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LOCAL, STRUCTURAL AND SYSTEMIC CUES FOR VASCULAR TISSUE ENGINEERING
By
VIDHYA RAMASWAMY
A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT
OF THE REQUIREMENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY
UNIVERSITY OF FLORIDA
2017
© 2017 Vidhya Ramaswamy
To every little girl with a dream of a career in science and engineering
4
ACKNOWLEDGMENTS
சிறு துளி பெரு பெள்ளம் Little drops make the mighty ocean
This humble contribution of mine to the ocean of science would not have been
possible without, the immense support from friends, family, colleagues and esteemed
mentors. I would first like to acknowledge my advisor, Dr. Josephine Allen, for taking a
chance with me and giving me this opportunity and freedom to expand my skills and
research methodologies. Her patient guidance and faith in me, has been instrumental in
shaping my doctoral research experience. I would also like to thank my committee
members, Dr. Antonio Webb, Dr. Jennifer Andrew, Dr. Christopher Batich and Dr.
Chelsey Simmons for the valuable interactions and insightful questions, that have
provided me the framework for developing this dissertation. I am also greatly thankful
for the constant support and guidance I have received from each of my committee
members to shape my career. I would also like to specially acknowledge Dr. Nancy
Ruzycki, for her continual support and encouragement throughout my tenure in
graduate school. I have knocked on her door on countless occasions seeking help, and
I have always left her office with hope, enthusiasm and sweet treats. I would also like to
acknowledge my lab alumni and current members who have helped greatly and
contributed valuably to my research projects: Taylor Repetto, Elliott Dirr, Krista Dulany,
Allison Goins and Dr. Agata Zupanska.
Any achievement of mine is the result of the great amount of sacrifice, patience
and hard work of my parents, K. Ramaswamy and R. Vijayalakshmi and grandparents,
R. Krishnamoorthy and K. Pattammal. It is their constant unwavering belief in me and
limitless affection, that has fueled my career so far. I would also like to acknowledge
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with great appreciation my brother, Dr. Bharath Ramaswamy, who has been, a great
source of inspiration, a dear friend, philosopher and a wise guide. Without his immense
patience, guidance and encouragement, my dream of a career in science would not
have been possible. A special thanks to my cousin, Sumathy who has been a big
source of inspiration for work-life balance.
Last but not the least, I want to thank all my wonderful friends: Praveen, Shalini,
Koushik, Pavithra, Neha, Aparna, Meena and Laura for being my personal
cheerleaders, and listening to my endless woes about graduate school and life. I am
permanently indebted to their belief in me, constant encouragement and positivity.
6
TABLE OF CONTENTS Page
ACKNOWLEDGMENTS .................................................................................................. 4
LIST OF TABLES .......................................................................................................... 10
LIST OF FIGURES ........................................................................................................ 11
LIST OF ABBREVIATIONS ........................................................................................... 14
ABSTRACT ................................................................................................................... 19
CHAPTER
1 INTRODUCTION .................................................................................................... 21
Specific Aim 1- Nanoscale Local Cues ................................................................... 23
Specific Aim 2- Mesoscale Structural Cues ............................................................ 23 Sub-aim 1 ......................................................................................................... 24 Sub-aim 2 ......................................................................................................... 24
Specific Aim 3- Macroscale Systemic Cues ............................................................ 24
2 BACKGROUND ...................................................................................................... 25
Cardiovascular Diseases and Tissue Engineering .................................................. 25 Nanoscale Local Cues ............................................................................................ 29
Mesoscale Structural Cues ..................................................................................... 35 Small Diameter Vascular Grafts ....................................................................... 38 Vascular Stents ................................................................................................ 42
Macroscale Systemic Cues ..................................................................................... 46
3 NANOSCALE LOCAL CUES- DEVELOPMENT OF A NOVEL VEGFR2 AGONIST APTAMER ASSEMBLY ......................................................................... 53
Introduction ............................................................................................................. 53 Specific Aim 1 ......................................................................................................... 56
Materials and Methods............................................................................................ 56 Aptamer Design ................................................................................................ 56 Aptamer-Endothelial Cell Binding ..................................................................... 58 VEGFR2 Phosphorylation ................................................................................ 60
Signal Pathway Activation ................................................................................ 61 Endothelial Nitric Oxide Synthase Expression .................................................. 61 Capillary Tube Formation (Angiogenesis) ........................................................ 62 Chorioallantoic Membrane (CAM) Assay.......................................................... 63
Results and Discussion........................................................................................... 64 Summary ................................................................................................................ 74
7
4 MESOSCALE STRUCTURAL CUES- PART 1- DEVELOPMENT OF A BIOMIMETIC, BILAYERED VASCULAR SCAFFOLD FOR TISSUE ENGINEERING SMALL DIAMETER BLOOD VESSELS ........................................ 76
Introduction ............................................................................................................. 76 Specific Aim 2 ......................................................................................................... 77
Sub-aim 1 ......................................................................................................... 77 Sub-aim 2 ......................................................................................................... 77
Materials and Methods............................................................................................ 81
Materials ........................................................................................................... 81 Solid Lumen Fabrication ................................................................................... 82 Fabrication of Electrospun Medial Layer .......................................................... 82
Glutaraldehyde Crosslinking ............................................................................. 83 FTIR Characterization ...................................................................................... 84 SEM Characterization ...................................................................................... 84
Contact Angle Measurement ............................................................................ 84 Uniaxial Tensile Testing ................................................................................... 85
Biocompatibility Characterization ..................................................................... 85 Cell culture ................................................................................................. 85 Scaffold preparation for cell studies ........................................................... 85
Cellular attachment .................................................................................... 86 Lactate dehydrogenase cytotoxicity assay ................................................. 86
Cell proliferation assessment ..................................................................... 87 Optimizing Bilayer Fabrication with Electrospun POC-PAA.............................. 88 Fabrication of Bilayered Construct ................................................................... 88
Mechanical Properties ...................................................................................... 89 Compliance ................................................................................................ 89
Circumferential hoop stress ....................................................................... 89 Statistical Analysis ............................................................................................ 90
Results and Discussion........................................................................................... 90
5 MESOSCALE STRUCTURAL CUES- PART 2- HEMOCOMPATIBILITY EVALUATION OF A NOVEL MG-CA-SR ALLOY FOR VASCULAR STENT APPLICATIONS .................................................................................................... 106
Introduction ........................................................................................................... 106 Materials and Methods.......................................................................................... 110
Alloy Synthesis and Sample Preparation ....................................................... 110
Cytotoxicity Evaluation ................................................................................... 110 Hemolysis ....................................................................................................... 112 Evaluation of Platelet Adhesion and Activation .............................................. 113 Whole Blood Clotting Assessment ................................................................. 114 Release of Monocytic Inflammatory Cytokines ............................................... 114
Statistical Analysis .......................................................................................... 115 Results And Discussion ........................................................................................ 115 Summary .............................................................................................................. 125
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6 MACROSCALE SYSTEMIC CUES- PART 1- EFFECT OF SIMULATED MICROGRAVITY ON DIFFERENTIATION AND DOWNSTREAM FUNCTION OF PORCINE BLOOD DERIVED ENDOTHELIAL PROGENITOR CELLS .......... 128
Introduction ........................................................................................................... 128 Specific Aim 3 ....................................................................................................... 131 Methods and Materials.......................................................................................... 133
Isolation of Peripheral Blood Mononuclear Cells from Porcine Blood ............. 133 Rotating Wall Vessel Bioreactor System ........................................................ 134
Microgravity Culture of PBMNCs .................................................................... 134 Differentiation, Colony Formation and Immunocytochemistry ........................ 135 Endothelial-Like Cell Migration and Proliferation ............................................ 135
Nitric Oxide and Prostaglandin Secretion ....................................................... 136 Clotting Kinetics .............................................................................................. 137 Statistical Analysis .......................................................................................... 137
Results And Discussion ........................................................................................ 137
7 MACROSCALE SYSTEMIC CUES- PART 2- EFFECT OF SIMULATED MICROGRAVITY ON FUNCTION OF HUMAN BLOOD DERIVED ENDOTHELIAL CELLS......................................................................................... 147
Introduction ........................................................................................................... 147
Materials and Methods.......................................................................................... 149 Isolation of Peripheral Blood Mononuclear Cells from Human Blood ............. 149
Culture and Differentiation of Endothelial Progenitor Cells ............................. 150
Immunocytochemical Assessment of Differentiated Endothelial Cells ........... 150
Simulated Microgravity Culture of HE-Like Cells ............................................ 151 Cell Proliferation Assessment ......................................................................... 152 Gene Expression ............................................................................................ 152
Endothelial Cell Secretory Function ............................................................... 153 Plasma Clotting Kinetics ................................................................................. 153
Statistical Analysis .......................................................................................... 154 Results and Discussion......................................................................................... 154 Summary .............................................................................................................. 166
8 SUMMARY, CONCLUSIONS AND FUTURE WORK ........................................... 171
Summary of Results.............................................................................................. 171 Conclusions .......................................................................................................... 177 Future Work .......................................................................................................... 178
APPENDIX: PROTEIN CROSSLINKING USING CINNAMALDEHYDE VAPORS ...... 180
Electrospinning Gelatin Nanofibers ....................................................................... 181 Cinnamaldehyde Crosslinking and Preliminary Stabilization Assessment ............ 181 FTIR Characterization ........................................................................................... 182 SEM Characterization ........................................................................................... 183
9
LIST OF REFERENCES ............................................................................................. 184
BIOGRAPHICAL SKETCH .......................................................................................... 212
10
LIST OF TABLES
Table page 3-1 Aptamer sequences and configuration ............................................................... 57
4-1 Average fiber diameter of Collagen:POC:Elastin fibers with different cross-linking times. ....................................................................................................... 95
4-2 Contact angles at the water – biomaterial interface ............................................ 96
4-3 Uniaxial tensile test. Assessment of tensile properties of the glutaraldehyde vapor crosslinked electrospun mat ..................................................................... 97
4-4 Circumferential mechanical testing of bilayered scaffolds compared to native tissue and ePTFE.. ........................................................................................... 103
A-1 CA cross-linking conditions and resistance to degradation .............................. 181
11
LIST OF FIGURES
Figure page 1-1 Multi-scale cues presented to cells ..................................................................... 24
2-1 Cardiovascular diseases and tissue engineering strategies. .............................. 27
2-2 Different types of cues that can direct stem cell behavior for tissue engineering. ........................................................................................................ 28
2-3 A representative schematic of a single iteration of the SELEX process. ............ 32
2-4 Mechanism of aptamer binding to complex substrate. ........................................ 33
2-5 Schematic representation of cellular response to microstructural and physical properties of scaffold matrices. ........................................................................... 36
2-6 A schematic representation of the structure of native artery.. ............................. 42
2-7 Some representative bioreactor systems for tissue engineering. ....................... 48
2-8 Path of a particle or cell aggregate in a RWV bioreactor illustrating the net forces acting on the particle in rotation. .............................................................. 50
3-1 Aptamer structure. .............................................................................................. 58
3-2 Schematic representation of the proposed method of action of the divalent aptamer assembly. Binding to target and activation via dimerization, membrane bound VEGFR2. ............................................................................... 65
3-3 Aptamer-cell binding interactions.. ...................................................................... 66
3-4 Aptamer Induce VEGFR2 activation and downstream molecular events ........... 68
3-5 Functional assessment of VEGFR2 aptamer binding. ........................................ 72
3-6 Chorio allantoic membrane assay. ..................................................................... 73
4-1 Schematic representation of the bilayered vascular graft. .................................. 81
4-2 Glutaraldehyde vapor crosslinking.. .................................................................... 83
4-3 Synthesis of poly diol citrates. ............................................................................ 91
4-4 SEM characterization as spun collagen:POC:elastin fibers. ............................... 93
4-5 SEM characterization of all crosslinking conditions using glutaraldehyde vapor.. ................................................................................................................ 94
12
4-6 ATR-FTIR spectra of collagen: POC: elastin crosslinked fibers. ......................... 95
4-7 Contact angle measurements. ............................................................................ 96
4-8 SEM images of cell attachment. hASMCs seeded onto collagen: POC: elastin fibers. ...................................................................................................... 98
4-9 Relative cytotoxicity measurements of collagen: POC: elastin with hASMCs and POC lumen with HUVECs. .......................................................................... 99
4-10 Proliferation kinetics. ........................................................................................ 100
4-10 SEM characterization of POC; POC-PAA bilayered construct. ......................... 101
4-12 SEM characterization of POC; collagen: POC: Elastin bilayered construct ...... 102
5-1 Evaluation of cytotoxicity. ................................................................................. 116
5-2 Hemolysis assessment. .................................................................................... 118
5-3 Whole Blood Clotting assessment.. .................................................................. 120
5-4 Platelet adhesion and activation. ...................................................................... 122
5-5 Release of inflammatory cytokines/chemokines by THP-1 cells reported as absorbance at 450 nm. ..................................................................................... 124
6-1 RWV bioreactor set-up. .................................................................................... 132
6-2 Differentiation and characterization of PE-like cells. ......................................... 139
6-3 Proliferative and migratory functions of simulated microgravity pre-exposed PE-like cells. ..................................................................................................... 141
6-4 Representative results of quantification of release of soluble factors. .............. 143
6-5 Representative result showing quantification of blood clotting studies. ............ 145
7-1 Bioreactor set-up and operation. ...................................................................... 151
7-2 Isolation of PBMNCs and differentiation into ECs. ............................................ 156
7-3 Proliferation kinetics ......................................................................................... 158
7-4 Gene expression. ............................................................................................. 160
7-5 Secretory functions of cells.. ............................................................................. 161
7-6 Plasma clotting kinetics. ................................................................................... 164
13
A-1 FTIR spectra of cross linked fibers ................................................................... 182
A-2 SEM characterization of CA crosslinking. ......................................................... 183
14
LIST OF ABBREVIATIONS
µg Microgravity
ACD Acid Citrate Dextrose
AHA American Heart Association
ANOVA Analysis of Variance
Apt80mer 80 Mer VEGFR2 Binding Aptamer Sequence
AptDIVALENT Divalent VEGFR2 Binding Aptamer Assembly
AptM80mer Modified Base 80 Mer VEGFR2 Binding Aptamer Sequence
ASTM American Society for Testing and Materials
BME Basement Membrane Extract
BMP-2 Bone Morphogenic Protein
BSA Bovine Serum Albumin
CA Cinnamaldehyde
CaCl2 Calcium Chloride
CAM Chorio Allantoic Membrane
CD 309 Kinase Insert Domain Receptor Or VEGFR2
CD 31 Cluster of Differentiation 31
cDNA Complementary DNA
CPC Circulating Progenitor Cells
CSC Circulating Stem Cells
CVD Cardiovascular Diseases
DACRON Polyethylene Terephthalate
DMEM Dulbecco's Minimum Essential Medium
DNA Deoxyribonucleic Acid
15
dsDNA Double Strand DNA
E Elastic Modulus
EBM-2 Endothelial Basal Medium-2
EC Endothelial Cells
ECM Extracellular Matrix
EGM-2 Endothelial Growth Medium-2
ELISA Enzyme Linked Immunosorbent Assay
eNOS Endothelial Nitric Oxide Synthase
EPC Endothelial Progenitor Cells
ePTFE Expanded Polytetrafluoroethylene
FBS Fetal Bovine Serum
FDA Food and Drug Administration
FITC Fluorescein Isothiocyanate
FTIR Fourier Transform Infrared Spectroscopy
GAPDH Glyceraldehyde 3-Phosphate Dehydrogenase
HARV High Aspect Rotating Vessel
hASMCs Human Aortic Smooth Muscle Cells
hEGF-2 Human Epidermal Growth Factor
HE-like cells Human Endothelial-Like Cells
hFGF Human Fibroblast Growth Factor
HRP Horseradish Peroxidase
HSP70 Heat Shock 70 Kda Protein
HSPA4 Gene Encoding Heat Shock 70 Kda Protein 4
HUVECs Human Umbilical Vein Endothelial Cells
IGF Insulin-Like Growth Factor
16
IgG Immunoglobulin G
IL-1β Interleukin 1 Beta
IL-6 Interleukin 6
IL-8 Interleukin 8
ISO International Organization for Standardization
ISS International Space Station
LDH Lactate Dehydrogenase
LED Light Emitting Diode
MTT 3-(4,5-Dimethylthiazol-2-Yl)-2,5-Diphenyltetrazolium Bromide
NASA National Aeronautics and Space Administration
NO Nitric Oxide
PAA Poly Acrylic Acid
PAEC Porcine Aortic Endothelial Cells
PBMNCs Peripheral Blood Mononuclear Cells
PBS Phosphate Buffered Saline
Pburst Burst Pressure
PCL Polycaprolactone
PCR Polymerase Chain Reaction
PDC Poly Diol Citrates
PDGF-BB Platelet Derived Growth Factor BB
PE Phycoerythrin
PEG Polyethylene Glycol
PE-Like cells Porcine Endothelial-Like Cells
PF 4 Platelet Factor 4
PGA Polyglycolic Acid
17
PGE2 Prostaglandin E2
PLGA Poly (Lactic-Co-Glycolic Acid)
PLLA Poly (L-Lactic Acid)
POC Poly (1,8 Octane Diol-Co-Citrate)
PRP Platelet Rich Plasma
PU Polyurethane
PVDF Polyvinylidene Fluoride
qRT-PCR Quantitative Reverse Transcriptase - Polymerase Chain Reaction
r Radius of tubular scaffold
R3-IGF-1 Human Recombinant Insulin-Like Growth Factor
RBC Red Blood Cells
RNA Ribonucleic Acid
RWV Rotating Wall Vessel
S.S Stainless Steel
SELEX Systematic Evolution of Ligands by Exponential Enrichment
SEM Scanning Electron Microscopy
SMC Smooth Muscle Cells
STLV Slow Turning Lateral Vessel
t Thickness of tubular scaffold
TCP Tissue Culture Plastic
THP-1 Human Monocytic Cell Line
TNF-α Tumor Necrosis Factor Alpha
tPA Tissue Plasminogen Activator
VEGF Vascular Endothelial Growth Factor
VEGFR2 Vascular Endothelial Growth Factor Receptor-2
18
vWF Von Willebrand Factor
ε Strain
σ Stress
19
Abstract of Dissertation Presented to the Graduate School of the University of Florida in Partial Fulfillment of the Requirements for the Degree of Doctor of Philosophy
LOCAL, STRUCTURAL AND SYSTEMIC CUES FOR VASCULAR TISSUE
ENGINEERING
By
Vidhya Ramaswamy
August 2017
Chair: Josephine Allen Major: Materials Science and Engineering
The chronic scarcity of autologous donor organs to treat cardiovascular diseases
has led to increased efforts in alternate therapeutic strategies. One such approach is
the development of novel tissue engineering methods, to create functional replacements
for diseased or damaged parts of the vasculature. Living cells alter their behavior as a
direct response to intrinsic and extrinsic cues presented to them. The underlying
principle of tissue engineering is that the living cell system is the only true “tissue
engineer,” and the goal is to deliver appropriate cues to modulate the cellular response.
