) is incorporated into this instrument to give the highest sensitivity simultaneously with the highest size and concentration range. Scanning Electron Microscopy (SEM) SEM was used to examine polymer structure and morphology on the nanometer scale. This technique was used show th e morphology of the NP s alone and encapsulated within t he

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37 SRTG. For particle morphology assessment, dry micelles wer e sputter coated with 5 nm Au and examined by SEM. Proton Nuclear Magnetic Resonance ( 1 H NMR) Proton nuclear magnetic resonance ( 1 H NMR) was completed with a Var ian Inova 500 NMR Spectrometer. 1 H NMR was used to confirm the structure of PSHU, the struct ure of PEG PSHU PEG, the removal of BOC protecting groups on dPSHU and the successful conjugation of PNIPAm to dPSHU. The PSHU, dPSHU, PSHU PNIP Am and PEG PSHU PEG samples (3 5 mg) to be analyzed were dissolved in 600 L of DMSO d 6 Spectra were process ed and analyzed using ACD 1D NMR Processor software (Advanced Chemistry Development, Inc.). Fourier Transform Infrared (FTIR) Spectroscopy Sulfonated PSHU PNIPAm and PSHU PNIPAm samples were evaluated using FTIR spectroscopy. Samples were dissolved in te trahydrofuran (THF) and placed on polyethylene windowed cards which were then analyzed by a Nicolet 6700 (Thermo Fisher Scientific). In Vitro Biological Factor Release Study 5 wt% polymeric (RTG + VEGF + NPs + PDGF + IL 10 and SRTG + VEGF + NPs + PDGF + IL 10 ) solutions were created using PBS with 0.2% BSA Solutions were mixed and left to dis solve overnight at 4 C. 500 ng of each factor was added to the gels at different times. VEGF was added directly to the polymer solutions 2 hours before the gelling was to occur. The micelle NPs loaded with PDGF were added 30 minutes before gelling and IL 10 was added 10 minutes before gelation. vials ( Figure 14 ) in a 37 C incubator for 2 minutes and then 1 mL of warm (37 C) release solution (1x PBS with 0.2% BSA) was added to the vials After 2 minutes, 1 m L of release solution was removed from each sample for the day 0 timepoint and 1 mL of fre sh warm release solution was

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38 added. Samples were then taken every 24 hours and immediately placed in a 80 C freezer for further analysis. Figure 14 Formed gels in 2 mL vials for in vitro release study. Enzyme Linked Immunosorbent Assay (ELISA) ELISA is a laboratory technique for quantifying the concentration of specific antigens or antibodies in various solutions. A sandwich ELISA format was used to quantify the concentration of factors in the release solution samples from the in vitro polymer delivery systems. The protocol from the manufacture was followed for each ELISA kit. First, the primary antibodies are attached to the wells. Once the antibodies have attached overnight, the samples are added to the wells and the protein of interest attaches to the primary antibody. A secondary antibody is then added, which will attach to the protein of interest. Next, an enzyme that attaches to the secondary antibody is added. Subsequently, a liq uid substrate is added that intera cts with the enzyme which produces a color change. The color development is detected using a spectrophotometer set to 405 nm with wavelength correction set at 650 nm. The plate is monitored at 5 minute intervals for approximately 20 minutes. These measurem ents correspond to the concentration of the protein of interest in the initial sample