The goal of this dissertation is to study the underlying mechanism of cell-material
interactions at different scales to develop and deliver novel biomaterial based cues to
repair, preserve, and regenerate vascular tissue. Starting at the nanoscale, the first part
of this work involved the development of a novel divalent aptamer assembly that acts as
an agonist towards vascular endothelial growth factor receptor-2 in endothelial cells.
Besides activation, the assembly was also able to cause a positive downstream
angiogenic response, which proves that this is a promising strategy for vascular tissue
engineering. The second part of this work consisted of studying cues designed at the
meso scale including interaction of cells with surrounding scaffolds. This section
20
covered two classes of cardiovascular biomaterials; small diameter vascular grafts and
vascular stents. The interaction of these biomaterials with in vitro models of the vascular
environment (blood and vascular cells) was studied to develop scaffold material based
strategies for tissue engineering. The third and final part of this dissertation was focused
on macroscale systemic cues, that provide a unique environment for modulating the
response of vascular cells. One such specialized environment is simulated microgravity,
which can be used to effectively study tissue engineering of 3D structures. The goal of
this work was to understand the behavior of circulating stem cells in this low shear
stress environment, to assess their therapeutic potential for vascular repair and
regeneration. Overall, this work is expected to contribute towards widening the
knowledge base in the study of cell-material interactions and open avenues for research
in this field.
21
CHAPTER 1 INTRODUCTION
Despite remarkable advances in research in the management and prevention of
complications arising from cardiovascular disease (CVD), it still remains the leading
cause of death worldwide1–3. In the United States alone, nearly $300 billion is spent on
healthcare costs annually towards managing CVD4. In most cases, these diseases
cause irreversible organ damage where the body is not able to regenerate or repair the
damage by itself. These complications have created a crucial need for novel methods to
stimulate repair and regeneration of the cardiovascular system. Traditional treatments
for CVD include medications and surgical interventions. However, most surgical options
involve circumventing the diseased or damaged vascular tissue, (bypass stents or
grafts), or completely replacing them (transplants). But the chronic shortage of available
donors, has steered the treatment of these conditions towards developing alternative
therapies such as tissue engineering, which can significantly expand patient care
options. The goal of such strategies is to efficiently provide means for repair,
preservation and regeneration of vasculature to establish normal functioning of the
cardiovascular system4,5 .This approach will also greatly diminish the need for scarce
donor organs, thereby improving clinical point of care for patients affected by these
diseases. This need is further fueled by aging population and increased expectations for
a better quality of life. Besides their intended use, the new engineered tissues also find
application in disease studies, drug testing and in vitro development organotypic model
systems6,7.
Vascular tissue engineering is aimed at applying principles of engineering and
life sciences towards the development of novel strategies to restore, maintain and
22
enhance cardiovascular tissue function6,8,9. Living cells react and alter their behavior as
a direct response to intrinsic and extrinsic cues presented to them. This behavior makes
it possible to design vehicles to deliver these cues to cell systems10. The use of
biomaterials as potential tools to achieve these end goals has been gaining popularity
recently. This is mainly due to the flexibility to alter physical, chemical and structural
properties of biomaterials to invoke a specific cellular response11. These cues can be a
variety of different factors such as pH, temperature, pressure, nutrient supply and waste
removal, mechanical, electrical and magnetic stimuli, etc. These in turn can cause the
modulation of important cellular functions such as cell growth, migration, differentiation,
survival or even apoptosis. These cues can also be applied across various length
scales varying from subcellular molecular cues to systemic macroscale cues12–14.
Biomaterials engineering exploits the structure-property relationship of
biomaterials to present these specific cues that can modulate a favorable response from
cell systems. This field makes use of a large and extensive toolkit comprising of material
parameters and properties available to modify or control an exact cellular function15,16.
Extensive research over the past few decades has shown that while designing
strategies to create living replacement for existing vascular tissue, the complexity of the
organ system itself must be taken into consideration. The underlying principle is that the
living cell system is the only true “tissue engineer,” and the goal is to deliver appropriate
stimuli, to the cellular microenvironment on multiple hierarchical levels to effectively
recreate tissue formation17,18.
Tissue formation in the body is a direct result of highly synergistic and
coordinated sequences of cell renewal, growth, differentiation, and assembly, within a
23
dynamic environment characterized by spatial and temporal gradients of multiple
factors6,8,17. Therefore, to replicate and stimulate these functions of cells, it is imperative
to study and understand the biological stimuli at multiple levels to design and deliver the
appropriate cues. Thus, there is a need for a multifaceted approach spread across
different scales while studying strategies to enhance tissue engineering. The goal of this
dissertation is to study the underlying mechanism of cell-material interactions at
different scales (nano, meso and macroscales), to develop novel strategies to repair,
preserve and regenerate vascular tissue. A thorough and comprehensive understanding
of the complex and vast mechanisms of complete tissue regeneration is challenging and
is out of the scope of this work. However, this study provides essential elements of
understanding, of distinct systems towards furthering the knowledge in the field of
vascular tissue engineering. The dissertation has been divided into three parts with
each part discussing studies dealing with a specific length scale as shown in Figure 1-1
below. To probe cells at each of the length scales of interest, I have established the
following three specific aims pertaining to each specific scale of interest.
Specific Aim 1- Nanoscale Local Cues
Design and develop of a novel divalent DNA oligonucleotide sequence (aptamer),
that when assembled into a unique divalent assembly acts as a receptor agonist towards
VEGFR2.
Specific Aim 2- Mesoscale Structural Cues
Understand and develop biomaterial based strategies towards vascular tissue
engineering and regeneration by studying the interactions of biomaterials with in vitro
vascular microenvironments.
24
Sub-aim 1
Develop a novel biomimetic, bilayered, biomaterial construct and evaluate its
applicability as a small diameter vascular graft.
Sub-aim 2
Evaluate the biocompatibility and hemocompatibility of an Mg-Ca-Sr alloy to
assess it suitability for use in designing vascular stents.
Specific Aim 3- Macroscale Systemic Cues
Study and understand the effect of simulated microgravity on the function of
vascular stem cells to assess their potential to be used as a therapeutic strategy for
vascular repair and regeneration.
Figure 1-1. Multi-scale cues presented to cells
25
CHAPTER 2 BACKGROUND
Cardiovascular Diseases and Tissue Engineering
In the statistical update summary released by the American Heart Association
(AHA), the overall rate of death attributed to Cardiovascular diseases (CVD) alone was
reported to be 235.5 per 100,000 in 20101,19. This data revealed that more than 2150
Americans die of CVD every day with an average of 1 death every 40 seconds. The
total cost of medical procedures performed in the treatment of CVD, amounts to over
300 billion dollars each year3,19,20. The treatment costs of CVD alone were more than
any other disease group. With these numbers growing every year, treatment strategies
to manage and alleviate the complications arising from CVD has become a national
concern.
Cardiovascular diseases can manifest in different ways in the body, often
resulting in death. Some of the conditions that can be classified under the broad
umbrella of cardiovascular diseases include rheumatic heart disease, congenital heart
disease, coronary heart disease, cardiomyopathy, deep vein thrombosis, inflammatory
heart disease and other disorders19. Although, significant efforts are taken towards
preventing the occurrence of these conditions by closely monitoring the underlying
causative factors, and providing additional care for patients in the high-risk category,
there has not been a significant reduction in mortality due to CVD. Therefore, a great
portion of healthcare costs are being invested in CVD treatment every year.
Current treatment strategies include, pharmaceutical drugs and novel medical
devices which have significantly improved the quality of life for patients with CVD. But
these have failed to help patients with end-stage CVD. Moreover, in some cases
26
despite drug therapy the condition eventually becomes refractory to therapy and
disease progression is inevitable4,21. The golden standard for treatment of CVD has
been complete organ replacement or partial tissue transplantation. But this approach is
very much limited by the scarcity of suitable autologous tissue that is healthy and
available for transplantation. With the exponential increase in the need for donors which
is not matched by the availability of donors, more recent approaches have been focused
on developing tissue engineering based strategies to replace diseased and damaged
part of cardiovascular tissue4,6.
Tissue engineering is a broad field of science that combines the principles of
engineering, material science, and biology into development of functional tissue
substitutes that can repair, maintain, or regenerate tissue function in damaged
segments of the body6. Within the field of cardiovascular tissue engineering, a great
amount of progress has been made. Tissue engineered replacements for various
constituents of the cardiovascular system, including blood vessels, heart valves, and
cardiac muscle have been developed over the years4,8,22,23. The extent of progress
made in this field of research has created a promising environment for widespread
application of tissue-engineered therapy for CVD.
Many noteworthy studies have reported new and improved techniques including
endothelial cell seeding of vascular grafts, tissue-engineered vascular conduits,
generation of heart valve leaflets, cardiomyopathy, genetic manipulation, and in vitro
conditions for optimizing tissue-engineered cardiovascular constructs11,22–25. Many of
these cross-over concepts have been simultaneously applied to treat multiple disorders
of the heart. The figure below (Figure 2-1) summarizes some of the recent tissue
27
engineering based therapy developed for various complications due to CVD. From the
tissue engineering point of view, the schematic below also highlights the importance of
using biomaterial scaffolds as delivery vehicles for appropriate vascular cells on the site
of injury, to enable regeneration. The figure also shows the need for diverse scaffold
designs as well cell types, used for different tissues in the cardiovascular system.
Figure 2-1. Cardiovascular diseases and tissue engineering strategies.Image reference :http://www.sciencecodex.com/files/different%20ways%20in%20which%20tissue%20engineers%20believe%20that%20heart%20muscle.jpg
The underlying concept behind tissue engineering therapy is that, the living cell
system is the ‘true tissue engineer’. Any tissue engineering strategy is focused on
exploiting the resilient ability of this system to recuperate and regenerate damaged
tissue by modulating its response when provided with the appropriate stimuli6,17,26.
28
Living cells are highly sensitive to extrinsic cues that can control their functions.
Their complex sensory mechanisms can detect and respond to a broad variety of
signals that can be transduced into useful cell functions10,13,17,18. These cues can range
from environmental changes such as change in osmolarity, pH, temperature, etc., to
subcellular molecules such as growth factors, cytokines, etc. These in turn play a very
important role in regulating important cell functions, such as cell migration,
differentiation, proliferation and even apoptosis. One such example is represented in
Figure 2-2 below, showing a schematic summarizing different types of cues that can be
applied to a stem cell environment to modulate a favorable response for tissue
engineering.
Figure 2-2. Different types of cues that can direct stem cell behavior for tissue engineering. Image reference: Mashayekhan S, Hajiabbas M, Fallah A. Stem Cells in Tissue Engineering. In: Pluripotent Stem Cells. Vol InTech; 2013. doi:10.5772/54371.
29
Another important point to note is that, these extrinsic cues applied to cell
systems can either be employed separately, or in correlation with each other. These
cues can often work synergistically and generate a complex response otherwise
unachievable by modulating a single factor. This concept is represented by the
interconnected lines in the above schematic as well.
This sensitivity of living cell systems has been exploited in directing cell function
for tissue engineering and regeneration for many years now. Biomaterials engineers are
focused on designing appropriate cues to control cell functions for tissue engineering. In
the process of designing cues, the science behind the interaction of cell systems with
the external micro/macro environment has also been extensively studied and
reported10,26,27. Many significant contributions to this field have been made in the recent
past that have revolutionized biomaterial design in tissue engineering. This chapter will
briefly cover the major trends in technological advancement in this field. For better
organization and understanding, the following text has been divided into 3 subsections,
including discussion about specific cues at 3 different scales (nano, meso and macro).
The rationale behind this organization is to correlate this brief review with the theme of
this dissertation as described in the Chapter 1.
Nanoscale Local Cues
From the birth of nanotechnology in 1959, through Richard Feynman’s seminal
lecture, this field has metamorphosed into an all-encompassing technology applied
across the scientific realm28. The ability to precisely tailor bulk properties of materials by
hierarchical design of nanomaterial constituents has caused increased interest in this
field. In addition to design of materials, the vested interest in nanotechnology has also
caused studies into understanding phenomenon happening at the nanoscale.
30
Application of nanotechnology to tissue engineering and regeneration, stems
from the hierarchical nanostructure of cells and tissues and the extracellular
environment29,30. Initial attempts at tissue engineering failed to consider the dynamic
microenvironment of the cells and its effect on altering cell behavior. As interest grew in
this field, emphasis was placed on mimicking the extracellular environment in healthy
tissues to regenerate and repair damaged tissues10,31. Instead of simply delivering cells
in the diseased area, this approach was combined with stimuli that help engineer the
cells to repair and regenerate the tissue. These cues vary both in scale and
spatiotemporal functionality13.
To truly mimic the extracellular environment, it is essential to start from the nano
or subcellular scale. Many studies over the years have revealed the potential
importance of cues at the nanoscale to affect important functions of cells such as
signaling, migration, adhesion, proliferation, and differentiation18,30. Various strategies to
deliver nanoscale cues to cells have since been investigated. Some current strategies
include but are not limited to, nanopatterned biomaterial scaffolds, nanoparticles,
nanodevices, proteins, growth factors, chemokines and small molecules18,30,32,33.
In particular the delivery of biological and synthetic signaling molecules such as
low-molecular-weight drugs, proteins and oligonucleotides that can stimulate cell
migration, growth and differentiation for tissue engineering has been successfully tried
and tested in in vitro settings34–36. The design strategies of signaling molecules has
been based on the understanding of natural tissue morphogenic mechanisms which is a
coordinated action of multiple signaling molecules. Broadly, two categories of such
molecules have been studied in relation to tissue engineering therapy; 1) soluble
31
signaling molecules such as growth factors and other regulatory proteins that are
secreted within the cell system’ and 2) Insoluble signaling factors derived from the ECM
environment. Of these, the delivery of many growth factors such as vascular endothelial
growth factor (VEGF) which regulates and controls primarily all vascular repair and
angiogenic mechanisms, platelet derived growth factor (PDGF-ββ) which promotes
would healing in most tissues, bone morphogenic protein-2 (BMP-2) which has known
to improve bone regeneration mechanisms have been studied extensively34,37,38.
Although many of these growth factors have been used in clinical trials for tissue
regeneration therapy, only PDGF- ββ, BMP-2 and BMP-7 have shown promise37,39–42.
These factors however are not cost effective and often have safety limitations. For
example, despite successful preclinical and phase I studies, clinical trials of VEGF
therapy have not been very encouraging. Particularly the delivery of VEGF at safe
dosages to target tissues to induce tissue repair has been a major challenge37,43,44.
Other engineering challenges for delivering growth factors include the need to control
the microenvironmental tissue distribution, which is inherently difficult with gene delivery
techniques. Although delivery of recombinant proteins could overcome the distribution
issue, they have too short half-lives in vivo. Additionally, growth factors also cause non-
specific activation events and invoke immunogenic responses which are not desirable45.
Therefore, current work in this field has been focused on alternate synthetic molecules
and engineered signaling factors that can specifically target cellular domains and
stimulate tissue regeneration. A promising approach to this challenge is the delivery of
synthetic peptides and engineered RNA and DNA sequences, such as aptamers, to
32
target and activate specific targets within the cells35,43,46. Some preliminary studies have
shown encouraging results via the use of aptamers towards this goal47,48.
Nucleic acid aptamers are short single-stranded oligonucleotides which bind to a
pre-determined target with very high specificity and binding affinity. Aptamers were first
discovered nearly 3 decades ago, in 1990, when several groups are attributed to have
isolated RNA aptamers. In the pivotal work by Tuerk and Gold where two high affinity
RNA ligands were enriched from a randomized library49. The iterative selection method
used for enriching specific ligands was named as SELEX, short for systematic evolution
of ligands by exponential enrichment. In a separate work by Ellington and Szostak, the
term aptamer was coined to refer to high affinity molecules35. A representative
schematic explaining the SELEX process is shown below in Figure 2-3. The figure
shows a single iteration process in SELEX where high affinity aptamers are selected
through a sequentially purifying candidate sequences by binding to a target molecule.
Figure 2-3. A representative schematic of a single iteration of the SELEX process.
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The group was able to identify and isolate RNA aptamers that could specifically
bind to organic dyes. Ever since their initial discovery and isolation, thousands of
aptamers have been isolated for a variety of targets including but not limited to peptides,
proteins, viruses, bacteria, receptors, organic molecules, metallic ions and even whole
cells47,48,50–53.
Aptamers have a unique advantage when compared to other oligonucleotides
such as siRNAs in therapeutics. While siRNAs can only be designed to intracellular
targets, aptamers are more flexible and can be developed to target cell-surface, intra
and extracellular substrates54,55. The therapeutic ability of aptamers have been largely
based on their ability to form complex 3D secondary structures such as hairpins, loops,
quadruplexes etc. that enables them to dock on and bind to complex domains on
substrates51. The mechanism of specific binding of aptamers to substrates on cells is
represented in the schematic below.
Figure 2-4. Mechanism of aptamer binding to complex substrate.
This characteristic of aptamers has been largely used to inhibit protein-protein
interactions such as receptor-ligand interactions thereby acting as antagonists of
receptor function. Although the binding affinity and specificity of aptamers make them
ideal therapeutic agents, in vivo translation and delivery of aptamers has proven to be a
34
challenge. The clinical studies for the use of aptamers have been largely limited by their
susceptibility to nuclease-mediated degradation in the presence of serum and renal
filtration. This reduces their circulation time and affects their efficacy in the body.
However, more recent studies showed that this challenge can be overcome by
chemically modifying the nucleotide sugars or internucleotide phosphodiester linkages
55. Further, these modified aptamers have also been screened using the SELEX
process. Another promising approach has been the capping of the oligonucleotide
termini, thereby increasing its stability in serum. With these important advancements,
many aptamer ligands have been identified for variety of targets such as Angiopoietin1,
Angiopoietin 2, VEGF, Nucleolin, Thrombin, PDGF etc47,53. Of which many aptamers
such as Pegaptanib (Targeting VEGF), AS1411 (targeting nucleolin in cancer cells),
NU172 (targeting thrombin) etc. have been approved for use clinically as well 48,56–60.
With increased advancement in the therapeutic applications of aptamers which
exploit their antagonistic ability, their potential application in tissue engineering and
regeneration has also been investigated. Aptamers are being increasingly used to
capture cells and proteins on biomaterial scaffolds to promote cell infiltration and
endothelization of scaffolds. Additionally, a new and exciting field of research with
aptamers has been their use to modulate cell function for tissue engineering
applications. Remarkable work by Dollins et al.61, reported the synthesis of an OX-40
specific aptamer assembly that acted as a receptor agonist. Since then there has been
increased interest in developing aptamer based agonists. A recent report has also
identified an aptamer that can cause activation of insulin receptors62. This work
describes the design and development of a divalent aptamer assembly for controlled
35
activation of VEGFR2 receptor in endothelial cells in Chapter 3. This study opens
avenues to develop aptamer based strategies to control and modulate cellular functions
for tissue engineering applications. Due to the flexibility of designing and developing
aptamer substrates and the wide variety of target molecules that they can be designed
for, a potential direction for future work in this field, is developing supra-molecular
multimer assembly that can deliver specific subcellular cues to cells to program their
function to repair and regenerate tissue in vivo60.