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39 solution. To quantify the amount of protein from the absorbance readings for each well, a standard curve is created with each ELISA plate to compare values to the unknown concentrations Additionally, the plate is washed several times between each step of the process to remove any residual materials not attached to the bottom of the wells. In Vivo Mouse Subcutaneous Biological Factor Injections The immune response and ang iogenesis analysis procedure was approved by the Institutional Animal Care and Use Committee (IACUC). A total of 36 adult male C57BL/6 mice were used for the study, 3 mice per ea ch injection group ( saline, SRTG, SRTG + VEGF, SRTG + VEGF + PDGF, SRTG + VEGF + NPs + PDGF or SRTG + VEGF + NPs + PDGF + IL 10) for 2 time points (7, 21 days ). The mice were allowed 7 d to acclimate prior to injections and maintained on a 14/10 hour light/dark cycle with access to food and water ad libitum. The mice were anaesthetized using continuous isoflurane and oxygen inhalation. Initial induction was at 5 % isoflurane in oxygen and then maintained at 2 % isoflurane in oxygen. Under anesthesia SRTG + VEGF + PDGF, SRTG + VEGF + NPs + PDGF or SRTG + VEGF + NPs + PDGF + IL 10 was subcutaneously injected in the middle back of the mice through a 25G insulin needle ( Figure 15 ) The amoun t of protein injected for each factor was 250 ng. The injection bolus was clearly noticeable under the subcutaneous tissue.

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40 Figure 15 Demonstration of subcutaneous injection in the middle back of a mouse [82] Euthanasia and T issue H arvest 7 or 21 days after injection the mice were euthanized by carbon d ioxide and cervical dislocation. After removing the hair around the injection site with depilatory cream, t he subcutaneous tissue (2 cm 2 cm) was harvested at the site of injection. The site was clearly visible as a result of shaving the backs of the mice at the time of the initial injection. Histology The s ubcutaneous tissue was fixed using 4% Paraformaldehyde in PBS for 24 h, cryoprotected with 30% sucrose in PBS for 24 h embedded in OCT compound, and frozen at 80 C. The tissue was sectioned transversely with a thickness of 5 m and placed on glass slide s. The sections were fixed in acetone for 10 min and washed 3 times in Wash Buffer (1X PBS, 0.1% Tween 20) for 5 min each The sections were permeabilized in P ermeabilizing Buffer (1X PBS, 0.5% Triton X 100) for 10 min and washed 3 times in Wash Buffer for 5 min each. The sections were blocked in Blocking/Dilution Buffer (1X PBS 0.25% Triton X 100, 2% BSA, 4% Bovine Gamma Globulins) for 60 min and washed 3 times in Wash Buffer for 5 min each. All antibodies were diluted in Blocking/ Dilution Buffer The sec tions were

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41 stained with primary antibodies (single or double stains) overnight washed 3 times in Wash Buffer for 5 min each, stained with the designated secondary antibody for 60 min, washed 3 times in Wash Buffer for 5 min each and washed 3 times in mill iQ water for 5 min each Dapi flouromount G was used to mount the slides. The sections we re double immunostained with either CD31 (1:50 ) SMA (1:250 ), or CD31 (1:50) and vWF (1:5 0). The sections were single immunostained with CD68 (1:500). For the secondary antibodies, CD31 sections were stained with Alexa Fluor 488 (1:500) SMA sections were stained with Alexa Flour 594 (1:500), vWF sections were stained with Alexa Flour 594 (1 :500) and CD68 sections were stained with Alexa Flour 594 (1:500). Images were taken using confocal microscopy. Quantitative Analysis of Immunofluorescent Staining 5 low magnification (200x) fields containing the highest number of CD31 SMA vWF or C D68 positive cells were selected for each group following previously published criteria [8] The number of CD31 SMA vWF or CD68 positive cells in the field was counted and confirmed by DAPI positive nuclei Statistical A nalysis A ll results are reported as means standard error of the mean. Analysis of variance (ANOVA) was utilized to determine significant differences between groups and followed by a two tailed t test test when applicable. Statistical significance (*) was indicate d when p < 0.05.

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42 CHAPTER VI RESULTS AND DISCUSSI ON Synthesis of S ulfonated PSHU P NIPAm Reaction S equence The synthesized sulfonated PSHU PNIPAm was designed to mimic heparin and hold similar biofunction to natural heparin, while having the advantage of being a reverse thermal gel with consistent reproducibility. Moreover it contains amide and ester groups to provide sites for degradation and the hydrophobic a lkyl chains allow for decreased degradation rates and allows for improved cell attachment. The sulfonated PSHU PNIPAm was synthesized using PSHU PNIPAm copolymer ( Figure 16 ). Figure 16 Synthesis Reaction sequence of sulfonated PSHU PNIPAm synthesis [22] The primary amine in PSHU is normally protected by a BOC group but it can easily be deprotected to a free amine group in the mixture of TFA/DCM. Then these deprotected

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43 group s can be utilized for further conjugations. PNIPAm was conjugated to 25% of the available primary amine groups and the remaining primary amines were then available for the attachment of sulfonate groups. Synthesis of PEG PSHU PEG Reaction S equence The no vel synthesized PEG PSHU PEG was based on the same PSHU backbone as sulfonated PSHU PNIPAm which makes the polymer biocompatible and biodegradable. As this polymer is used t o fabricate micelle NP s, it must have differing sections of hydrophobic and hydrop hilic chains to form these particles. The PSHU backbone provides the hydrophobic core of the micelles, while the terminal PEG chains provide the hydrophilic interactions on the exterior shell which can ionically bind to GF s ( Figure 17 ). PEG adds an NPs for an extended half life in circulation. When the micelles are injected together with sulfonated PSHU PNIP Am, the micelles provide long term and sequential release of PDGF while the SRTG provides a scaffold for their retention at the injection site. Figure 17 Reaction sequence of PEG PSHU PEG synthesis.

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44 PSHU and dPSHU C haracterization 1 H NMR The synthesis of PSHU from N BOC serinol, urea, and HDI was confirmed using 1 H NMR ( Figure 18 ). Figure 18 1 H NMR sp ectrum of PSHU confirming the molecular struct ure [83] 1 H NMR was used to confirm the removal of BOC protecting groups. The resulting free amines were used to conjugate P NIPAm and sulfonate groups. The loss of the b peak confirms the remov al of the BOC groups ( Figure 19 ). Figure 19 1 H NMR spectrum of PSHU and dPSHU confirming the removal of t he BOC protecting group with the loss of the b peak [83]

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45 PSHU P NIPAm and PEG PSHU PEG Characterization U sing 1 H NMR The synthesis of PSHU PNIPAm from N BOC serinol, urea, HDI and PNIPAm was confirmed using 1 H NMR ( Figure 20 ). Figure 20 1 H NMR spectrum of PSHU PNIPAm Successful conjugation of PNIP Am was confirmed by the presence of methylene and methyl protons at 1.55 and 1.09 ppm, respectively. The synthesis of PEG PSHU PEG from N BOC serinol, urea, HDI and PEG was confirmed using 1 H NMR ( Figure 21 ). Figure 21 1 H NMR spectra of PEG PSHU PEG. Successful conjugation of PEG was confirmed by the presence o f the peak at 3.51 ppm which identifies the protons on the PEG repeating unit of the polymer backbone [84]

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46 PSHU PNIPAm and S ulfonated PSHU PNIAPAm Characterization U sing FTIR In addition to 1 H NMR FTIR was used to further characterize the chemical structure of PSHU PNIPAm and s ulfonated PSHU PNIAPAm. This method was used to confirm the conjugation of the sulfonate groups to s ulfonated PSHU PNIAPAm ( Figure 22 ) Figure 22 FTIR spectra of sulfona ted PSHU P NIPAm and PSHU P NIPAm FTIR can identify the bond between a single sulfur molecule and oxygen on the sulfonate groups. PSHU PNIPAm was used as a baseline spectrum to compare with the sulfonated PSHU PNIPAm to identify the sulfonate peak. Thus, peaks appearing f rom the sulfonated polymer spectrum represent changes in bonds due to the sulfonation process. Figure 23 highlights a peak area between 1050 1025 cm 1 that has been identified as a sulfonate peak, which confirms tha t the SRTG has been sulfonated as this shift is not present in the plain RTG.