Mesoscale Structural Cues
The previous section discussed the importance of designing stimuli at the
nanoscale. Especially the use of small molecule triggers to control and modulate cellular
function. Most tissue systems are known to be anchorage dependent which means that
they require a substrate to attach to and grow. In healthy tissues this support is provided
by the extra cellular matrix that surrounds the cellular environment.
As discussed earlier, the extra cellular environment is a complex hierarchical
system constituting of multiple levels of functional components. This complex structure
makes it essential to study the cues presented by each hierarchical level to mimic the
existing extracellular environment63. Moving up on the scale of hierarchy, this section
describes strategies to modulate cellular response by designing appropriate cues on the
mesoscale. The mesoscale is defined as being inclusive of constituents ranging
between 10 to 100 microns in size14. This scale includes the interactions between
biomaterial substrates and cells. These interactions can be divided broadly into two
categories, (a) cells that are in direct contact with biomaterials, and (b) cells that are
indirectly exposed to an environment which consists of biomaterials.
36
Important cellular responses and characteristics including alteration of cell shape,
alignment, differentiation, growth and tissue formation have been known to be
modulated by architectural cues that are presented to the cell systems by biomaterials
that are in contact with them14,30,33. This concept has inspired the design of biomaterials
for tissue engineering. For example mesoscale structural parameters such as fibers,
grooves, patterns, roughness as well as physical cues such as stiffness, surface
chemistry, elasticity and micro hardness of substrates have known to play an important
role in cellular morphology, alignment, growth, migration and differentiation14,26. This is
depicted in the schematic shown in Figure 2-5 below. A typical example of this
phenomenon is the behavior of stem cells in substrates of varying stiffness. In the
seminal work by Adam Engler in 200664, the dependence of matrix stiffness in directing
stem cell lineage was reported. Similarly many studies since then have shown the effect
of substrate mechanical and topographic properties on stem cell fate65.
Figure 2-5. Schematic representation of cellular response to microstructural and physical properties of scaffold matrices. Image reference: Kollmannsberger P, Bidan CM, Dunlop JWC, et al. The physics of tissue patterning and extracellular matrix organisation: how cells join forces. Soft Matter. 2011;7(20):9549.
37
In the vascular world, biomaterial substitutes have been used to engineer a
variety of scaffolds and devices that are used to treat many vascular diseases as shown
in Figure 2-1. Early vascular biomaterial substrates and devices were devised as a
means of support to damaged or diseased vasculature and to house transplanted
cells/tissues to repair and regenerate the area. However, over the years, the inherent
functionality and cues provided by these scaffolds have also gained importance.
Biomaterials used for vascular tissue engineering can broadly be classified into two
categories; synthetic and natural materials66,67. Synthetic materials used in vascular
tissue engineering for cardiovascular system primarily consists of polymers and metals
or sometimes even a composite of both. Natural materials are mostly derived from plant
or animal tissues. They are inherently the preferred choice for biomaterial implants due
to their excellent biological functionality that makes them acceptable for tissue
engineering applications. Natural materials have the advantage of possessing excellent
biocompatibility, being non-immunogenic and eliciting minimal inflammatory response
from the body. However, they do not match up to the expected levels of mechanical
robustness for a cardiovascular implant in most cases. Additionally they are also
expensive and difficult to process68,69. Therefore, synthetic materials have been
researched as an alternative. One of the major advantages of synthetic materials is the
ability to tailor their physical, chemical and topographical properties. They are stronger
and more mechanically robust in comparison to natural biomaterials. They also are
cheaper to obtain and possess more flexibility of processing. The application of
synthetic materials is however greatly limited by concerns of long term cytotoxicity and
inflammatory response by the body66,67,70,71.
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The definition of biocompatibility of biomaterials is dynamic and has been
changing with advancement in research in this field. However, the general requirements
of a cardiovascular biomaterial include, mechanical compatibility, low cytotoxicity,
noncalicification, compatibility with blood and blood components, minimal inflammatory
or immunogenic response and in most cases capability to support vascular cell
systems66,67. Traditionally, preferred biomaterials used for medical applications have
been inert systems that do not cause adverse reactions in the body. However more
recent research has been focused on developing functional bioactive scaffolds that can
deliver cues to cell systems that come in contact with them. The idea behind this
approach is to design and tailor the chemical, physical or topographic features of
biomaterial substrates to effectively influence the interaction at the cell-biomaterial
interface.
Cardiovascular biomaterials in general encompasses a wide variety of material
applications ranging from vascular grafts to hemodialysis tubes. These biomaterials can
be designed to alleviate complications arising due to different cardiovascular diseases,
such as atherosclerosis, myocardial infarction, peripheral artery disease etc. This
section will briefly discuss two classes of cardiovascular biomaterial applications,
namely, small diameter vascular grafts and vascular stents, respectively. This section
will also briefly summarize the interaction of these biomaterials with in vitro models of
the vascular environment to understand and develop strategies to repair damaged
vasculature and enable regeneration.
Small Diameter Vascular Grafts
Owing to the clinical difficulty in obtaining autograft blood vessels to replace
diseased vasculature, alternative strategies are much sought after. Early studies dating
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back to 1980’s reported the development of completely synthetic grafts25,72,73. These
studies led to the development two important candidate materials for vascular grafts;
polyethylene terephthalate (Dacron®; INVISTA Technologies S.à.r.l., Saint Gallen,
Switzerland) and expanded polytetrafluoroethylene (Teflon®; E.I. du Pont de Nemours
and Company, Wilmington, DE, USA). Both these materials have been used clinically,
with good success rates as replacement materials for large diameter blood vessels73,74.
However, they failed to perform adequately in small-diameter vessels with disappointing
results from clinical attempts. Many synthetic grafts since then have been developed but
they have proven to be largely unsuccessful resulting in low patency and clinical
complications such as neointimal hyperplasia and increased thrombosis71,75–77. As an
improvement on the previous materials, polyurethanes were then identified as potential
candidates for the vascular grafts. Polyurethanes exhibit good mechanical compliance
thromboresistive characteristics and support the growth of living tissue. However their
use was limited by their poor biostability71.
After the initial setbacks in the use of purely synthetic substrates, emerging
needs steered the research in this field from ’replacement’ to ‘regeneration’ of the
functional tissue. To this end, a biomimetic approach was employed while designing
scaffolds for regeneration. More and more natural materials were included in the
scaffold design. In an important work by Weinberg and Bell in 1986, collagen tubular
scaffolds seeded with cells was studied and this study became one of the first reports to
use protein based scaffolds78. Even though since then many collagen based scaffolds
have been fabricated and also have been utilized in animal studies, they lacked
mechanical strength and required the use of a synthetic support layer. Many other
40
natural materials such as fibrin, silk, gelatin and elastin have been used to create
vascular grafts but all of them have had limited success due to mechanical mismatch
with native tissue68,79. Therefore, more recently there has been interest in developing
composite materials which are mechanically compatible but also possess the biological
functionality of natural materials. Natural ECM proteins such as collagen and elastin
have been blended with synthetic polymeric systems such as poly(l-lactide) (PLLA),
polycaprolactone (PCL), Poly(l-lactide-co-glycolide) (PLGA) and polyurethanes (PU).
The first part of the work discussed in Chapter 4 of this dissertation describes the
development of one such protein-polymer hybrid scaffold for small diameter vascular
grafts. This study investigates the use of a combination of a biodegradable, elastomeric
polymer; poly (1,8 octanediol-co-citrate) (POC) with natural extracellular matrix proteins;
collagen and elastin. POC belongs to a family of elastomers composed of citric acid and
an aliphatic diol. POC has been identified as a hydrophilic and biodegradable polymeric
constituent that can substitute for elastin in synthetic fibres80–83. It is also anti-
thrombogenic and supports differentiated endothelial progenitor cell adhesion and
growth84,85.
Even though the base material of the scaffold largely determines its physical,
chemical characteristics and its mechanical properties, the processing of scaffold plays
a major role in determining its functional ability. The scaffold must be fabricated in the
correct shape and size matching the target tissue, and be sufficiently strong to
withstand repetitive mechanical loading. The scaffold must also promote appropriate
vascular cell phenotype and present a large surface area to volume ratio to enhance
cell attachment and growth86. Further the scaffold design should include porous areas to
41
improve cell migration, signaling and 3D vasculogenesis87.This porous network is
important for nutrient delivery, waste removal, and migration of cells and signaling
molecules33,66,88,89.
A biomimetically inspired scaffold design attempts to best represent the
architecture of the native ECM. Native artery as shown in Figure 2-6 consists of a solid
inner layer known as the lumen and a fibrous exterior layer also known as the medial
layer. The inner luminal layer is composed primarily of a monolayer of endothelial cells
that are in direct contact with circulating blood. The medial layer, consists of multiple
layer of smooth muscle cells. This layer is also rich in collagen and elastin among other
ECM proteins and provides a great amount of mechanical support to the vessel90–92.
This organizational structure is essential to enable vascularization and enhanced
reintegration into existing healthy vasculature. Therefore, some studies in the past
decade have adopted this concept and designed vascular scaffolds consisting of
multiple layers93–95. As mentioned above the scaffold design has also been made to
include porous layers to account for the 3D integration of cells and vascular growth.
Many techniques such as use of porogens, 3D printing, electrospinning, solvent casting
and phase separation has been employed to include porosity in scaffolds86,96,97. Of
these techniques, electrospinning offers robust material selection, low cost, simplicity,
high surface-to-volume ratios, and controlled porosity for vascular scaffold fabrication.
With these advantages, electrospinning is a proven framework allowing for the improved
endothelialization of vascular grafts. In this context, Chapter 3 describes the design and
development of a bilayered vascular graft that consists of a solid lumen layer and a
fibrous medial scaffold layer to mimic the native architecture of the blood vessel.
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Figure 2-6. A schematic representation of the structure of native artery. Image reference: Blausen.com staff (2014). "Medical gallery of Blausen Medical 2014". WikiJournal of Medicine 1 (2). DOI:10.15347/wjm/2014.010. ISSN 2002-4436.
Future work in this field is focused on identifying and utilizing an appropriate cell
source to promote the reendothelization of vascular grafts. Therefore, the next
generation grafts are being designed to recruit stem cells from circulating blood to mimic
natural blood vessel regeneration in the body. Additionally, further topographical
modulations, and functionalization of scaffold surface with bioactive molecules are being
investigated to improve graft functionality46,98–100.
Vascular Stents
Metals dominate the field of cardiovascular biomaterials next to if not equal to
polymeric materials67. The use of metals for biomedical application dates to as early as
1985. Majority of the applications of metals in medicine was in the form of heart valves,
43
vascular stents and stent-graft combinations. The excellent mechanical properties and
workability of metals have furthered the application of metals for high stress applications
within the body101. However, some of the commonly used bio-metals possess limited
biofunctionalities such as blood compatibility, bone conductivity and bioactivity. To be a
candidate material for cardiovascular applications, a material must exhibit compatibility
with the cellular environment including blood and its components. Failure of which,
leads to complications such as thrombosis and inflammation which not only prevents
incorporation of the implant into the existing vasculature but also causes significant
damage to it.
Balloon-expandable stents were first used in 1985 for coronary arteries to treat
complications arising from coronary artery diseases24,67,102. Many metals have been
used to fabricate vascular stents since then. Some of the earlier stent design utilized
inert materials such as stainless steel. Further modifications were made to the
composition of the stainless steel such as addition of Cr and Ni to improve the strength
of the alloy. It was then furthered altered by reducing the percentage of carbon and
include molybdenum to improve corrosion resistance101,103. This composition was
designated as 316 L which is still currently the golden standard for stents and several
316 L SS have also been FDA approved for stent applications. Several other metals
such cobalt, tantalum, molybdenum, titanium and their alloys were also used to develop
metallic stents. Stents made of Ti and its alloys were particularly preferred for these
applications due to their light weight, high strength and excellent corrosion resistance.
Even though these bare metal stents performed well in terms of mechanical behavior,
they had serious setbacks such as increased thrombotic or inflammatory response.
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More importantly, due the permanent nature of these implants, they also caused
damage to vessel geometry and caused negative remodeling101,102,104. Therefore,
another important development in this field was the noteworthy discovery of degradable
stents that gradually dissolve away and are replaced in entirety by the native tissue.
These stents are designed to assist in arterial remodeling, a process which has been
showed to take 6-12 months105–108. Therefore, the degradation rate of materials chosen
for this application was designed to match this rate to promote successful remodeling.
This strategy greatly improved the efficacy of these stents and reduced the need for
anticoagulation or anti inflammation therapies. However, with biodegradable alloys, it
becomes increasingly important that the degradation products as well as the degraded
stent be compatible with the vascular environment. It is especially important that the
stent material be compatible chemically with the vascular environment and does not
cause significant toxicity to the endothelium or the circulating blood107–109. Polymers
have been considered as potential candidates to develop degradable stents.
Polyamides, Polyolefins, Polyesters, Polytetrafluoroethylene, Poly(L-Lactide) and
Polyurethanes have all been utilized to make vascular stents. However the unknown
long term toxicity as well as poor mechanical behavior of the polymeric materials have
deterred development of polymeric stents71,110. Therefore, there is more recent interest
in biodegradable metals for the fabrication of stents.
The first patent for a degradable metallic stent was issued in 2002. Several iron
and magnesium based metallic alloys have been investigated since then107,111. The
rationale behind the selection of these metals was their natural presence in the body
and their positive contribution to bodily functions. Iron is an essential element in the
45
body and has known to cause relatively low cytotoxicity and therefore was among the
first degradable element investigated as candidate material for stents105,112. Magnesium
is also an essential element in the body with relatively low toxicity levels and it is also
efficiently excreted out in the urine. Additionally, magnesium has known to positively
impact cardiovascular mechanisms in the body113,114. Therefore, more recently
magnesium is gaining popularity as a base element for vascular stents.
In Chapter 4 of this dissertation, a novel magnesium alloy is evaluated for its
potential use as a vascular stent material. This chapter deals with assessment of
hemocompatibility of this alloy for its suitability for use in a vascular environment.
Parameters such as surface finishing conditions greatly influence the chemical
composition of the surface. For example, electropolishing of 316 L stainless steel
leaches the surface of most metallic elements except chromium. This process leads to a
formation of a layer of chromium oxide when the stent is exposed to external
environment, preventing further corrosion115. Other finishing processes such as
abrasive polishing cause an enrichment of elements on the surface as well as randomly
oriented crystals at the interface. These parameters such as the microtopography of the
surface, local chemical composition, roughness and crystal size offer important
mesoscale cues which subsequently alter cell behavior. Surface defects as well as trace
elements such as elements at grain boundaries have been known to cause alterations
in proteins and affect cellular response. Therefore the formation of an oxide layer is
beneficial in degradable metallic implants107.
46
Future strategies for development of biodegradable metallic stents are focused
on structured, long-term and multidisciplinary investigation of implant efficacy in vivo.
These studies are particularly useful in clinical translation of this technology.
Macroscale Systemic Cues
The capability of cells to sense environment and produce signals that influence
cell behavior, phenotype and cell fate is one of the most important ways to modulate cell
response. The previous sections described the effect of soluble small molecule triggers
and biomaterial scaffolds on cellular responses critical for tissue engineering and
regeneration. However, cells exhibit a high level of sensitivity to biophysical cues from
the environment such as mechanical forces, physical alterations and gravitational
stresses. In the past decade, the mechanobiological cues that influence cell behavior
have been identified and used favorably to promote regeneration27,116. The major
roadblock to generation of functional tissue engineered tissues that can be used in a
clinical setting has been the limited understanding of physicochemical cues that affect
tissue development as well as the high cost of the existing tissue engineered
products25,117. Bioreactors provide reproducible, standardized, and controlled
biomechanical and environmental stimuli to cells in a closely monitored system. Some
of these environmental cues include, pH, temperature, waste removal, nutrition supply,
mechanical stresses, gravitational changes, etc. The great degree of reproducibility,
environment control and automation provided by bioreactors has led to their application
in large scale applications118–121.
Bioreactor systems historically have been employed in various applications such
as fermentation, food processing, waste water treatment, drug testing as well as protein
synthesis. They eliminate the need for in vivo testing at initial stages of development of
47
therapeutics by simulating an organotypic environment118,122,123. With respect to tissue
engineering, 3D tissue engineered structures are typically cultured on porous scaffolds
which provide a support to house the cells. While initially these were initially cultured on
a static 2D culture environment, with increased interest in biomimetic designs,
bioreactor culture became more popular. Long-term bioreactor culture of these
structures promoted efficient nutrient transfer, ability to provide mechanotransducive
forces as well as close monitoring of tissue growth in an in vitro model that mimicked
the body more closely than tissue culture plates. In the human body, cells are constantly
subjected to mechanical, electrical and chemical forces that are essential for their
synchronized metabolic functions. Many types of bioreactor systems have been
developed in the past. Some of the earliest important bioreactors include, the spinner
flasks, rotating wall vessel bioreactor, hollow fiber bioreactors, direct perfusion systems
as well as custom designed bioreactor systems 122–124 as shown in Figure 2-7 below.
Each of the bioreactor systems mentioned have unique properties such as increased
turbulence as well as convective mass transfer (Spinner flasks), low shear environment
(Rotating wall vessels), precise mechanical control (Custom designed bioreactors and
Perfusion systems). The limitations as well as the advantages of bioreactors depend on
their specific applications. However, any bioreactor should follow one of the following
functions according to Freed : (1) provide a framework for uniform distribution of cells on
a three-dimensional scaffold; (2) maintain the desirable concentration of gases and
nutrients in the culture medium; (3) stimulate efficient mass transfer to the growing
tissue; (4) presenting developing tissue with controlled physical stimuli; and (5) provide
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information regarding the formation process of 3D tissues, which originate from the
isolated cells124,125.
Figure 2-7. Some representative bioreactor systems for tissue engineering. (A) Spinner flasks; (B) Rotating wall vessel bioreactors (C) Hollow-fiber bioreactor (D) Perfusion systems (E) Custom designed computer-controlled system. Image reference: Martin I, Wendt D, Heberer M. The role of bioreactors in tissue engineering. Trends Biotechnol. 2004;22(2):80-86. doi:10.1016/j.tibtech.2003.12.001.
One such bioreactor system gaining popularity as an ideal vehicle for tissue
engineering are rotating wall vessel bioreactors (RWV)119,126,127. Initially, the motive for
the fabrication of microgravity bioreactors was to create a system that can simulate the
49
changes in gravitational forces in outer space on earth. This was developed by NASA
as a cost-effective model system to study and preliminarily experiments space flight
experiments. But over the recent years, space research has also contributed to
application of some of its principles in tissue engineering127. Microgravity provides a
unique environment that promotes efficient nutrient transfer, cell-cell signaling in a low
shear stress environment. In real microgravity, cells have been known to aggregate and
initiate differentiative signaling which can cause 3D tissue constructs. Further
microgravity has also known to promote stem cell proliferation which also is an
important for tissue engineering126,128–130.