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47 Figure 23 FTIR spectra of sulfona ted PSHU PNIPA m and PSHU PNIPA m with an enlargement of the sulfona te peak. SRTG Sulfonation Detection Using E leme ntal A nalysis After confirmation of sulfonation with FTIR, elemental analysis was performed to further verify the sulfona tion of the SRTG, while also determining the amount of sulfonation. Elemental analysis is a process where a sample of a compound is analyzed to determine its elemental composition for carbon, hydrogen, nitrogen, oxygen and sulfur (Micro Analysis Inc.). The sample material is combusted and the gases produced are routed through a membrane drying system to remove all water and then the detector modules quantify the chemicals in the compound. Table 1 shows the results of the elemental analysis for the RTG and SRTG. The RTG was used as a baseline for how much sulfur is in the polymer without being sulfonated. The results show that 0.11% wt of the SRTG contains sulfur groups which further verifies that the polymer is sulfona ted when compared to the plain RTG with 0.02% wt sulfur content.

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48 Table 1 Elemental analysis of PSHU PNIPAm (RTG) and sulf onated PSHU PNIPAm (SRTG) Polymers were analyzed for car bon, hydrogen, nitrogen, oxygen and sulfur content and were determined by weight percent for each element. Polymer C [%] H [%] N [%] O [%] S [%] SRTG 57.09 10.07 11.86 16.65 0.11 RTG 56.44 9.95 12.09 16.15 0.02 PEG PSHU PEG Micelle Nanoparticles Size Distribution U sing DLS DLS is a technique that can be used to determine the size distribution profile of different particles, or polymers, in solution. The micelle size was determined using a Zetasizer Nano ZS. Three separate batches of micelles were made using the same protocol and the particle size was measured ( Figure 24 ). T he DLS data are plotted using an intensity weighted distribution. Micelles had a mean diameter of 216.5 nm with a relative standard deviation (RSD) of 8.62 %. As shown with a low RSD value, the micelles consist of a monodispersed population of particles an d batch to batch variability appears to be low. Figure 24 Size distribution of micelles from DLS measurements.

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49 Morphology of S RTG Embedded with N P s U sing SEM Scanning electron microscopy (SEM) was used to characterize the morphologies of the SRTG, micelles and the combined delivery system. SEM revealed that the SRTG al one consists of polymer sheets ( Figure 25 A), while the micelle NPs were uniformly spherical with the expected core shell structure ( Figure 25 B). Figure 25 SEM images of the SRTG and micelle NPs A: 5 % (w /v) of SRTG cross section showing polymer sheets (scale bar = 10 m). B: Micelle nanoparticles confirming the spherical structure (scale bar = 2 m). C: SRTG encapsulating micelles cross section showing porous configuration (scale bar = 20 m). D: Enlargem ent of SRTG encapsulating micelles with black arrows indicating micelles (scale bar = 1 m). Figure 25C shows the morphology of the SRTG encapsulating the micelle NPs The morphology of the polymeric gel embedded with micelles is substantially altered from the original polymer sheets, as the cross sectional image sho ws the gel in a highly porous configuration Figure 25D is an enlargement of the porous area of the same polymeric gel

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50 showing that the micelles are embedded within the gel and their uniform spherical shape has been maintained. The black arrows indicate the micelles in the gel. I n Vitro Multiple Biological Factor s Release Study An in vitro release test study was performed to compare the cumulative release rates of the RTG to the SRTG wit h all three biological factors and the micelles loaded into the gels. 5 wt% polymeric (RTG + VEGF + NPs + PDGF + IL 10 and SRTG + VEGF + NPs + PDGF + IL 10) solutions were created using PBS with 0.2% BSA 500 ng of each factor was adde d to the gels and the samples were incubated at 37 with samples taken daily. The samples were then analyzed using ELISAs to quantify the cumulative release for each protein ( Figure 26 ). Figure 26 C umulative release profile showing the sequential release of all three factors from the SRTG and plain RTG (n=3 samples) The release profiles for the RTG and SRTG show that the sulfonate groups are reducing the a mount of protein released for each factor. The SRTG shows a 12.5 % reduction

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51 in the burst release of VEGF compared to the plain RTG for the day 0 timepoint. This confirms that the sulfonate groups on the SRTG are mimicking heparin sulfate function and elec trostatically bind to VEGF to reduce its release rate. Furthermore, the SRTG slowed the release rate for IL 10 and PDGF slightly. Heparin sulfate binds to areas rich in basic amino acid residues such as arginine and lysine, and it has been shown that IL 10 and PDGF contain these regions and bind to the sulfonate groups on heparin [85,86] Additionally, both polymer systems displayed sequential release of all three factors, VEGF released first, f ollowed by IL 10 and then PDGF released last. This demonstration of sequentially releasing factors f ro m the polymer scaffold is essential because if PDGF released first, it could interfere with the therapeutic angiogenesis process during the later stage of vessel maturation. If the blood vessels do not mature properly, then this could cause leaky and non functional vessels. The factors also showed sustained release from the delivery system as the samples showed a consistent release rate of about 0.25 0.5 ng per day through 17 days after the initial burst release. The GFs need to have this sustained rel ease to help stabilize the neo vessels during their initial formation through their final development into functional blood vessels. Demonstrating sustained release of IL 10 is also important to reduce an inflammatory response from the polymer system itsel f, and to limit the amount of inflammation from an MI to minimize the damaging effects of cardiac remodeling. In Vivo Mouse Subcutaneous Biological Factor Injections In order t o assess the angiogenesis and immune response efficacy in vivo the SRTG polymer delivery system was subcutaneously injected in the middle back of mice. The amount of protein injected for each factor was 250 ng and the following 6 groups were injected: saline, SRTG, SRTG + VEGF, SRTG + VEGF + PDGF, SRTG + VEGF + NPs +

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52 PDGF or SRTG + VEGF + NPs + PD GF + IL 10) for 2 time points (7 days, 21 days ). Immunohistochemistry (IHC) was performed to directly identify functional and mature blood vessels and to identify infiltrating macrophages to quantify a change in the immune response. IHC Analysis of Therapeutic Angiogenesis Response IHC was used to identify and quantify the different cell types involved in the process of blood vessel formation. Vascular endothelial cells (ECs) were stained with CD31 and Alexa Flour 488, vascular smoot h muscle cells were stained SMA and Alexa Flour 594, and ECs associated with the functionality of blood vessels were stained wi th vWF and Alexa Flour 594. Saline was used as a negative control to compare how many blood vessels normally exist in the subcu taneous tissue and the number of cells (SMCs, ECs, vWF+ cells) that constitute those vessels. Positively stained cells were counted and quantified when DAPI was identified with the stains. Functional Vascular Endothelial Cell Analysis Co staining for CD31 and vWF, a factor involved in hemostasis, was examined to identify the amount of functional vascular endothe lial cells within blood vessels for 7 and 21 days after injection ( Figure 27 ). In all experimental injection groups, an increase in functional blood vessels was observed qualitatively by comparing organized rings of ECs (CD31+ and v WF+ cells) to saline injections (Figure 27).

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53 Figure 27

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54 Figure 27 Representative images of IHC co staining for ECs (CD31) and blood vessel functionality (vWF) after 7 and 21 days for all 6 groups Subcutaneous tissue was stained with: CD31 and Alexa Flour 488 and appear green vWF and Alexa Flour 594 and appear red and DAPI which appears blue. Scale bar represents 100 m. After 21 days, SRT G + VEGF + PDGF, SRTG + VEGF + NPs + PDGF and SRTG + VEGF + NPs + PDGF + IL 10 groups revealed a higher quantity of functional ECs compared to the day 7 time point demonstrating that the cells are maturing over time (Figure 27) Staining for CD31 was firs t exa mined to compare the proliferation of E Cs associated with blood vessel lumen formation After 7 days all 5 experimental SRTG groups showed significantly more ECs ( CD31 + cells) than saline ( Figure 28 ) After 21 days all 4 SRTG groups loaded with biological factors showed significantly more ECs than saline ( Figure 28 ) The SRTG group without any factors showing significantly more ECs at 7 days may be attributed to an immune response induced by the polymer injection. It has been shown that tumor necrosis factor alpha (TNF macrophages during inflammation and immune response [87] TNF differentiation and by stimulating the production of angiogenic factors from other cells [88] Figure 28 Comparison of CD31+ cells between the 6 injected groups quantified from co staining of CD31 and vWF Error bars represent standard error of the mean. indicates p < 0.05.