In 1970, NASA’s Johnson center assembled a team of scientists to develop an
earth based system for simulating some aspects of microgravity to pretest space based
tissue engineering studies. The first two bioreactor models developed by NASA were
called Slow turning lateral vessels (STLV) and High aspect ratio vessel (HARV)119,127.
The RWV bioreactors of two concentric cylinders which rotate at a constant
angular speed. The reactor also consists of an axial silicone membrane for gas
exchange. The vessels can be fabricated in different volumes and speed of rotation is
also adjusted per the system used. Initial studies with simulated microgravity
bioreactors hypothesized that the expected damaging hydrodynamic forces on cells
would be from high velocity gradients and associated shear stress involved in the
bioreactors. The appearance of this shear stress is through the boundary layer that
exists between fluid moving inside the vessel relative to the vessel wall. The RWV
bioreactor however eliminates this gradient resulting in a reduced shear stress
environment131,132.
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Figure 2-8. Path of a particle or cell aggregate in a RWV bioreactor illustrating the net forces acting on the particle in rotation.
The diagram above represents the path of a particle or cell aggregate and the net
forces acting on the particle in rotation is also illustrated. The sedimentation velocity due
to gravity, Vs, is composed of an inwardly directed component, Vsr, and a tangential
component, Vst. There is an outwardly directed motion, Vcenr, produced by centrifugal
force, and a tangential component, Vcort, from the Coriolis force. Coriolis force is an
inertial force (also called a fictitious force) that acts on objects that are in motion relative
to a rotating reference frame the resolution of these forces on cells or aggregates as
shown produces a slow descent through the culture media as the vessel turns at a pre-
determined speed.
Because the net forces on the cells are substantially reduced, this culture
environment is sometimes referred to as simulated or modeled microgravity. The results
obtained using simulated microgravity conditions utilizing the RWV have been
confirmed and validated by experimental models in space flights to outer space133–136.
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The initial attempts to engineer tissues in simulated microgravity, led to culturing
of cartilaginous tissue constructs consisting of a collagen and glycosaminoglycan base
in 1997126,137. This study showed positive results with higher regenerative tissue
formation compared to other methods. Interestingly, the study showed that the cells
formed 3D aggregates as the cells were cultured for longer times in microgravity. This
system has also been utilized to culture stem cells and vascular cells for vascular tissue
engineering. In a recent study by Grimm et al., EA.hy926 endothelial cell lines cultured
under simulated microgravity formed tube-like structures with walls mostly consisting of
single-cell layers. They were found to have a lumen and resembled the intima of blood
vessels in vivo, reaching a length of up to several centimeter138,139. More interestingly, a
recent study by Chiu et al., showed that human cord blood stem cells when cultured
under simulated microgravity proliferated and trans-differentiated into 3D structures
resembling vascular tubules while exhibiting vascular endothelial phenotype
expressions140. These studies have opened exciting avenues for the use of simulated
microgravity bioreactors as a controlled environment for vascular tissue engineering.
Additionally, this environment has also known to support bone tissue growth,
which was studied by culturing osteoblasts cell in the RWV reactor141. These systems
have also gained popularity in developing cancer models due to the improved mass
transfer characteristics and 3D growth environment129,142. Many excellent recent reviews
have summarized the effect of simulated microgravity on the behavior of various cell
systems128,129,143,144. In Chapters 6 and 7 of this dissertation, the effect of simulated
microgravity on blood derived endothelial cell systems are investigated and reported.
The goal of this study was to use simulated microgravity as a model system to
52
understand the behavior of these cells to further study the vascular repair mechanisms
enabled by these cells. This study has the potential to develop regenerative therapy for
tissue engineering.
Future studies using microgravity mediated tissue engineering rely on confirming
the efficacy of these constructs in enabling vascular repair in patients. Current studies in
this field is focused on translating these strategies to production of whole sensory
organs and other high-density structures.
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CHAPTER 3 NANOSCALE LOCAL CUES- DEVELOPMENT OF A NOVEL VEGFR2 AGONIST
APTAMER ASSEMBLY
Introduction
In addition to intrinsic factors produced within cells that regulate the normal
functioning of cell systems, living cells are highly sensitive to extrinsic cues presented to
them at the molecular scale by the surrounding environment10,17,18. The interactions of
these nanoscale triggers with cellular systems are highly efficient in eliciting a multiscale
and complex response within the body. These triggers can range from biochemical cues
such as soluble growth factors, cytokines and other small molecules to nano-structural
cues provided by the ECM. Due to the small scale of these interactions, there is still a
dearth of information in understanding these mechanisms in vivo10,14,18. However,
recent advances in tissue engineering, regenerative medicine, and cell based therapies
necessitate the development of novel strategies based on delivery of these nanoscale
cues that allow for controlling cellular processes for an application of interest. The end
goal is to design approaches that are targeted, specific, robust, and ultimately result in
molecular pathway activation events and the transduction of signals that direct cellular
responses, such as cell proliferation, differentiation, tissue matrix production or even cell
death. For receptor-mediated cell processes, the identification of stable receptor
substrate ligands that exhibit precise binding, high specificity, and targeted activation of
a particular receptor represents an area of active research145. Current efforts in this
field, aim to use receptor-specific growth factors and cytokines or their respective
receptor-binding peptide sequences, as these are responsible for native interactions
that result in receptor activation. While inherently successful as receptor agonists,
these agents, often being difficult to control spatially, can lead to off-target effects
54
through the activation of nonspecific receptors and binding to non-specific cell
types34,146. Unfortunately, identification of therapeutic agents that selectively activate
specific cell surface receptor targets has proven to be quite a challenge. Overcoming
these limitations in the control of receptor-mediated processes will result in distinct
improvements in directing cellular processes such as proliferation, migration, and
differentiation.
A promising approach to achieving this goal that is gaining in popularity in the
recent times is the use of nucleic acid aptamers. Aptamers are short single-stranded
oligonucleotides that exhibit highly specific binding to intracellular and extracellular
target sites147. The high affinity and specificity of aptamer binding can be attributed to
their ability to form complex three dimensional structures51. This is typically due to their
capability to form complementary base pairs allowing them to form a variety of
secondary structures that can be further formed into three dimensional assemblies.
These assemblies are recognized by their cognate targets and form complexes, similar
to that formed when antibodies interact with antigens. The aptamer-target interaction
can be via hydrophobic and electrostatic interactions, hydrogen bonding, van der Waals
forces, shape complementarity and base stacking. Such interactions enable the specific
binding of the aptamer to the target with high affinity.
Ever since their isolation using the SELEX (Systematic Evolution of Ligands by
Exponential Enrichment) process, which was developed in the lab of Larry Gold at the
University of Colorado148, aptamers have been studied extensively as a diagnostic and
therapeutic tool for a wide range of applications35,53,149. In addition, using this method,
aptamers have been selected against a wide array of targets from small molecules,
55
proteins and bacteria, to whole cells (i.e. whole-cell SELEX)36,47,150. The major
advantages in using aptamers over traditional antibodies and growth factors are their
high affinity and targeting specificity as well as their non-immunogenic nature. Aptamers
can target and bind to small hidden domains within a large biomolecule that can be
inaccessible to larger antibodies. Aptamers are also easily synthesized and can be
modified to confer nuclease resistance, and thus stability to withstand filtration by the
kidney and improves circulation time in vivo47. With regard to aptamer-receptor
interactions, most work with aptamers exploits the ability of aptamers to inhibit protein-
protein interactions, such as receptor-ligand interactions, thereby rendering them
receptor antagonists47. This approach has been quite successful in the area of
neovascularization and angiogenesis inhibition151,152. Pegaptanib (Macugen) targeted
against VEGF, is currently the only federally approved aptamer therapeutic drug153. The
higher affinity of aptamers to their targets as compared to other receptor ligands, such
as antibodies or growth factors, makes them excellent at blocking receptor activation.
However, the therapeutic ability of agonist or receptor- activating aptamers is yet to be
explored completely.
As a departure from current strategies, this work shows the ability of an
engineered aptamer assembly to act as a receptor agonist that activates VEGFR2 on
human endothelial cells. The activation of VEGFRs by VEGF plays a crucial role in
many processes of vascular stem and mature cells, including cell proliferation,
migration, differentiation, survival, tissue factor production, and nitric oxide production,
as well as functions such as angiogenesis39. The delivery of VEGF, local as well as
systemic, has been extensively investigated in order to improve angiogenesis for
56
neovascularization in vascular tissue engineering. Although the strategy has proven to
be highly effective in vitro, the clinical translation still has some gaps yet to be filled. For
example, in a study by Simón-Yarza et al. (2012)37, VEGF delivery for clinical trials for
myocardial regeneration revealed no significant improvement with respect to placebo.
Intravenous administration of the growth factor also proved to be ineffective due to its
short in vivo half-life (~30 min) and overall dose was limited by off-target site toxicity
issues. Therefore, there in a need to identify alternate strategies to specifically cause
receptor activation to improve angiogenesis and other downstream functional
responses. With that in mind, the following aim was established and tested.
Specific Aim 1
Design and develop of a novel divalent DNA oligonucleotide sequence (aptamer),
that when assembled into a unique divalent assembly acts as a receptor agonist towards
VEGFR2.
Materials and Methods
Aptamer Design
The aptamers used in this study include an 80 mer VEGFR2 binding DNA
aptamer with modified nucleic acids that was obtained from Aptamer Sciences (AptSci,
Pohang, Gyeongbuk, South Korea) and was used to confirm endothelial cell binding.
The sequence (Apt80mer) is shown in Table 3-1. Following validation of endothelial cell
binding with the commercially available aptamer sequence, subsequent receptor
agonist studies were conducted with a modified 80mer DNA aptamer of similar
sequence, however replacing the benzyl containing modified bases (BndU) (indicated
as 5 in the sequence) with thymine (Integrated DNA Technologies, Coralville, Iowa).
The modification of the bases results in slightly lower binding affinity, however the
57
binding specificity towards VEGFR2 is maintained. The modified sequence (AptM80mer) is
also shown in Table 3-1.
Table 3-1. Aptamer sequences and configuration
Towards the goal of assessing aptamer agonist function, aptamer assembly
composed of two identical AptM80mer monomers tethered to each end of an 18-atom
hexa-ethylene glycol spacer thus creating a divalent assembly (AptDivalent) (Integrated
DNA Technologies) was fabricated. Polyethylene glycol (PEG) was chosen to be the
spacer of the divalent aptamer assembly because it exhibits many qualities desirable for
the functionality of the assembly. First, PEG shows favorable biological properties in
that it is non-toxic and non-immunogenic. Additionally, the hydrophilic nature of PEG
helps to support water solubility of the assembly. Further, PEG is flexible and does not
impose steric hindrance on aptamer molecules bound to it. Finally, PEG molecules
have been shown extensively in the literature to prevent non-specific binding and thus
may prevent non-specific aptamer binding154. As another selection criterion, the root
mean square distance between the two subunits of VEGFR2 is reported as 3.5 Å155.
Therefore, to make possible the binding of both receptor subunits, the 18-atom hexa-
Aptamer Description Sequence
Apt80mer 80mer monomer 5’ GAT GTG AGT GTG TGA CGA GC5 ACG ACG 5C5 GG5 G5A A55 5A5 AAA GAC AC5 G5G 5A5 A5C AAC AAC AGA ACA AGG AAA GG 3’
AptM80mer Modified 80mer monomer
5’ GAT GTG AGT GTG TGA CGA GCT ACG ACG TCT GGT GTA ATT TAT AAA GAC ACT GTG TAT ATC AAC AAC AGA ACA AGG AAA GG 3’
AptDivalent 5’-3’/spacer/5’-3’ 5’ GAT GTG AGT GTG TGA CGA GCT ACG ACG TCT GGT GTA ATT TAT AAA GAC ACT GTG TAT ATC AAC AAC AGA ACA AGG AAA GG 3’/ (PEG)6 spacer /5’ GAT GTG AGT GTG TGA CGA GCT ACG ACG TCT GGT GTA ATT TAT AAA GAC ACT GTG TAT ATC AAC AAC AGA ACA AGG AAA GG 3’
58
ethylene glycol spacer with an end-to-end chain length of 21 Å and a calculated Flory
radius of 10.26 Å was selected, as this length would make such a dimerization
mechanism possible156. It should be noted that while this mechanism is not
demonstrated in this work, based upon the length of the spacer, it is possible with the
AptDivalent assembly. The predicted secondary structure with the lowest free energy, and
a schematic representation of the divalent aptamer assembly are shown in Figure 3-1 A
& B.
Figure 3-1. Aptamer structure. (A) The predicted secondary structure with the lowest free energy of the AptM80mer monomer (structure predicted by UNAfold) (B) Schematic representation of the divalent aptamer assembly (AptDivalent) showing two AptM80mer monomers tethered through the 3’ and 5’ ends to an 18-atom hexa-ethylene glycol spacer.
Aptamer-Endothelial Cell Binding
To determine the ability of the 80mer aptamer and the modified base 80mer,
Apt80mer and AptM80mer, respectively, to bind to the VEGFR2 receptor, flow cytometric
analysis was performed using human umbilical vein endothelial cells (HUVEC) (Lonza,
Baltimore, MD). HUVEC between passages 5-9 were cultured in endothelial cell basal
media (EBM-2) (Lonza), supplemented with hEGF-2, hydrocortisone, hFGF-b, VEGF,
IGF-1, ascorbic acid, and gentamycin/amphotericin-B, and 2% fetal bovine serum (FBS)
in 37°C CO2 incubator. The VEGFR2 binding aptamer (Apt80mer) and the modified base
aptamer (AptM80mer) containing a terminal amine group were obtained to allow for
A B
59
fluorescein isothiocyanate (FITC) modification. The FITC tag was incubated with the
amine-terminated aptamers separately in borate buffer (pH 8.5) for 1 h at room
temperature with gentle mixing. The excess FITC was removed via the use of three
consecutive Zeba desalting columns (Pierce biotechnology, Rockford, IL) to ensure
complete removal of the dye. The amine group on the aptamer was then linked to the
isothiocyanate group forming a thiourea type bond between the FITC molecule and the
aptamer. The FITC conjugated aptamer solution was suspended in phosphate buffered
saline (PBS) (pH 7.4) and stored in the dark. Control experiments included the use of
VEGFR2 antibody (CD309) and an IgG isotype control antibody for testing non-specific
binding (BD Biosciences, San Jose, CA). In preparation for flow cytometry, HUVEC
were detached from tissue culture polystyrene (TCPS) flasks using cell dissociation
media, then split into separate micro-centrifuge tubes, each containing a suspension of
300,000 cells. The FITC conjugated aptamers i.e.; Apt80mer and AptM80mer previously
prepared were then incubated with HUVEC for 1 h the in suspension to allow for the
aptamer binding to the cell surface VEGFR2. This was followed by a washing step
(three centrifugal spins and washes) utilizing PBS containing 0.1 % Tween-20 to ensure
the removal of non-specifically bound aptamer. The cells were suspended in PBS
containing 4% FBS and run on an Eclipse EC800 flow cytometer (Sony Biotechnology,
Champaign, IL). The data was analyzed using FlowJo (Treestar, Ashland, OR) and
plotted as a histogram. To confirm binding specificity of the Apt80mer and AptM80mer, in a
separate experiment sample, HUVEC were treated as described, except the cells were
incubated with 2 ng/mL soluble VEGF prior to incubation with the aptamer to block the
target VEGFR2 receptor. In another experiment, AptM80mer binding when in competition
60
with soluble VEGF was assessed. In this experiment, HUVEC were incubated
simultaneously with 150nM AptM80mer and 50, 75, 100 or 150nM soluble VEGF. Flow
cytometry was conducted to detect changes in aptamer binding affinity in the presence
of soluble VEGF and the data was analyzed using FlowJo.
VEGFR2 Phosphorylation
The activation of VEGFR2 by both the monomeric form and divalent aptamer
assembly was tested by assessing phosphorylation events of VEGFR2 on HUVEC.
Phospho-VEGFR2 was quantitatively measured spectrophotometrically using Human
Phospho-VEGF R2/KDR DuoSet IC (R&D Systems, Minneapolis, MN). Briefly, HUVEC
were grown in complete media previously described to nearly 100% confluence on
TCPS. The cells were then serum starved overnight in basal media containing 0.2%
FBS and 0.1 % Bovine serum albumin (BSA) (Sigma Aldrich, St. Louis, MO).
Immediately prior to stimulation, the HUVEC were washed with PBS, and then
stimulated for 10 min in either unsupplemented basal media (BM), unsupplemented
basal media containing 100 ng/mL VEGF, unsupplemented basal media containing 300
nM AptM80mer or unsupplemented basal media containing 150 nM AptDivalent. Following
stimulation, the cells were first rinsed with ice cold PBS, and then incubated on ice for
15 minutes in ice cold lysis buffer containing 10 μg/mL each of protease inhibitors
Aprotinin and Leupeptin (R&D Systems). The lysates were collected and assayed by
following the standard ELISA protocol provided by the manufacturer (R&D Systems).
The optical density of the samples was immediately determined using a Synergy H1
microplate reader (Biotek) with the absorbance measured at 450 nm and wavelength
correction set at 540 nm.
61
Signal Pathway Activation
Activation of the downstream Akt pathway by the divalent aptamer assembly was
assessed via immunological labeling and flow cytometry. HUVEC were grown to 80%
confluence on TCPS prior to stimulation, and then subsequently, serum starved
overnight in EBM-2 devoid of FBS and growth factors. Following culture overnight in
serum free media, the cells were stimulated for 20 min in either unsupplemented serum
free basal media (EBM-2), unsupplemented basal media containing 2 ng/mL VEGF, or
in unsupplemented basal media containing 300 nM AptDivalent. Following the stimulation,
the cells were washed with PBS, trypsinized and split into microcentrifuge tubes
containing 300,000 cells. The stimulated cells were permeabilized using a cold 100%
methanol solution at 4°C for 15 minutes, washed 3 times to remove excess methanol
and fixed using fluorofix buffer (Sony Biotechnology, Champaign, IL) for 30 minutes at
room temperature. The cells were again washed using PBS three times to remove any
excess fixative in preparation for staining. The cells were then probed using Anti-
phospho Akt antibody (Ser473) (EMD Millipore, Billerica, MA) (250 µg/sample) for 1 h in
the dark with occasional gentle mixing. The cells were subsequently washed 3 times
with PBS containing 0.05 % Tween-20 to remove non-specifically bound antibody. The
cells re-suspended in PBS containing 4% FBS, were subsequently run on the Eclipse
EC800 flow cytometer and the data was examined and plotted using FlowJo.