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55 Additionally, the SRTG + VEGF + NPs + PDGF + IL 10 group displayed significantly more ECs than saline, SRTG, and SRTG + VEGF + PDGF af ter 7 days and it showed significantly more ECs th an saline and SRTG after 21 days This demonstrates that the biological factors being released from the polymer delivery system s stimulates the proliferation of EC s Co staining for CD31 and vWF a factor involved in hemostasis, was examined to quantify the amount of functional vascular endothelial cel ls within blood vessels. After 7 and 21 days the SRTG + VEGF + NPs + PDGF + IL 1 0 and SRTG + VEGF + NPs + PDGF groups showed significantly more functional ECs (CD31+ and vWF+ cells) than saline, SRTG and SRTG + VEGF ( Figure 29 ). The SRTG + VEGF g roup did not show a statistically significant difference of functional ECs compared to the SRTG and saline groups which shows that the delivery of PDGF is important to increase the proliferation of vascular ECs ( Figure 29 ) Figure 29 Comparison of CD31+ and vWF+ cells between the 6 injected groups quantified from co staining of CD31 and vWF Error bars represent standard error of the mean. indicates p < 0.05.

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56 Mature Blood Vessel Formation Analysis Co staining for CD31 and SMA, a marker for vascular SMCs (mural cells) was e xamined to identify mature and stable blood vessels for 7 and 21 days after injection ( Figure 30). More mature vessels should show an organized circle of ECs surrounded by mural cells. After 21 days, an increase in blood vessels surrounded by mural cells was observed qualitatively for the SRTG + VEGF + NPs + PDGF + IL 10 and SRTG + VEGF + NPs + PDGF groups compared to sal ine and SRTG injections ( Figure 30). Additionally, it appeared that all biological factor injection groups stimulated the proliferation of ECs compared to saline and SRTG, sh owing that the GFs are inducing a potent an giogenic response Staining for CD31 w as first examined again to compare the proliferation of ECs associated with blood vessels and similar results were found compared to the CD31 and vWF co staining results. After 7 days, the SRTG + VEGF, SRTG + VEGF + NPs + PDGF and SRTG + VEGF + NPs + PDGF + IL 10 groups showed a statistically significant increase in EC proliferation compared to the saline and SRTG groups ( Figure 31 ). After 21 days, all 4 biological fac tor SRTG groups showed significantly more ECs than saline and SRTG groups ( Figure 31 ). Although the SRTG group did show an increase in ECs after 7 and 21 days compare d to saline, it was not significantly different which shows that the TNF macrophages may not be having as much of an angiogenic response as previously thought. Furthermore, SRTG + VEGF + NPs + PDGF group had significantly more ECs after 21 da ys compared to 7 days ( Figure 31 ).