Endothelial Nitric Oxide Synthase Expression
Changes in expression of endothelial nitric oxide synthase (eNOS) were
assessed using western blot analysis. Briefly, HUVEC were grown in complete media
previously described, to 80% confluence. Prior to stimulation, the HUVEC were washed
with PBS, and then stimulated for 24 h in either unsupplemented basal media (EBM-2
62
devoid of serum and growth factors), unsupplemented basal media containing 2 ng/mL
VEGF, or in unsupplemented basal media containing 300 nM AptDivalent. Following
stimulation, cells were lysed on ice for 15 min using 0.1% triton-X in PBS, containing
10μg/mL protease inhibitor cocktail (Sigma Aldrich). Cellular proteins were resolved by
7.5% sodium dodecyl sulfate–polyacrylamide gel electrophoresis and subsequently
transferred to a polyvinylidene difluoride (PVDF) membrane (Fisher Scientific, Waltham,
MA). After transfer, the membranes were divided with the top portion being
immunoblotted with anti-eNOS primary antibody (1:500, Abcam, Cambridge, UK)
overnight at 4°C. The lower portion of the membrane was probed with anti β-tubulin
primary antibody to serve as a loading control (1:500, Abcam). eNOS and β-tubulin
protein expression was visualized using a horseradish peroxidase–conjugated goat anti-
rabbit secondary antibody (1:5000, Abcam) and colorimetrically detected using the Opti-
4CN HRP substrate detection kit (Bio-Rad, Hercules, CA). The membrane was digitally
imaged and the intensity of the resulting bands was quantified using Image J software
Version 1.46.
Capillary Tube Formation (Angiogenesis)
To assess the function of the divalent aptamer assembly, a capillary tube
formation assay was conducted. Recall that, the divalent aptamer is composed of two
AptM80mer tethered at both ends of a PEG spacer. Therefore, to confirm the effectiveness
of the divalent assembly structure over the monomeric aptamer, the AptM80mer was also
included. Briefly, HUVEC were plated on a 48-well tissue culture plate pre-coated with
250 µl of Cultrex basement membrane extract (BME) (Trevigen, Gaithersburg, MD). The
cells were then treated with 100 µl each of either unsupplemented basal media (EBM-2
devoid of serum and growth factors), unsupplemented basal media containing 2 ng/mL
63
VEGF, unsupplemented basal media containing 300 nM AptM80mer or unsupplemented
basal media containing 150 nM AptDivalent. After 12 h incubation in a 5% CO2 humidified
atmosphere at 37°C, the three-dimensional structures formed by the cells in this matrix
were examined using an inverted phase-contrast microscope and digital images were
recorded. Tube-like structures were further analyzed by measuring the sum of the
lengths all tubules per field using the image analysis software AxioVision, (Zeiss, V4.5)
(Carl Zeiss Microsystems, Thornwood, NY). Four randomly selected low-power fields
were examined for each sample.
Chorioallantoic Membrane (CAM) Assay
Fertilized chicken eggs (30 pieces) were obtained from a local farm (Laughing
Chicken Farm, Gainesville, FL) and placed in trays in a horizontal position in an
incubator maintained at 37 °C with approximately 60–65 % humidity after cleaning with
70 % ethanol. The eggs were incubated for 3 days in horizontal position and were
rotated along their axis twice per day. On day three of incubation, 2 mL of albumen was
aspirated using a syringe needle (25 G) from the transverse end of the egg, to detach
the developing CAM attached to top part of the shell of the egg. The top part of the shell
was also gently poked with the syringe to further help the dropping of the CAM. Then, a
small window of around 1.5 cm2 was gently opened on the top shell with sterile forceps
on the blunt end of the egg without damaging the embryo. Whatman Type I filter paper
was used as carriers for testing angiogenic response. The filter was punched using a ¼
‘paper punch. The paper discs were gas sterilized using EtOH gas and soaked in 100 µl
each of either normal saline solution, saline solution containing 2 ng/mL VEGF, saline
containing 150 nM AptDivalent. The discs were then placed directly on top of the CAM.
The window was then sealed with transparent sterile band-aid to prevent dehydration
64
and possible infections and returned to the incubator. The eggs were incubated for
another 3 days, the covering band-aids were then removed, and the filter paper was
then gently removed. The surrounding vessels of the CAM were then examined under
the dissecting microscope and recorded (Leica M320). The images were then
qualitatively analyzed to assess angiogenic response of the aptamer in comparison to
the controls.
Results and Discussion
Herein the design of a supramolecular aptamer assembly for targeted VEGFR2
binding and activation has been reported. The design of the aptamer assembly is based
upon the native interaction of soluble VEGF with both subunits of surface bound
VEGFR2. When bound, soluble VEGF acts as a bridge between the two subunits and
results in their dimerization. The rationale behind this approach has proven successful
in other surface receptors, such as OX40, a receptor expressed on activated T-cells61.
In this work, an RNA aptamer scaffold was designed with agonist function by binding to
the receptor subunits of the OX40 receptor. In a similar manner, it is hypothesized that
the aptamer assembly described herein allows for the VEGFR2 binding aptamer at each
end of the 18 atom spacer to bind to each of the 2 subunits comprising the VEGFR2
and facilitate dimerization. A schematic demonstrating the proposed mode of action of
the divalent aptamer assembly is shown in Figure 3-2. Receptor agonist activity is
predicated on the ability of the aptamer assembly to first bind the target receptor.
65
Figure 3-2. Schematic representation of the proposed method of action of the divalent aptamer assembly. Binding to target and activation via dimerization, membrane bound VEGFR2.
The Apt80mer used in this work has a high VEGFR2 binding affinity with a Kd of
0.12 nM. (AptSci) However the divalent aptamer assembly is composed of two identical
monomeric units of AptM80mer tethered to each end of an 18-atom hexa-ethylene glycol
spacer. To demonstrate specificity of the Apt80mer as well as the AptM80mer monomer
towards HUVEC, known for high levels of VEGFR2 expression, flow cytometry was
performed on HUVEC probed with FITC conjugated Apt80mer (FITC-Apt80mer) and FITC
conjugated AptM80mer (FITC-AptM80mer). Controls included phycoerythrin (PE) labeled
VEGFR2 specific antibody and PE-labeled isotype control. The results from the study
are shown below in Figure 3-3. The images below show representative histograms from
flow cytometric analysis. The results confirm VEGFR2 expression on HUVEC as
evidenced by the rightward shift (~82% positive) for the VEGFR2 antibody, compared to
~6% positive for the isotype control (Figure 3-3 A and B).
66
Figure 3-3. Aptamer-cell binding interactions. Images A-F show flow cytometry data of aptamer-endothelial cell binding. Un-probed HUVEC are shown in each panel as the blue histogram. The red histogram shows (A) positive binding of CD309 (VEGFR2) antibody; (B) negative binding from the isotype control; (C) positive binding of FITC conjugated Apt80mer aptamer; and (D) inhibition or decreased binding of the aptamer when blocked with soluble VEGF prior to aptamer incubation. (E) positive binding of FITC conjugated modified AptM80mer aptamer (F) inhibition or decreased binding of the aptamer when blocked with soluble VEGF prior to aptamer incubation. Images G-J show flow cytometry data of aptamer-endothelial cell binding in competition with soluble VEGF. Un-probed HUVEC are shown in each panel as the blue histogram. The red histogram in each figure shows positive binding of 150 nM of the fluorescent tagged AptM80mer solution in competition with (G) 50 nM (H) 75 nM (I) 100 nM (J) 150 nM VEGF solutions.
PE -18.0
A
PE +
82.0 PE + 6.59
B
PE -93.4
FITC - 10.6
C
FITC+ 89.4
FITC - 79.5
D
FITC+ 20.5
FITC – 37.5
E
FITC + 62.5
FITC - 84.2
FITC + 15.8
F
G
I
H
J
67
To demonstrate binding specificity of the Apt80mer and AptM80mer towards the
VEGFR2 target receptor, HUVEC were incubated with Apt80mer and AptM80mer with and
without pre-treatment with soluble VEGF. The flow cytometry data demonstrate ~89%
positive targeted binding of Apt80mer to VEGFR2 compared to the diminished ~20%
positive Apt80mer binding when pre-treated with soluble VEGF (Figure 3-3 C and D).
Similarly, the data also showed ~62.5% positive targeted binding of AptM80mer to
VEGFR2 while the cells pre-treated with soluble VEGF showed diminished binding
(~15.8% positive) (Figure 3-3 E and F). This data demonstrates successful receptor-
ligand binding, which is the critical first step in the complex process of membrane
receptor activation157. In addition, the cells pre-incubated with VEGF showed diminished
aptamer binding, which indicates specificity of the binding aptamer to VEGF receptor.
Finally, the assessment of competitive binding of the aptamer in the presence of
VEGF at varying concentrations showed that VEGF did not interfere with the aptamer
binding at concentrations ranging from 50 to 150 nM (Figure 3-3 G-J). Also, preliminary
data from a separate experiment conducted, indicated that over a period of 20 h, the
bound monomeric aptamer (AptM80mer) (FITC tagged) is internalized by the cell. This
early finding is consistent with reports in the literature that showed internalization of
aptamers by the target cell50,57. Further investigation is required to study the specific
localization of the aptamer assembly within a cell.
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Figure 3-4. Aptamer Induce VEGFR2 activation and downstream molecular events (A) Quantification of HUVEC VEGFR2 phosphorylation upon exposure to either unsupplemented basal media (BM), or BM supplemented with 100 ng/ml soluble VEGF, 150 nM soluble divalent Aptamer (AptDivalent) or 300 nM Monomer (AptM80mer). Data shown are Mean ± SD; Images B and C show activation of Akt signal transduction pathway. HUVEC are shown in each panel as the blue histogram. The red histogram shows positive binding of Anti-Phospho Akt (Ser473) antibody, thus indicating pathway activation upon stimulation with (B) 2 ng/ml soluble VEGF; and (C) 300 nM AptDivalent. Image D represents eNOS protein expression from HUVEC treated with either unstimulated basal media (BM), or BM supplemented with 2 ng/ml soluble VEGF, or 300nM soluble divalent Aptamer (AptDivalent). The top membrane shows eNOS specific protein bands and the bottom membrane shows beta tubulin specific bands that were used as a loading control. eNOS protein band
intensity was normalized to β-tubulin and quantified using image J software
version 1.46. In figures A and D, based upon One-way ANOVA with Tukey’s multiple comparison test, N.S. indicates no significant difference (p>0.05), and * asterisks indicate a statistically significant increase (p<0.05).
The binding of soluble VEGF to VEGFR2, initiates a sequence of events resulting
in receptor dimerization, kinase activation, and auto-phosphorylation of specific tyrosine
kinase residues within the dimeric complex158. The activation of tyrosine kinase through
C D
69
trans-phosphorylation between receptor molecules is the first step leading to activation
of intracellular signal transduction pathways157. Therefore, to assess the activation of a
cell surface receptor, assessment of phosphorylation events is important. Towards this
goal HUVEC were treated with basal media supplemented with either the AptDivalent,
AptM80mer, or soluble VEGF and assessed for VEGFR2 receptor phosphorylation. As
expected, treatment of HUVEC with soluble VEGF results in activation of VEGF
receptor-2 phosphorylation events (Figure 3-5 A).
The treatment of HUVEC with AptDivalent also results in phosphorylation of
VEGFR2 (Figure 3-4 A). Further, treatment of HUVEC with the divalent assembly
results in significantly greater phosphorylation relative to the amount of phosphorylation
from basal media or basal media containing monomeric AptM80mer (Figure 3-4 A,
p<0.05). These data reaffirm the specificity of the divalent aptamer assembly and
demonstrating the promotion of VEGFR2 phosphorylation events on HUVEC.
More importantly these data demonstrate the unique advanced ability of the
divalent assembly in receptor activation that is not seen in the monomeric form. In fact,
despite its ability to bind VEGFR2, the addition of the monomeric aptamer (AptM80mer),
from which the divalent assembly is composed, resulted in no difference in
phosphorylation over unsupplemented basal media alone (Figure 3-4 A; p>0.05). The
induction of receptor phosphorylation via an aptamer assemblage shown here
represents one of the early and few examples in the literature of receptor activation via
an aptamer complex.
The downstream events in VEGF receptor activation are quite complex, involving
multiple signal molecules and multiple molecular pathways, including the well
70
characterized Akt pathway39. To further assess the effect of aptamer-receptor binding,
the activation of this downstream molecular pathway was examined. Akt pathway
activation is involved in numerous diverse biological processes, such as cell
survival/apoptosis, cell cycle control, angiogenesis, differentiation, and cell growth and
proliferation159,160. In addition, Akt pathway activation has been associated with
activation of VEGFR2, through treatment with soluble VEGF161,162. In this study, serum
starved HUVEC were treated with AptDivalent or soluble VEGF, then phosphorylation of
Serine 473 is detected with a phosphor-Akt antibody and analyzed via flow cytometry.
Data reveal that treatment with the AptDivalent results in activation of the Akt pathway in
levels similar to that seen with soluble VEGF (Figure 3-4 B and C). The cells treated
with either AptDivalent or soluble VEGF showed positive staining for phosphor Akt, ~97%
and 93% respectively. Although the activation of Akt pathway by soluble VEGF is
expected and has been shown by others, there are currently no reports of Akt pathway
activation triggered by treatment with a DNA aptamer, nor an aptamer assembly. These
promising results provide early evidence that the AptDivalent assembly as designed
promotes activation of VEGFR2 and causes downstream activation of pathways that
control critical cell functions.
eNOS is an important protein involved in the production of Nitric oxide in
endothelial cells, and is critical for regulation of cardiac function and angiogenesis. The
protein expression experiments reveal aptamer induced upregulation of eNOS to levels
comparable with what is seen in VEGF induced upregulation (Figure 3-4 D).
Quantification of the protein levels shows a statistically relevant increase in eNOS
expression in HUVEC stimulated with VEGF or the divalent aptamer assembly. These
71
results are encouraging and indicate a potential angiogenic downstream effect of
receptor activation.
As a final functional assessment of the agonist behavior of the AptDivalent, a tube
formation assay was utilized. The tube formation assay is well characterized and
regarded as a powerful tool to screen for various factors that promote or inhibit in vitro
angiogenesis. In addition, angiogenesis as measured in the tube formation assay, is
mediated by VEGF signaling through activation of VEGFR2163. The soluble AptDivalent
promotes HUVEC migration and capillary like tube formation (Figure 3-5 A-D). When
compared to HUVEC stimulated with soluble VEGF, cells stimulated with AptDivalent
shows statistically similar tube length measurements (Figure 3-5 B-C; p>0.05). The data
also indicates that the AptDivalent is more effective in induc ing capillary tube formation
than the AptM80mer (Figure 3-5 E; p<0.05). These data are particularly impactful as this is
the first report of a divalent aptamer assembly activating cellular processes involved in
promoting angiogenesis. In addition, this data also show the potency of the divalent
assembly to act as a receptor agonist while the monomeric form (AptM80mer) shows very
little or no receptor phosphorylation and downstream angiogenic effect, despite
exhibiting excellent binding affinity towards VEGFR2. These data reaffirm the rationale
behind the design of the divalent aptamer assembly. Although many robust in vitro
assays have been developed to assess angiogenic potential of various biopolymers as
well as natural substances, the process is so dynamic that it is best characterized by in
vivo studies. Although, animal studies using a mammalian subject is the golden
standard for such experiments, these are often very expensive and time consuming.
72
Figure 3-5. Functional assessment of VEGFR2 aptamer binding. Representative phase contrast images of HUVEC cultured on matrigel treated with either (A) unsupplemented basal media (BM); (B) BM supplemented with 2 ng/ml soluble VEGF; (C) 150 nM soluble divalent Aptamer (AptDivalent); or (D) 300 nM Monomer (AptM80mer). Images in each panel show tube like structures formed. Quantification of tube length in µm was taken from 4 separate low powered magnification images using AxioVision, v.4.5 software. Data represent the Mean ± SD of between 18-75 measurements taken from each of 4 images. Based upon One-way ANOVA with Tukey’s multiple comparison test, N.S. indicates no significant difference (p>0.05), and * asterisks indicates a statistically significant increase (p<0.05).
The chorio allantoic membrane assay (CAM) assay ever since its development
has been widely used as an in vivo model to study angiogenesis. It has many
advantages over typical mammalian systems including low cost, and ease of
experimentation, as well as the absence of a mature immune system164,165. In this study,
C D
E
A B
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the CAM assay was used to determine the angiogenic potential of the divalent aptamer
assembly in comparison with VEGF. Saline was used as a control in this study as a
blank control. Preliminary data as shown through digital images in Figure 3-6 of the
CAM after 3 days of incubation with samples and controls indicated that the aptamer
assembly is able to induce angiogenic response comparable to VEGF.
Figure 3-6. Chorio allantoic membrane assay. Digital images of the CAM after 3 days of incubation with samples. The dotted circle marks the position of sample loaded filter paper. The arrows show neoangiogenic sprouting of vessels in the developing CAM.
The response was assessed via appearance of neovasculature around the area
where the sample was placed as shown in the images.The data presented here
demonstrate the ability of an aptamer assembly to act as a receptor agonist for
VEGFR2 on human endothelial cells. Further, the function of the aptamer to promote
endothelial cell capillary like tube formation, an important criterion demonstrating
functional agonist activity was shown. As an extension of this behavior, the CAM assay
also showed preliminary evidence of the aptamer assembly acting as a pro-angiogenic
factor in an in vivo model165. This mechanism of utilizing an aptamer agonist to
modulate receptor response and activate downstream signaling pathways in endothelial
cells has not been shown previously in the literature.
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Summary
Controlling and modulating receptor mediated response in cells is of paramount
importance in many areas such as tissue engineering, regenerative medicine,
diagnostics, etc30,33. Due to the small subcellular scale of the interactions involving
receptor mediated processes, strategies to invoke favorable cellular response via
cellular receptors are often developed using biomolecules at the nanoscale (<100 nm).
Traditional strategies for such processes historically have involved the use of cytokines
and growth factors since these represent the natural mechanism of receptor activation
in the body31,45. However, these approaches have been shown to cause nonspecific
response when used as a therapeutic means. Therefore, it is imperative to identify and
develop strategies that are targeted and specific for controlling receptor mediated cell
processes. The present work hypothesizes the use of a divalent aptamer assembly for
activation of VEGFR2 receptor in endothelial cells. This work can further be extended to
develop aptamer based receptor agonists to control other cell mediated processes such
as differentiation, proliferation, apoptosis etc. This novel strategy also has tremendous
potential for the development of novel pharmaceuticals in the treatment of a wide range
of diseases, including inflammation, autoimmune diseases, and vascular disease.
Activation of the VEGFR2 receptor has been shown to facilitate various processes such
as proliferation, migration, neoangiogenesis and differentiation in endothelial cells and
various stem cell populations. These positive outcomes of the activation event have
been exploited in designing pro-angiogenic therapies via targeted activation of VEGFR2
for vascular tissue engineering to cause repair and regeneration of vascular
tissue155,163,166.
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The results presented in this study show that this novel strategy of using nucleic
aptamers based divalent assembly to target VEGFR2 was effective in triggering the
activation cascade in these receptors in endothelial cells. The assembly was also able
to cause positive downstream angiogenic response which proves that this is a
promising strategy to modulate cell behavior favorably for vascular tissue engineering
and regenerative medicine, where controlling cell processes is of paramount
importance. The ability to modulate a functional response such as cell growth,
differentiation and angiogenesis is critical for success in these areas37,166. Therefore,
this novel, specific and targeted strategy has tremendous potential towards designing
novel nanoscale biomaterial cues for vascular tissue engineering.
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CHAPTER 4 MESOSCALE STRUCTURAL CUES- PART 1- DEVELOPMENT OF A BIOMIMETIC,
BILAYERED VASCULAR SCAFFOLD FOR TISSUE ENGINEERING SMALL DIAMETER BLOOD VESSELS
Introduction
In Chapter 3, the response of sub-cellular and molecular cues was addressed.
This chapter describes strategies to modulate cellular response by designing
appropriate cues on the mesoscale. Important cellular responses and characteristics
including alteration of cell shape, alignment, differentiation, growth and tissue formation
have been known to be modulated by architectural cues that are presented to the cell
systems by biomaterials that are in contact with them14,30,167. These structural stimuli
may be provided to them across multiple length scales. Each stimulus can be designed
to elicit a specific response from the cell system. This concept is made use of while
designing biomaterials for tissue engineering.
The biological mesoscale range has been described to be inclusive of biological
structures ranging from 10 to several 100 microns in size. Although there have been
some discrepancies in the numerical definition of the scale size itself, mesoscale is
broadly used to describe phenomenon occurring in cellular microenvironments at a
relatively larger scale than the sub-cellular level. This includes interaction of cells with
surrounding biomaterials14. Proper understanding of these interactions is essential
especially in fields such as tissue engineering and regenerative medicine. Manipulating
the cell environment using artificially engineered matrices is a primary focus in this field
of science15,168. Biomaterial environments provide a variety of physical, mechanical,
chemical as well as structural cues that can be altered to suit a specific application.
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Cardiovascular biomaterials in general encompasses a wide variety of material
applications ranging from vascular grafts to hemodialysis tubes. These biomaterials can
be designed to alleviate complications arising due to different cardiovascular diseases,
such as artherosclerosis, myocardial infarction, peripheral artery disease etc. The
current chapter and Chapter 5 will discuss two classes of cardiovascular biomaterial
applications, namely, small diameter vascular grafts and vascular stents, respectively.
This work is focused on studying the interaction of these biomaterials with in vitro
models of the vascular environment in order to understand and develop strategies to
repair damaged vasculature and enable regeneration.
With this goal in mind, the following aim was established and studied. Within the
broad scope of this aim, specific sub-goals were also listed to study the two classes of
biomaterial applications as described above.
Specific Aim 2
Understand and develop biomaterial based strategies towards vascular tissue
engineering and regeneration by studying the interactions of biomaterials with vascular
microenvironments.
Sub-aim 1
Develop a novel biomimetic, bilayered, biomaterial construct and evaluate its
applicability as a small diameter vascular graft.
Sub-aim 2
Evaluate the biocompatibility and hemocompatibility of an Mg-Ca-Sr alloy to
assess it suitability for use in designing vascular stents.
As mentioned above, this chapter will deal with the design and development of a
biomimetic vascular scaffold for small diameter blood vessel engineering. The field of
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vascular tissue engineering aims to create replacement vascular tissue for diseased or
damaged segments of the vasculature. The end goal of this research is to restore
adequate blood flow to regions of the body where blood flow is lacking because of
damage or diseased state of the existing vasculature, or in some cases, the complete
destruction of the vascular network6,8. Current efforts in the field of vascular tissue
engineering are focused on creating functional replacements for small diameter blood
vessels (<6mm). Furthermore, the golden standard for vascular grafts is the use of an
autologous vessel that would be biologically acceptable for the patient. But the
availability of autologous vessels is limited due to pre-existing conditions or failure to
identify or extract a vessel of appropriate size with minimal damage to the existing
system. These clinical limitations have steered research in this field towards developing
synthetic tissue engineered vascular grafts6,8,117. Current grafts are not completely
successful and often fail due to complications after implantation, such as mechanical
failure, thrombosis or increased inflammation. An ideal vascular graft should be
nonthrombogenic or thromboresistant, non-toxic, cause minimum inflammatory
response, ease of processing and handling, and mechanically compatible with existing
vasculature75. Regardless of whether the material used to create the vascular graft is
synthetic or natural, it is also essential that the scaffold can support vascular cell
adhesion and proliferation to enable reintegration into the vascular network. Although
there have been decades of advancement in vascular cell biology, there remains the
challenge of identifying a suitable material and an ideal graft design. Recent
approaches aim to achieve acceptable patency rates, through the development of grafts
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that best mimic or promote the extracellular environment and mechanical properties of
native blood vessels4,169.
The identification of a single material or a combination of materials for fabrication
of grafts has been a pivotal research focus in the field. A popular and widely practiced
approach is the use of proteins derived from plant and animal sources as the base
material for scaffold18,79,89,170,171. However, this method is limited by the low mechanical
strength of the scaffolds, high cost of proteins as well as difficulty in processing them.
Historically, many synthetic materials have been identified as suitable materials in
vascular graft design due to the ease and flexibility of tailoring their mechanical and
structural properties. Due to the poor patency rates of most synthetic material grafts,
recent efforts focusing on the use of biodegradable polymers as scaffolds to improve
the clinical viability have been made. In this design, the scaffold will degrade and be
replaced with extracellular matrix proteins that are secreted by the cells over time.
Some common materials that have been investigated include polyglycolic acid, poly-L-
lactic acid, and co-polymers of the two70,77,109. However, synthetic materials have
increased risk of thrombogenicity, toxicity, mechanical failure and inflammatory
response. Novel strategies including surface functionalization to impart greater anti-
thrombogenic properties of the material, via coatings, chemical and protein
modifications, and cellular seeding on these otherwise inert materials, have also been
attempted172–175. But the long-term toxicity as well as anti-thrombogenicity of purely
synthetic grafts are still not completely verified. These challenges have directed
research towards novel hybrid material systems that combine the positive
characteristics of natural as well as synthetic materials. This study will investigate the
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use of a combination of a biodegradable, elastomeric polymer; poly (1,8 octanediol-co-
citrate) (POC) with natural extracellular matrix proteins; collagen and elastin. POC
belongs to a family of elastomers composed of citric acid and an aliphatic diol. POC has
been identified as a hydrophilic and biodegradable polymeric constituent that can
substitute for elastin in synthetic fibres.80–83 It is also anti-thrombogenic and supports
differentiated endothelial progenitor cell adhesion growth.84,85
This polymer-protein mixture was fabricated into a fibrous construct using
electrospinning. Electrospinning offers robust material selection, low cost, simplicity,
high surface-to-volume ratios, and controlled porosity for vascular scaffold fabrication.
With these advantages, electrospinning is a proven framework allowing for the improved
endothelialization of vascular grafts89,96,170. While some encouraging data using
synthetic electrospun polymers have been obtained, these approaches are still limited
and some of the challenges pertaining to their lack of bioactive functionality are still
present97. Electrospinning of POC with natural co polymers and their suitability as
scaffolds for vascular grafts has been successfully demonstrated before81. This study is
focused on electrospinning POC with collagen to closely mimic the architecture of the
native vessel wall.
Apart from selection of appropriate biomaterials, microstructure as well
dimensions of the scaffold are an important criterion for success14,27,176. Current efforts
have focused on mimicking the native structure of the arteries to improve cellular
compatibility of the constructs by providing essential structural cues. The design of the
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construct in this study is of a scaffold structure consisting of two layers as shown in the
illustration below (Figure 4-1).
Figure 4-1. Schematic representation of the bilayered vascular graft.
Native artery as shown in Figure 2-3, consists of a solid inner layer known as the
lumen and a fibrous exterior layer also known as the medial layer. The inner luminal
layer is composed primarily of a monolayer of endothelial cells that are in direct contact
with circulating blood. The medial layer, consists of multiple layer of smooth muscle
cells. This layer is also rich in collagen and elastin among other ECM proteins and
provides a great amount of mechanical support to the vessel91,177.
The goal of this study was to mimic this structure of as closely as possible.
Therefore, the design as described above was chosen to support a monolayer of
endothelium and an interconnected multi-layer of SMCs. This report describes the
development of a bilayered, biomimetic scaffold. This strategy is currently being
developed to recruit stem cells to improve regeneration and reintegration into the native
vasculature.
Materials and Methods
Materials
POC pre- polymer was synthesized using a protocol reported by Yang et al. 178.
Briefly, equimolar ratios of 1, 8 octanediol (Sigma-Aldrich, St. Louis, MO) and citric acid
(Sigma-Aldrich, St. Louis, MO) were blended together and melted at 165˚C. Once the
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monomers were completely molten, the poly condensation reaction was carried out at
140˚C. The pre-polymer solution was then purified reconstituted at concentration of 50
wt.% in ethanol and 10 wt.% in 1,1,1,3,3,3-hexafluoro-2-propanol (HFIP) (Sigma-
Aldrich, St. Louis, MO) for lumen solvent casting and electrospinning respectively.
Lyophilized collagen type I from calf skin (Elastin Products Company, Inc., Owensville,
MO) and soluble elastin (Elastin Products Company, Inc., Owensville, MO) were used to
make the protein-polymer mixture for electrospinning. Polyacrylic acid (PAA) powder
was obtained from Sigma Aldrich and was reconstituted to 10 wt.% in ethanol.
Solid Lumen Fabrication
To fabricate a solid lumen consisting of POC, First, a 5 wt. % Polyethylene oxide
(Peo) solution in water was electrospun on an aluminum rod. The polymer solution was
fed through a 25G needle at 2 mL/ hr with a syringe pump (Chemyx Fusion 100,
Stafford, TX). A voltage potential of 25 kV was applied and the rod was placed 15 cm
away from the syringe tip. The rod was rotated at a speed of 500 RPM to ensure
uniform distribution of fibers The Peo spun rod was dipped in a 50 wt. % solution of
POC pre-polymer in ethanol and the solvent was evaporated overnight in air at room
temperature. To prevent irregular flowing of the pre-polymer from the rod while
subsequent fabrication steps are carried out, the POC coated rod was then partially
crosslinked at 60˚C and 80˚C for one day each.
Fabrication of Electrospun Medial Layer
10 wt. % solutions of Collagen, POC and Elastin in HFIP were prepared for
electrospinning. The collagen: POC: elastin solution was made with a 60:20:20 weight
ratio of each of the components respectively. The mixture was vortexed for 15 min
under low speed to thoroughly mix the constituents. The collagen: POC: elastin solution
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was fed through a 25G needle at 1.25 mL/ h with a syringe pump and a voltage
potential of 22-25 kV was used; the working distance was set at 15 cm.
Glutaraldehyde Crosslinking
The electrospun scaffolds were then crosslinked using glutaraldehyde vapor. The
scaffolds were crosslinked immediately after electrospinning to stabilize the proteins in
the scaffold and to improve mechanical strength of scaffold. The rods were suspended
in a set up as shown in schematic below (Figure 4-2 A). The cylinder contained 30 ml of
glutaraldehyde (25% in water). For flat sheets, a similar set up was shown where the
scaffold is placed in an enclosed setting (Figure 4-2 B). The scaffolds were crosslinked
for 2, 4, 6 and 8 h. After crosslinking the scaffolds were soaked in water to wash away
any residual glutaraldehyde and to enable demolding of the bilayered construct
Figure 4-2. Glutaraldehyde vapor crosslinking. (A) Electrospun rod crosslinking set up (B) Electrospun mat crosslinking set up.
A
B
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FTIR Characterization
The crosslinking of scaffolds was confirmed by FTIR. Spectro chemical analysis
of the electrospun scaffolds crosslinked at various times was performed using
attenuated total reflectance Fourier transformation infrared spectroscopy (ATR-FTIR).
ATR-FTIR spectra was obtained for each of the samples using the Nicolet 6700 FTIR
Spectrometer (Thermo Scientific) and a diamond tip window. The spectra were read
over a range of 600-4000 cm-1 for each of the spectra and a total of 32 scans were used
with a resolution of 4 cm-1.
SEM Characterization
The scaffolds were washed and frozen at -80˚C for 18 h before being lyophilized.
The dried scaffolds were then mounted for SEM analysis and the nanofibrous structure
of the scaffolds were observed using a table top SEM (PhenomWorld ProX, Eindhoven,
Netherlands). Fiber size distribution was quantified from SEM images using ImageJ
(National Institutes of Health Bethesda, MD).10 images taken at 10000 X magnification
were used and 15 measurements were taken from each of the images. The average of
the fiber size measurements was reported and a distribution was plotted.
Contact Angle Measurement
The contact angle of the solid POC lumen as well and the collagen: POC: elastin
layer were characterized using a custom goniometer (courtesy: Brennan lab, UF). Solid
films of POC and the 60:20:20 collagen: POC: elastin were fabricated by solvent casting
the polymer solutions on glass slides and solvent was allowed to evaporate overnight.
The films were crosslinked using the conditions reported earlier. The glass slides were
placed on the custom goniometer and a droplet of 5 µL of deionized water was placed
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on the film. This was repeated three times for each of the films. The contact angle
observed was averaged and reported and correlated with hydrophilic behavior.
Uniaxial Tensile Testing
Mechanical properties of the electrospun protein polymer construct was
determined using a tabletop uniaxial testing machine (INSTRON 3345) equipped with a
10-N load cell. A cross-head speed of 10 mm/min at room temperature was applied to
hydrated samples after being punched into dog bone shaped specimens. A sample size
of at least 5 was used for the testing procedure.
Biocompatibility Characterization
Cell culture
Human umbilical vein endothelial cells (HUVECs) (Lonza, Walkersville, MD) and
human aortic smooth muscle cells (hASMCs) (Cell Applications Inc., San Diego, CA)
were used as representative mature vascular cell populations to evaluate the response
of the different layers on the construct with relevant vascular cells. Both cell types were
cultured in a cell culture incubator maintained at 37˚C, 5% CO2. HUVECs were cultured
in endothelial growth media-2 (EGM-2) (Lonza) and were used between passages 3-5
and while the hASMCs were cultured in smooth muscle cell growth media (SmGM) (Cell
Applications Inc) and used between passages 8-12.
Scaffold preparation for cell studies
For biocompatibility evaluation, the scaffolds were punched into discs of diameter
(1/16”) following fabrication. The scaffolds were then gas sterilized with ethylene oxide
gas prior to use, in a gas sterilizer (Anaprolene AN741, Vernon Hills, IL). Unreacted
monomers and residual glutaraldehyde was leached from the scaffold using Dulbecco’s
Modified Eagle Medium (DMEM) (Corning Cell Grow, Mediatech Inc., Manassas, VA)
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for one week with constant replacement of medium. The scaffolds were then washed
with 1x phosphate buffered saline (PBS) and then incubated with cell culture medium
overnight before cell seeding.
Cellular attachment
To assess cellular attachment and spreading, appropriate cells were seeded on
the solid lumen layer and the electrospun medial layer separately. HUVECs were
seeded at a density of 5000 cells/mm2 on the POC discs and hASMCs were seeded at a
density of 5000 cells/mm2 on the collagen: POC: Elastin discs respectively. The cells
were seeded by placing a 50 µl cells suspension directly on the scaffold surface.
Special care was taken to avoid any spillage. The seeded scaffolds were then carefully
placed in a cell culture incubator and allowed to attach for 45 minutes. Following this
incubation, scaffolds containing wells were filled with about 300 µl of media. Cells were
then cultured on the scaffolds for two days. After two days of culture, scaffolds were
rinsed twice with PBS to remove any unattached cells, fixed with 2% glutaraldehyde.
The fixed scaffolds were the serially dehydrated with graded series of ethanol (30%,
50%, 70%, 90%, 95% and 100% ethanol in water) by incubating the scaffold in each
solution for 10 min. Scaffolds were then lyophilized, mounted, and coated with Au-Pd
via sputter coating for SEM analysis.
Lactate dehydrogenase cytotoxicity assay
Preliminary cytotoxic response of the scaffolds was evaluated by quantifying the
concentration of lactate dehydrogenase (LDH) released in cell culture supernatant
media. LDH cytotoxicity assay kit (Thermo Scientific, Waltman, MA) was used to
colorimetrically detect the concentration of LDH in the media. Scaffolds were placed in
24 well ultra-low attachment plates (Fisher Scientific, Waltman, MA). Briefly, 5000
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HUVECs were seeded onto the POC layer and 5000 hASMCs were seeded on the
fibrous scaffold layer cultured in 96 well plate with 100 µl of media in each well. The
optimal cell density for LDH release was previously assessed using a cell density vs
LDH release calibration curve. After two days of culture, the media was transferred into
a 96 well plate and assayed for LDH concentration. A subset of wells was treated with
lysis buffer provided to serve as a control for maximum LDH concentration. The relative
percentage cytotoxicity was reported per scaffold.
Cell proliferation assessment
The assessment of proliferation of cells on the medial layer consisting of
collagen: POC: Elastin, was studied using a 3-(4,5-dimethylthiazol-2-yl)-2, 5-diphenyl
tetrazolium bromide (MTT) assay (Trevigen, Gaithersburg, MD). The assay is designed
to spectrocolorimetrically indicate the metabolic activity of the cells through the
assessment of cellular ability to reduce the yellow tetrazolium salt to purple formazan
crystals by mitochondrial enzymes. Healthy proliferating cells can produce a greater
amount of crystals when compared to quiescent cells or cells undergoing apoptosis. For
proliferation assessment, cells were seeded as described in the LDH assay section
above. The initial cell seeding density was determined to be 2500 cells/mm2. Each well
was filled with 100 µl of cell culture medium. At time points of 1, 3, 5, and 7 days, 10 µl
of MTT reagent was added to media, allowed to react for 4 h, following which 100 µl of
detergent was added to solubilize the crystals produced overnight. The supernatant was
collected and the absorbance was measured using a plate reader. The absorbance
values were correlated with cell number using a previously determined standard curve
that correlated absorbance to cell number. The cell number versus culture time was
plotted.
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Optimizing Bilayer Fabrication with Electrospun POC-PAA
Once the biocompatibility and cellular compatibility of the electrospun scaffold
was confirmed. As the next step was focused on the fabrication of the bilayered
construct. The fabrication of the bilayer was optimized using an established
electrospinning system consisting of 10 wt. % POC: PAA (50:50) solution in ethanol.
This system has already been characterized and reported for wound healing
applications by the lab previously. The POC: PAA solution was electrospun on the
partially crosslinked lumen on the aluminum rod. The polymer solution was fed through
a 25 G needle at the rate of 2 mL/h. The collector plate distance was fixed at 20 cm and
the rod was rotated at 500 RPM. The fabricated construct was thermally crosslinked at
80˚C for 2 days and 100˚C for one day. The scaffolds were demolded from the rods by
soaking in water. The integration between layers as well as the fabrication of two
distinct phases was confirmed visually using SEM analysis of cross-sections of the
scaffold.
Fabrication of Bilayered Construct
To fabricate a bilayered construct, a layer of collagen: POC: elastin was spun on
a rotating mandrel consisting of the partially crosslinked lumen as described above. The
final biphasic scaffold was crosslinked in glutaraldehyde vapor for 6 h as determined
earlier. Once crosslinking was completed the scaffold was placed in deionized water for
two days then removed from the mandrel. The construct was also characterized using
SEM analysis as described above.