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57 Figure 30

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58 Figure 30 Representative images of IHC co staining for ECs (CD31) and SMA) to show blood vessel maturation after 7 and 21 days for all 6 groups. Subcutaneous tissue was stained with: CD31 and Alexa Flour 488 and appear green SMA and Alexa Flour 594 and appear red and DAPI which appears blue. Scale bar represents 100 m. Figure 31 Comparison of CD31+ cells between the 6 injected groups quantified from co SMA Error bars represent standard error of the mean. indicates p < 0.05. SMA was examined to quantify the formation of mature blood vessels induced by the polymer delivery systems. After 7 days, the SRTG + VEGF + NPs + PDGF + IL 10 group induced significantly more mural cells than the saline, SRTG and SRTG + VE GF groups ( Figure 32 ). After 21 days, the SRTG + VEGF + NPs + PDGF + IL 10 and SRTG + VEGF + NPs + PDGF groups both displayed significantly more mural cells compared with the saline, SRTG and SRTG + VEGF + PDGF groups ( Figure 32 ) Furthermore, the SRTG + VEGF, SRTG + VEGF + NPs + PDGF and SRTG + VEGF + NPs + PDGF + IL 10 groups al l induced significantly more mural cells over t ime when comparing the 7 and 21 day time points ( Figure 32 ) This demonstrates that the GFs are

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59 inducing more mature neo vasculature from the early to late stages of the angiogenesis process. Figure 32 SMA+ cells between the 6 injected groups quantified from co SMA Error bars represent standard error of the mean. indicates p < 0.05. The SRTG +VEGF + PDGF group did not show any increase in mural cells between days 7 and 21, unlike the oth er delivery systems loaded with GFs ( Figure 32 ) This group does not have PDGF encapsulated within NP s so both VEGF and PDGF are most likely burst releasing simultaneously. As stated earlier, it has been shown that early stage angiogenic factors can have inhibitory effects on late stage GFs and vice versa, when presented concurrently [5,17,18,89] This data for the SRTG +VEGF + PDGF group further confirms the importance of sequentially delivering angiogenic GF s to develop stable and mature blood vessels.

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60 IHC Analysis of Immune Res ponse IHC was used to identify and quantify the number of macrophages present in the subcutaneous tissue and whether there was a reduced immune response from the polymer delivery system loaded with IL 10. Macrophages were stained with CD68 and Alexa Flour 594. Saline was used as a negative control to compare how many macrophages infiltrate the tissue from the needle injection. Positively stained cells were counted and quantified when DAPI was identified with the stains. The IH C stains qualitatively show a d ecrease in the number of macrophages for the SRTG + VEGF + NPs + PDGF + IL 10 group compared to all other polymer delivery groups at 7 and 21 days ( Figure 33 ). Moreove r, the inflammatory marker showed a large number of CD68+ cells for all polymer groups ( Figure 33 ). Figure 33 Representative images of IHC staining for macrophages (CD68) to show immune response after 7 and 21 days for all 6 groups. Subcutaneous tissue was stained with CD68 and Alexa Flour 594 and appear red and DAPI which appears blue. Scale bar represents 50 m.

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61 The number of macrophages we re additionally quantified by counting CD68+ and DAPI+ cells at the boundary area of the polymer injection site. The saline injection group showed significantly less macrophages than all polymer delivery groups at 7 and 21 days ( Figure 34 ). After 21 days, the SRTG + VEGF + NPs + PDGF + IL 10 group showed significantly less CD68+ cells than the 4 other polymer biological factor delivery groups ( Figure 34 ) This result demonstrates that IL 10 is significantly reducing the immune response caused by the polymer gel. Additionally, the SRTG + VEGF, SRTG + VEGF + PDGF and SRTG + VEGF + N Ps + PDGF + IL 10 all showed a significant reduction i s macrophages from 7 to 21 days ( Figure 34 ). Figure 34 Comparison of CD68+ cells between the 6 injected groups. bars represent standard error of the mean. indicates p < 0.05.