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Mechanical Properties
Compliance
Dynamic compliance of the scaffolds was measured using a custom bioreactor
apparatus (42 Bio, Gainesville, FL). Scaffolds were mounted on the metal adapters of
the custom bioreactor apparatus using UV glue, Loctite 3942 (Henkel Adhesives, Rocky
Hill, CT). Compliance was measured per ISO 7198 section 8.10 defined as the change
in volume divided by the change in pressure at the physiological diastole and systole
(approximately 120/ 80 mmHg). Using an LED micrometer (Keyence LS7000, Itasca, IL)
the change in outer diameter was measured. The following equation provided in the
ISO, Equation 4-1 was used to calculate the % compliance of the scaffolds.
% 𝑐𝑜𝑚𝑝𝑙𝑖𝑎𝑛𝑐𝑒 =
𝑅𝑝2−𝑅𝑝1
𝑅𝑝1
𝑝2−𝑝1× 104 (4-1)
Where Rp2 represents the radius at p2, the diastolic pressure and Rp1 represents the
radius at p1, the systolic pressure.
Circumferential hoop stress
Circumferential hoop stress was measured per ISO 7198 section 8.3. The
biphasic scaffolds were cut into ringlets, hydrated for 1 h, and loaded into stainless steel
hooks in a uniaxial mechanical testing apparatus (Instron 5942, Norwood, MA). Ringlet
width, wall thickness, and outer diameter were measured using a caliper for each
sample. A minimum of three measurements was taken per sample. The cross-head
speed was set at 50 mm/min, applying tensile force until failure of the scaffold was
observed. Load and displacement data were collected to determine the elastic modulus
of the scaffolds, using Equation 4-2 as shown below.
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𝐸 =𝜎
𝜀 (4-2)
The burst pressure of the scaffolds was extrapolated from the circumferential
hoop stress using the Laplace law which states burst pressure increases linearly with
decreasing diameter if the wall thickness is assumed to be constant. Equation 4-3,
where Pburst was the burst pressure, σy was the yield stress, t was wall thickness, and r
was the radius of the scaffold was used to calculate extrapolated burst pressure.
𝑃𝑏𝑢𝑟𝑠𝑡 =𝜎𝑦×𝑡
𝑟 (4-3)
Statistical Analysis
All numerical data are reported as average ± standard deviation. Statistical
significance was determined using one-way analysis of variance test with Fisher’s LSD
as the post hoc analysis. P values < 0.05 were reported as significant while P> 0.05
were reported as statistically similar.
Results and Discussion
Despite immense efforts, the search for a fully functional scaffold system that
enhances vascular regeneration is still a major focus of this field. One of the major
challenges has been the identification of a suitable material to develop small diameter
vascular graft replacements. Historically, many synthetic as well as natural materials
have been utilized for the fabrication of vascular grafts. Additionally, various changes to
the processing conditions, surface properties and material geometry have also been
investigated towards the development of patent and compliant vascular grafts for small
diameter vessel applications. Even though natural materials, are most suited for these
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applications due to their bio functionality, the weak mechanical properties as well as the
high cost and difficulty in processing limits their application. Synthetic materials despite
having high mechanical properties have limited bioactivity. Therefore, the use of a
hybrid protein-polymer system has been evaluated in this study.
The overall goal of this study to fabricate a bilayered biomimetic scaffold for
potential application as a vascular graft. The rationale behind this approach is to
develop a scaffold system that mimics architectural and functional aspects of an existing
arterial environment to promote vascular cell attachment and growth. The design as
shown in Figure 4-1 consists of solid inner layer consisting of POC and a fibrous
electrospun exterior layer consisting of POC combined with ECM proteins (collagen and
elastin). POC belongs to an elastomeric family of synthetic polymers, termed poly (diol
citrates) (PDC) composed of citric acid and an aliphatic diol. The condensation reaction
involved in the synthesis of polydiol citrates is shown below in Figure 4-3.
Figure 4-3. Synthesis of poly diol citrates.
Amongst these elastomeric polymers poly (1, 8- octane diol) citrate (POC) has
been identified as a hydrophilic and biodegradable polymeric constituent that can
substitute for elastin in synthetic fibers. It is also anti-thrombogenic and supports
endothelial progenitor cell adhesion growth which is required for the blood contacting
lumen of the graft83,179. Therefore, this material was chosen to form the base lumen
layer to support growth of a neoendothelium. This layer was fabricated by solvent
Citric acid Poly diol
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casting a viscous solution of POC in ethanol and partially crosslinking the layer to
preserve the structure of the lumen during subsequent processing stages.
The idea behind fabrication of the bilayered scaffold was to fabricate by solvent
casting a solid POC lumen onto a mandrel and electrospinning a nanofibrous collagen:
POC: elastin layer on top of the lumen. Towards this goal, an electrospun composite
consisting of collagen, POC, elastin was first fabricated. Electrospinning offers robust
material selection, low cost, simplicity, high surface-to-volume ratios, and controlled
porosity for vascular scaffold fabrication. With these advantages, electrospinning is a
proven framework allowing for the improved reintegration of vascular grafts. While some
encouraging data using synthetic electrospun polymers have been obtained, these
approaches are still limited and some of the challenges pertaining to their lack of
bioactive functionality are still present97. Electrospinning of POC with natural co
polymers and their suitability as scaffolds for vascular grafts has been successfully
demonstrated before79,81,180–182. Natural ECM proteins, collagen and elastin have been
known to be present in the ratio of 80:20 in native arteries177,181,183. This ratio has been
adopted while designing collagen and elastin based vascular scaffolds68,184,185. In this
study, the idea is to substitute a part of the collagen for elastomeric POC to improve the
mechanical robustness and compliance of the scaffold. Therefore, collagen: POC:
elastin was mixed in a ratio of 60 : 20 :20 by weight and dissolved in a common solvent
HFIP to electrospin the mixture.
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The electrospun mixture produced uniform fibers of size 0.816 nm as shown in
Figure 4-4 below and the electrospun scaffolds were water soluble and had to undergo
chemical crosslinking with glutaraldehyde vapors to stabilize the proteins.
Figure 4-4. SEM characterization as spun collagen: POC: elastin fibers. (A) SEM image of as spun Collagen POC Elastin fibers, 5000X magnification (B) Fiber size distribution analysis.
Glutaraldehyde has been used historically as a strong crosslinking agent to
crosslink and stabilize proteins186. It reacts with the amino groups present in the
proteins to form Schiff base intermediates which then link further to form a large
crosslinked network. Glutaraldehyde crosslinking has been shown to be a cost effective,
fast and an efficient process which increases the mechanical strength and denaturing
temperature of the crosslinked proteins. However, some studies have shown that this
process can leave behind unreacted aldehydes, which when released causes local
toxicity to surrounding tissue187–189. Therefore, care needs to be taken to remove
unreacted aldehydes released by this process.
A B
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In this study, the scaffolds were crosslinked at various crosslinking times to
determine an optimal crosslinking time and their stability was preliminarily tested by
observing their behavior when immersed in water. Scaffolds crosslinked for shorter
times, less than 1 h seemed to readily dissolve in water. Therefore, in this study,
scaffolds were crosslinked for 2, 4, 6 and 8 h. The crosslinked scaffolds were all stable
in water and retained their structural characteristics of crosslinked fibers as shown in the
Figure 4-5 below.
Figure 4-5. SEM characterization of all crosslinking conditions using glutaraldehyde vapor. All images are at a magnification of 5000 X.
Further FTIR analysis confirmed the crosslinking of scaffolds as indicated by a
characteristic Amide II (NH) bending vibrations peak observed at around 1550 cm-1.
Additionally, Amide I (C=O stretching) and Amide A (N-H stretching) signatures were
also seen respectively at 1632- 1664 and 3320-3340 cm-1 for all scaffolds tested. As
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time of crosslinking increased, an increase in peak absorbance was observed in the
Amide bands indicating increase in crosslinking (Figure 4-6).
Figure 4-6. ATR-FTIR spectra of collagen: POC: elastin crosslinked fibers.
Table 4-1. Average fiber diameter of collagen: POC: elastin fibers with different cross-linking times.
Time of crosslinking Average fiber diameter (µm)
Uncrosslinked 0.816 2h 1.044 4h 1.436 6h 1.900 8h 2.414
Higher magnification SEM images of the nanofibrous layer revealed a uniform
bead-less nanofibrous network with microporous structures. Further SEM
characterization of the scaffolds showed that with higher crosslinking times the
openness and apparent porosity between fibers decreased. The average fiber diameter
of the scaffolds also increased with higher crosslinking time as shown in Table 4-1
above. The average fiber diameters ranged from 0.816 µm-2.414 µm. Previous work by
Rüder et al. 2013 demonstrated significant smooth muscle cell adhesion, spreading,
and proliferation on scaffolds with submicron sized fibers 190. As discussed in previous
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chapters, porosity and openness are important characteristics for adequate cell growth.
Therefore, to balance the openness of structure as well as to ensure acceptable
mechanical compatibility, a 6h crosslinking time was chosen for scaffolds prepared for
the rest of this study.
Another important physical characteristic of biomaterial scaffolds is their
hydrophilic nature. Many studies in the past have shown that a hydrophilic surface is
preferred for biomedical applications. The wettability of the surface plays an important
role in cell adhesion and spreading, thereby influencing the biocompatibility of the
biomaterial191–193. To assess the hydrophilicity of the surfaces used in this study, the
contact angles of the water – biomaterial interface on POC and collagen: POC: elastin
films were tested. The experiment showed that contact angle at the sample water
interface on POC and collagen: POC: elastin films, were 60.78 ± 5.94° and 55.62 ±
8.12° respectively. The image of the water droplet and the contact angle measure are
shown below in Figure 4-7 and Table 4-2 below. The fact that both contact angles were
below 90° implied that both layers of the biphasic scaffold exhibited hydrophilic nature
and therefore are suitable for cell adhesion.
Figure 4-7. Contact angle measurements. A representative digital image of a water droplet on (A) collagen: POC: elastin film and (B) POC film.
Table 4-2. Contact angles at the water – biomaterial interface
Material Composition Contact Angle
Collagen: POC: Elastin 55.62˚ ± 8.12˚
POC 60.78˚ ± 5.94˚
A B
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The medial layer of native arteries are enriched with collagen and have been
known to carry the majority of the mechanical load in an artery90,177,194. It was essential
that the collagen: POC: elastin layer was able to provide the required mechanical
support to the construct. The preliminary mechanical properties (as shown in table 4-3
below) of the electrospun scaffolds were measured by quantifying its uniaxial tensile
properties. The hydrated tensile testing of the scaffolds showed a young’s modulus
value of 1.89 +/- 0.5 MPa and UTS value of 0.76 +/- 0.04 MPa. These results align with
values reported for medial layer of human carotid artery in literature195–198. Further the
percentage elongation of the scaffold was measured to be 60.9 +/- 16.34 %. Although
this value was not comparable to arterial elongation in vivo, the addition of an
elastomeric POC layer is expected to improve the elasticity of the scaffold construct.
Table 4-3. Uniaxial tensile test. Assessment of tensile properties of the glutaraldehyde vapor crosslinked electrospun mat
Young’s modulus (MPa)
Ultimate Tensile Strength (MPa) % Elasticity
1.89 +/- 0.5 0.76 +/- 0.04 60.9 +/- 16.34
An essential requirement of a cardiovascular biomaterial is biocompatibility and
capability to house appropriate vascular cells and to promote their growth. It has been
shown the incorporation of natural ECM components like collagen and elastin in tissue
scaffolds provide a biomimetic substrate for scaffolds and promote cell functions
relevant for tissue regeneration78,81,89,184. Preliminary biocompatibility of the scaffold
construct was assessed by evaluating cell adhesion and proliferation on both layers of
the bilayered scaffold. Additionally, the cytotoxic response invoked by the scaffold
layers was also quantified using an LDH assay as described in the methods section.
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Previous reports have raised concerns about potential cytotoxic risk posed by
residual aldehydes from the glutaraldehyde crosslinking of proteins. Therefore, in this
study, the electrospun and glutaraldehyde vapor crosslinked scaffolds were leached by
soaking in DMEM for at least a week before use in biocompatibility experiments.
HUVECs and hASMCs were used to perform biocompatibility assessments because a
monolayer of endothelial cells makes up a significant portion of the intimal layer in
native vasculature and smooth muscle cells reside in the nanofibrous medial layer.
Corroborating the hypothesis from the contact angle measurements and previous work
that has reported adhesion of endothelial cells on POC films 178,179,199,200.
Figure 4-8. SEM images of cell attachment. hASMCs seeded onto collagen: POC: elastin fibers.
SEM imaging (Figure 4-8) revealed that the hASMCs adhered to the collagen:
POC: elastin and show preliminary evidence of migration under the electrospun fibers
into the bulk of the scaffold. But due to the presence of proteins in the matrix, fixation of
cells with glutaraldehyde introduced some artifacts in the SEM image as shown above.
Therefore, the morphology of the fibrous scaffold was not clearly visible. Current efforts
to image the cell seeded scaffolds using alternate techniques such as environmental
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SEM or confocal microscopy is underway. Nevertheless, the images indicated
preliminary evidence of cell attachment and spreading indicating healthy cell adhesion
on scaffold layers. These results were in accordance to enhanced smooth muscle cell
adhesion on collagen, elastin, and POC individually and in combination with other
materials as shown in literature.90,178,200,201
Figure 4-9. Relative cytotoxicity measurements of collagen: POC: elastin with hASMCs and POC lumen with HUVECs. Data are reported as mean ± standard deviation (n=3). N.S. indicates no statistically significant difference (p > 0.05) and * indicates a statistically significant increase (p < 0.05).
Colorimetric detection of the intracellular enzyme LDH was used to assess
cytotoxicity of the scaffold layers on the HUVECs and hASMCs. Relative cytotoxicity
measurements revealed a 5.83 ± 3.68 % cytotoxicity for the HUVECs on the POC
lumen and 5.49 ± 0.91 % on the tissue culture plastic. For the hASMCs the cytotoxicity
of the hASMCs on the collagen: POC: elastin electrospun layer was 14.24 ± 7.96 %
compared to 21.18 ± 11.12 % on tissue culture plastic (Figure 4-9). Compared to the
tissue culture plastic control there was no statistical significance (p > 0.05) in the
cytotoxicity of the two layers of the bilayered scaffold.
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After confirming the preliminary biocompatibility of the scaffold layers, the
proliferation of cells on the electrospun medial layer was assessed via metabolic activity
which was correlated to cell number. At time points of 1, 3, 5, and 7 days MTT reagent
was added and then the purple formazan crystal produced were solubilized. The
absorbance of the media was read for the hASMCs on the collagen: POC: elastin
(Figure 4-10).
Figure 4-10. Proliferation kinetics. Kinetic curve representing the proliferation of hASMCs on the collagen: POC: elastin. Data are reported as mean ± standard deviation (n=3). N.S. indicates no statistically significant difference (p > 0.05) and * indicates a statistically significant increase (p<0.05).
After the initial seeding at t = 1 day the cell number was calculated as 7133.33 ±
321.46 cells. At t = 3 days and t = 5 days the cell number increased to 7566.67 ± 57.73
and 8300 ± 264.58 cells respectively. From t =3 days to t = 5 days there was a
statistically significant increase in cell number (p < 0.05). At t = 7 days, the extrapolated
cell number was 9933.33 ± 513.16 and between t = 5 days and t = 7 days there was a
statistically significant increase (p < 0.05).
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The biocompatibility results indicate the scaffold is able to support relevant cell types
growth and proliferation and preliminarily promote relevant cell functions to support
tissue regeneration. To fully assess the regenerative capacity of the scaffold further
studies would need to be done with longer culture times and in a dynamic environment
or in vivo.
Figure 4-11. SEM characterization of POC; POC-PAA bilayered construct. (A) the biphasic scaffold that contains solid lumen and nanofibrous medial layer, 500X (B) interface between solid lumen and electrospun fibers, 2000X, and (C) the fibrous electrospun mat, 2000X.
Encouraged by these preliminary results, the next step was the optimization of
the bilayered construct fabrication. The methodology of this process was optimized
using a solid lumen layer consisting of POC surrounded by an electrospun POC: PAA
layer. The POC layer was partially thermally crosslinked for 2 days at 60 ̊C and 80 ̊C
respectively. A layer of POC PAA was then electrospun onto the POC layer using
predetermined electrospinning parameters. The scaffold construct was further
crosslinked thermally as described. The fabricated scaffold was then visually inspected
using SEM imaging (Figure 4-11) which showed complete integration between the two
layers and the clear distinction between the porous morphology of the medial
electrospun layer versus the solid non-porous POC luminal layer.
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The fabrication technique optimized above was applied to construct the bilayered
scaffold incorporating a medial layer of electrospun collagen: POC: Elastin. The
fabricated construct was also crosslinked for 6h in gluataraldehyde vapor and
crossections of the scaffold was SEM imaged. Images (Figure 4-12) also revealed
integration between the electrospun and solid polymer layer of the biphasic scaffold.
Figure 4-12. SEM characterization of POC; collagen: POC: Elastin bilayered construct. Representative SEM image of (a) the biphasic scaffold that contains solid lumen and nanofibrous medial layer, 500X (b) interface between solid lumen and electrospun fibers, 2000X, and (c) the fibrous electrospun mat, 2000X.
The mismatch in compliance between a vascular graft and its host artery is one
of the major reasons for failure after implantation. Compliance mismatch has been
known to cause (1) anastomotic rupture, (2) flow induced shear stress and, (3) reduced
distal perfusion. All of these, factors eventually lead to build up of uneven stresses and
affect long term patency of the grafts202–204. Therefore a compliant scaffold that closely
matches the mechanical properties of native vasculature is imperative to successfully
design a tissue engineered vascular construct 8,9,87. The assessment of the
circumferential mechanical properties of the bilayered construct revealed highly
promising results. Salacinski et al., showed based on literature values there is a linear
relationship between the patency of the scaffolds and how closely the compliance
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matches native arteries. In this study, the saphenous vein, an autologous replacement
with the most similar compliance(4.4 ± 0.8 %), compared to native arteries(5.9 ± 0.5%),
showed the greatest patency rate (75%).However, ePTFE, a synthetic graft material
with a significantly higher compliance, (1.6 ± 0.2 %) exhibited the lowest patency rate,
40% 205.
Table 4-4. Circumferential mechanical testing of bilayered scaffolds compared to native tissue and ePTFE. 202,203,206,207.
Bilayered Scaffold Native Human Artery
Human Saphenous Vein ePTFE
Compliance (%) 6.9 ± 0.14 5.9 ± 0.5 4.4 ± 0.8 1.6 ± 0.2
Burst Pressure (mmHg) 401.68 ± 125.31 3,196 ± 126 1,599 ± 877 -------
Young’s Modulus (MPa) 0.330 ± 0.08 9-12 2.25-4.2 2.2
Circumferential Tensile Strength (MPa)
0.201 ± 0.06 ------ 3.7 ± 2.0 -------
The mechanical compliance (as shown in Table 4-4) of the scaffold was 6.94 ±
0.14% which is very close to reported values for native vasculature and therefore has
the potential to maintain a good patency rate similar to that of the saphenous vein,
probably even better than the current clinical synthetic standard, ePTFE202,204,205.