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62 CHAPTER VII CONCLUSION In this study, a sulfonated reverse thermal gel encapsulating novel micelle NPs was investigated for the application of localized and controlled protein delivery to induce angiogenesis. The polymer delivery system was synthesized, characterized and evaluated for its therapeutic potential with an in vitr o release test and subsequent in vivo animal study. Sulfonated PSHU PNIPAm was successfully synthesized and the molecular structure was verified through 1 H NMR, FTIR and elemental analysis. Polymeric PEG PSHU PEG micelles were fabricated and characterized by 1 H NMR, DLS and SEM to validate the chemical structure and to ensure uniform spherical size distribution. Additionally, SEM was further used to demonstra te micelle encapsulation within the sulfonated thermal gel. The entire polymer delivery system was then evaluated for controlled and sequential release of three biological factors using an in vitro release study. Release test samples were taken daily and protein release was then quantified using ELISAs. The results showed that the sulfonated PSHU PNIPAm system significantly reduced the burst release of VEGF from the polymer scaffold compared to plain PSHU PNIPAm, demonstrating the effectiveness of sulfonat e groups controlling VEGF release. Furthermore, the release profile showed that the encapsulated micelles were being slowly released from the SRTG which provided the sequential release of PDGF. This spatiotemporal release of GFs from the polymer scaffold i s critical to promote stab le and mature angiogenic vessel formation The ability of the polymer system to induce new blood vessel formation was analyzed in vivo using a subcutaneous injection mouse model. Subcutaneous injections of the SRTG polymer system with 4 different combinations of factors was evaluated for angiogenesis

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63 activity against saline and SRTG injections Histological assessment was used to directly observe blood vessel formation. Functional vascular ECs and vascular SMCs were quantified usi ng three different stains and were subsequently counted as positive cells when identified with DAPI. After 21 days, t he SRTG + VEGF + NPs + PDGF + IL 10 group showed significantly more functional ECs and SMCs than every other group, except for the SRTG + V EGF + NPs + PDGF group. These results show that the polymer delivery system with all three biological factors is inducing a substantial angiogenic response with stable and more mature vessels. The polymer system was analyzed for a reduced immune response as well. IHC was used to examine the effects of IL 10 on suppressing macrophage infiltration at the polymer injection site. A stain for macrophages was quantified by counting the number of CD68+ cells identified with DAPI. After 21 days, the SRTG + VEGF + NPs + PDGF + IL 10 group showed significantly less macrophages present in the subcutaneous tissue compared to all other polymer groups. These results are encouraging as limiting an immune response is important for optimizing cardiac repair following an MI.

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64 CHAPTER VIII FUTURE WORK Increase Sample Size SMA+ cells were not seen between the groups themselves and between the time points. This is due to the high variability from differing amounts of blood vessels while counting positive c ells from the selected images Increasing the sample size would provide more tissue to assess histologically and this could reduce variability which may allow for observation of significant s tatistical differences between the groups time points. Additional Time Points Although statistical significant differences between some of the groups were not observed, different trends between the groups were seen from 7 to 21 days that could be further elucidated with a longer time point at 4 or 6 weeks. Many mature blood vessels were SMA+ cells after 21 days, but an additional time point to allow the vessels to mature further could provide a more apparent difference between the groups. Increase Biological Factor Loading Amount To further increase the statistical significant differences between the groups, increasing the amount of proteins loaded into the polymer delivery system could enhance the amount of functional and mature blood vessels. GFs have a short half life in vivo and are quickly degraded in circulation so increasing the local amount of GFs to the tissue could promote higher EC proliferation rates while also recruiting more SMCs to the vessels. Additionally,

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65 increasing the amount of IL 10 could help reduce the amount of inflammation caused by the polymer system. Acute Myocardial I nfarction Animal Model As the target of this polymer delivery system is for treatment of an MI, an acute MI mouse model should be implemented to examine the effects of revascularizing ischemic myocardium and reducing detrimental cardiac remodeling. Validating the efficacy of thi s biological factor delivery system for treatment of MI is essential to determine if this system can produce therapeutic angiogenesis effects in the heart. IL 10 has been shown to attenuate post MI left ventricular dysfunction, reduce infarct size and atte nuate MI induced cardiac cell death [33] so it is important to demonstrate that this polymer system can deliver IL 10 and provide those same benefits. Echocardiography, IH C, TUNEL staining and Western blot analysis will all be utilized to examine the therapeutic effects of this protein delivery on treatment for MI.

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