Further, the Young’s modulus of the bilayered scaffold was calculated to be
0.330 ± 0.08 MPa, and the burst pressure extrapolated using the Laplace relationship
was 401.68 ± 125.31 mmHg. The theoretical burst pressure was lower than native
arteries, (3,196 ± 126 mmHg), and saphenous veins (1,599 ± 877 mmHg)207. Although,
the bursting pressure was more than physiological levels of pressure, it is hypothesized
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that as the cells infiltrate and deposition ECM, they actively begin remodeling the
scaffold underneath. This may potentially lead to improved mechanical properties such
as Young’s modulus and burst pressure that align more with that of native tissue.
Niklason et al. showed from 5 to 7 weeks PGA scaffolds cultured dynamically with
vascular cells showed a remarkable increase in burst pressure from approximately 600
mmHg to 2200 mmHg117,206.
The main goal of designing tissue engineered vascular grafts is to find a
functional replacement for native vessels. These vessels consist of a dynamic,
hierarchical and complex structure and composition. In order to create a successful
replacement for native vessels, the design strategy to try to mimic not only the
composition but also the microarchitecture of the native vessel. The idea behind this
strategy is to best imitate the microarchitectural, chemical as well mechanical cues
provided by the ECM to a native blood vessel. With this rationale in mind, this study
involved design, fabrication and characterization of a novel bilayered biomimetic
construct consisting of a solid lumen and a fibrous medial layer. The medial layer was
enriched with native ECM proteins to improve bioactivity of the scaffold system. Overall
the results of this study indicate show that the bilayered scaffold was preserved the
architectural characteristics as shown by SEM imaging. It was also biocompatible and
was able to promote adhesion and proliferation of relevant vascular cell types.
More importantly the mechanical properties of the construct were compatible to
existing autologous replacements which are the golden standard in the field currently.
This scaffold is currently being developed to recruit circulating progenitor cells from the
blood and use them to endothelialize the graft surface. It is also believed that this
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strategy will promote integration of the scaffold into existing vasculature. Although many
previous reports have demonstrated the use of electrospun protein-polymer blends for
vascular grafts, mechanical compliance comparable to native tissue as well improved
biocompatibility has not been reported so far. This novel scaffold system is a promising
strategy for developing readily available vascular graft implants.
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CHAPTER 5 MESOSCALE STRUCTURAL CUES- PART 2- HEMOCOMPATIBILITY EVALUATION
OF A NOVEL MG-CA-SR ALLOY FOR VASCULAR STENT APPLICATIONS
Introduction
Another important class of cardiovascular biomaterials include metals which are
majorly used in designing heart valves, endovascular stents, and stent-graft
combinations. Metallic biomaterials are largely used in biomedical applications due to
their inertness and high mechanical strength67. However, some of the commonly used
bio-metals possess limited biofunctionalities, such as blood compatibility, bone
conductivity and bioactivity. To be used in cardiovascular applications, bloody
compatibility as well as compatibility with surrounding cellular environment (based on
the application), is a must. Failure of which leads to complications such as thrombosis
and inflammation which not only prevents incorporation of the implant into the existing
vasculature but also causes significant damage to it208,209. Therefore, the goal of
designing implants for vascular engineering is to design implants with cues that enable
repair of existing vasculature and prevent any susceptibility to damage due to the
implant placement.
As mentioned in the earlier chapters, the interaction of a biomaterial with its
surrounding environment is determined by its composition as well as structural features.
These factors govern the immediate as well as long-term responses after implantation.
In case of bio-inert metallic implants, a very thin layer of metal oxide formed, provides
the ultimate interface between stent and host tissue. The type of metal oxide formed
itself varies based on the composition of the material and the processing conditions
used. It has also been known that the thrombogenic response as well as tissue
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incorporation is largely determined by physical characteristics of the surface, such as
roughness, electrical change, free surface energy, and wettability101,103,107.
The initial interaction of metallic implants with the surroundings and dynamic
series of events that immediately follow triggers responses that facilitate the growth of
endothelial cells over the surface of the prosthesis107. Similarly, biomechanical behavior
also seems to affect the host’s response as well. Therefore, it is important to study and
understand the physical and chemical characteristics of metallic biomaterials before
they can be used for blood-contacting applications. This will be essential to design
structural, mechanical and chemical cues presented by the implant to the surrounding
system to elicit a favorable response.
Atherosclerosis is a life threatening cardiovascular disease that causes the
formation of plaques in arteries and is characterized by inflammatory infiltrates, lipid
accumulation, cell death and fibrosis91,210,211. To treat this condition various procedures
have been employed to mechanically improve perfusion in the artery without resorting to
surgery98,209,212. The most common procedure is percutaneous transluminal coronary
angioplasty, by implantation of an intracoronary stent to expand the artery. Despite
major breakthroughs and advancements in the research on permanent vascular stents
for treatment of coronary arterial diseases, major challenges such as stent thrombosis,
size mismatch, mechanical blockages and late development of restenosis and
malposition still exist213. These challenges have motivated and steered the research in
this field towards biodegradable and bioabsorbable stent materials. Biodegradable
vascular stents are currently the focus of research and clinical advancement, due to 1)
their ability to support the growth and remodeling of native vascular tissue; 2) the
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possibility to deliver therapeutic agents through degradation; and 3) their temporary
nature which eliminates the risk of late stent thrombosis213–216. Historically many
polymeric biomaterials such as, Poly L-lactic acid (PLLA), polyglycolic acid (PGA),
polycaprolactone (PCL) and poly (D, L-lactide/glycolide) copolymer (PDLA) have been
investigated as candidate materials for biodegradable stent applications 70,210,214,215.
While polymeric stents are advantageous because of their increased ability for targeted
drug delivery, higher drug loading capacity, and potential ability to allow percutaneous
revascularization and surgical intervention when necessary, there are some significant
disadvantages to the use of polymeric stents. The major disadvantage being their low
mechanical strength which may lead to stent fracture and migration over time77,216,217.
Moreover, polymeric stents have been known to cause a significant degree of local
inflammatory response and the long term biocompatibility of the degradation products of
polymeric materials remain unclear70,108.
These limitations have caused increased interest in biodegradable metallic stents
in the recent years, including magnesium (Mg)-based materials210,218. Magnesium-
based materials have been investigated for a variety of biomedical applications,
including but not limited to orthopedic and cardiovascular implants111,219–223. This
approach is supported by the fact that Mg is a naturally occurring element in the human
body, which is the second most abundant cation after sodium and has been reported to
exhibit very low to no toxicity219. When considering the use of Mg in vascular stent
applications, early studies have shown the beneficial effect of Mg ions to treat coronary
artery disease based upon their inherent anti-thrombogenic properties114. Mg was
shown to inhibit platelet activation through modulating the release of platelet activating
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factors, tromboxane A2 and prostaglandin224–228. The challenge in using pure Mg in
vascular stent applications is its very high degradation rate which results in a significant
loss of mechanical integrity, which is of critical importance for vascular stent
applications. This challenge can be overcome by alloying Mg with elements such as
Aluminum, Zinc, Yttrium and Cerium to decrease the degradation rate. While successful
in decreasing the rate of degradation, these alloys are inappropriate for biomedical
applications due to increased toxicity219. Herein, an alloy composed of Mg, Calcium and
Strontium (Mg-Ca-Sr) as a candidate material for vascular stent applications was
investigated. The high strength and low degradation rate of this alloy composition was
reported by Berglund et al. 2012222. These studies also characterized and reported the
low cytotoxicity and bioactivity of the degradation products from this alloy. Although
good biocompatibility and low degradation rate are primary requirements for materials
for stent applications, the compatibility of the material with blood and blood components
is of utmost importance213. Hemocompatibility of a vascular stent material is essential to
prevent fatal complications such as thrombosis, restenosis and inflammation post
implantation. Therefore, this study focuses on evaluating the hemocompatibility of the
Mg-Ca-Sr alloy with whole porcine blood and blood components as well as human
vascular and blood cells. The interaction of the alloy with porcine blood was evaluated
via assessment of hemolysis, whole blood clotting, platelet adhesion and activation.
Additionally, the inflammatory response of the alloy was assessed by quantifying the
release of pro-inflammatory cytokines/chemokines such as Interleukin 1 beta (IL-1β),
Interleukin 6 (IL-6), Interleukin 8 ( IL-8) & Tumor Necrosis factor alpha (TNFα) by a
human monocytic cell line (THP-1 cells) when in contact with the alloy. Lastly, the
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cytotoxicity of the alloy degradation products on vascular wall specific endothelial cells
was evaluated. This work represents the first report providing a comprehensive
evaluation of the hemocompatibility of this Mg-Ca-Sr alloy, and is the foundation for
future work on the use of this alloy material for vascular stent applications.
Materials and Methods
Alloy Synthesis and Sample Preparation
Pure elements (98.5 wt. % Mg, 1.0 wt.% Ca, and 0.5 wt.% Sr) were melted in an
induction furnace and direct chill cast to cylindrical billets 60 mm in diameter and 120
mm in length. The billet was homogenized at 450°C for 24 h under inert atmosphere. It
was then machined to 50 mm diameter and extruded in a 1.5 MN press at 0.5 mm/s to a
diameter of 10 mm for an extrusion ratio of 1:25. Temperature of 275ºC was maintained
by induction heating. Samples for cytotoxicity and hemocompatibility studies were
fabricated by cutting discs ~1mm thick on a low speed cutting saw, then grinding at 320
then 1200 grit SiC paper on the faces and sides. Controls of medical grade 316
stainless steel were fabricated in a similar fashion, then passivated per ASTM
A967/A967M-13 in 30% nitric acid for 1 the at room temperature. Poly(L-lactide) (PLLA)
discs were also included as a control for degradable polymeric material. The discs were
made by solution casting 5 % PLLA solution in chloroform. The cast film was washed 5
times with ethanol and De-ionised water to remove any residual chloroform.
Cytotoxicity Evaluation
Human umbilical vein endothelial cells (HUVECs) (Lonza, Walkersville, MD) were
used as a positive control for mature endothelial cell functionality. Endothelial cell basal
media-2 (EBM-2) (Lonza) supplemented with 2% fetal bovine serum, human fibroblast
growth factor (hFGF), vascular endothelial growth factor (VEGF), human epidermal
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growth factor (hEGF), human recombinant insulin-like growth factor (R3-IGF-1),
ascorbic acid, gentamicin, amphotericin, and heparin is referred to as endothelial cell
growth media-2 (EGM-2). HUVECs were cultured in EGM-2 culture medium and utilized
within passages 3-5. To assess the cytotoxic risk posed by the alloy degradation
products to the surrounding cellular environment, sample extract medium was prepared
according to ISO 10993-5 where 1 mL of extraction medium was added per 3 cm2 of
sample surface area. Briefly, 9.5mm diameter discs of Mg-Ca-Sr, S.S, and PLLA were
sterilized with 70% ethanol followed by UV radiation for 1 hr. The sterilized samples
were then placed in phosphate buffered saline (PBS) for 24 h, followed by another 24 h
incubation at 37̊C and 5% CO2 in complete EGM-2 media. After which, the media
extract containing degradation products was collected. Following collection, the cell
culture supernatant was serially diluted to 10, 30 and 50% in complete medium and
assessed for LDH release according to a modified protocol reported by Berglund et
al.222 . The cytotoxic response of HUVEC to alloy degradation products was evaluated
by quantifying release of lactate dehydrogenase (LDH) in the culture medium. LDH
released into the media from damaged cells is a biomarker for cellular cytotoxicity and
cytolysis. Briefly, HUVECs were seeded at a density of 5000 cells/well of a 96 well plate
and allowed to attach for 18 h. The culture medium in the wells were then replaced with
alloy extract and incubated for 24 h in a cell culture incubator. Following manufacturer’s
protocol, a colorimetric substrate for LDH (Pierce Biotechnology, Rockford, IL) was
incubated with the extracts for 30 min at 37 ̊C. The reaction was stopped with the
addition of 1N hydrochloric acid. The optical density was measured at 490 nm with a
reference wavelength of 650 nm.
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Hemolysis
The percentage of red blood cell lysis (RBC) or hemolysis was quantified using a
modified protocol reported by Motlagh et al. 2006. Circulating peripheral blood was
collected from the exterior jugular vein of domestic pigs (Department of Animal
Sciences, University of Florida) in the presence of acid citrate dextrose anticoagulant.
All blood collection procedures were performed in accordance with the rules and
regulations of the Institutional Animal Care and Use Committee (IACUC) at the
University of Florida. To assess hemolysis, 9.5 mm diameter discs of the Mg-Ca-Sr
alloy, S.S, and PLLA were immersed in PBS overnight, then placed in 24 well plates.
The anticoagulated blood was then diluted 1:4 in 0.9% saline, then 1mL of diluted blood
was added to the wells containing test samples and incubated under gentle agitation on
a shaker for 2 h at 37 C̊. As a positive control for complete RBC lysis, whole blood was
diluted in distilled water an incubated in the absence of any of the test materials. Saline
diluted blood also incubated in the absence of the test materials serves as a negative
control for no lysis. The diluted blood from each well was centrifuged at 1000 X g for 10
min to pellet the intact RBCs. The supernatant, containing the released hemoglobin was
transferred to a 96-well plate in triplicates then the absorbance was measured at 545
nm. The percentage hemolysis was then calculated using the following formula shown
in below in Equation 4-1 (abs denotes absorbance).
%𝐻𝑒𝑚𝑜𝑙𝑦𝑠𝑖𝑠 =(𝐴𝑏𝑠 𝑜𝑓 𝑡𝑒𝑠𝑡 𝑚𝑎𝑡𝑒𝑟𝑖𝑎𝑙)−(𝐴𝑏𝑠 𝑜𝑓 𝑛𝑒𝑔𝑎𝑡𝑖𝑣𝑒 𝑐𝑜𝑛𝑡𝑟𝑜𝑙)
(𝐴𝑏𝑠 𝑜𝑓 𝑝𝑜𝑠𝑖𝑡𝑖𝑣𝑒 𝑐𝑜𝑛𝑡𝑟𝑜𝑙)−(𝐴𝑏𝑠 𝑜𝑓 𝑛𝑒𝑔𝑎𝑡𝑖𝑣𝑒 𝑐𝑜𝑛𝑡𝑟𝑜𝑙)𝑋100 (4-1)
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Evaluation of Platelet Adhesion and Activation
Platelets isolated from porcine whole blood was prepared using a modified protocol
as reported by Motlagh et al. (2006)229 to assess and quantify platelet adhesion and
activation in response to the Mg-Ca-Sr alloy in comparison with control materials, S.S,
and PLLA. Collagen coated plastic was included as a positive control for platelet
adhesion and activation 84. To assess platelet adhesion, platelet rich plasma (PRP) and
quantified platelets by measuring LDH released by platelets upon lysis was obtained.
Briefly, whole blood was centrifuged at 250 X g for 15 min and the platelet-rich
supernatant was removed and the number of platelets were counted using a
hemocytometer. Samples of the alloy, S.S and PLLA (9.5mm diameter) immersed in
PBS overnight, were incubated with 100 µl of PRP containing 1 X 106 platelets per
sample for 1h at 37 ̊C. The platelets adhered on the samples were first lysed by
incubation with 2% Triton-PBS buffer for 45 min at 37 ̊C. The number of adherent
platelets was determined by detecting the amount of LDH present after cell lysis
following manufacturer’s protocol. A calibration curve was generated from a series of
serial dilutions of known platelet concentrations and was used to determine the number
of adhered platelets.
The activation of platelets on exposure to alloy samples and control materials was
assessed by quantifying the release of platelet factor 4 (PF4) by activated platelets.
Briefly, Mg-Ca-Sr alloy, S.S and PLLA samples were incubated with PBS overnight. The
samples and controls were then incubated with 100 µl of PRP containing ~6 X 106
platelets, for 2h at 37̊C. The PRP was then collected and assayed to quantify the
amount of PF4 released via an ELISA (ABclonal Science, Inc., Woburn, MA) per
manufacturer protocol.
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Whole Blood Clotting Assessment
The clotting behavior of recalcified whole blood in the presence of the Mg-Ca-Sr
alloy, S.S and PLLA was assessed using a modified protocol previously reported
elsewhere. 84,229–231 Briefly, whole blood collected in the presence of anticoagulant was
added onto 9.5 mm sample discs in a multiwell plate (20 µl per sample/control).
Following addition of blood, 10 µl of 0.2 M CaCl2 was added to recalcify the blood to
activate the clotting cascade. The addition of CaCl2 was omitted for a subset of wells as
a control for no clotting. At 1, 30 and 60 min after the addition of CaCl2, 1 mL of distilled
water was gently added to each well and mixed to lyse free RBCs that were not part of
the clot formed. The liquid was then removed from each well and the absorbance of the
supernatant was measured. Higher absorbance indicated higher thromboresistance of
the material. The absorbance at 30 and 60 min from each sample/control was recorded
and was normalized to initial absorbance at 2 min and the percentage decrease in
absorbance was calculated and correlated with percentage increase in clotting. In this
study, glass served as a positive control representing the kinetics of uninhibited whole
blood clotting.
Release of Monocytic Inflammatory Cytokines
The release of inflammatory cytokines (IL-1β, IL-6, IL-8 & TNFα) in the presence of
the alloy was assessed using a human acute monocytic leukemia cell line (THP-1 cells)
(ATCC, Manasass,VT). The cells were cultured in suspension in complete RPMI-1640
media (FisherScientific, Waltham,MA) for several days at 37 ̊C, 5% CO2 in a cell culture
incubator. The cells were then centrifuged and re-suspended at a concentration of 1 X
106 cells/ mL. A 1-mL cell suspension was placed in each well of a multiwell plate
containing 9.5mm diameter discs of Mg-Ca-Sr, S.S, and PLLA. Control cell population
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grown in plain growth media and media including lipopolysaccharide (LPS, 10 ng/mL)
were used as positive and negative controls for inflammatory response, respectively.
THP-1 cells were incubated in the presence of samples at 37̊C for 24 h under static
conditions. The cells were centrifuged and pelleted and the media supernatant was
collected and assayed. The concentrations of IL-1β, IL-6, IL-8 & TNFα in the
supernatants were determined using a multi analyte ELISA kit according to the
manufacturer’s instructions (Qiagen, Valencia, CA).
Statistical Analysis
All data in this study are reported as mean +/- standard deviations. An ordinary one
way ANOVA (Analysis of variance) with Dunnett’s multiple comparison post hoc
analysis was used to analyze and identify significant differences between the control
column and other columns. P-values <0.05 are reported as significant.
Results and Discussion
In spite of great advancement in the field, the major challenge while using
biodegradable metallic intracoronary stents is the prevention of thrombosis formation
and restenosis development after implantation108,110,232. This is primarily governed by
the interaction at the interface of the implanted biomaterial and the surrounding
biological environment. Therefore, the compatibility of the implanted stent material with
blood and blood components as well as surrounding vascular cells is of utmost
importance. In this study, the hemocompatibility and biocompatibility of a novel Mg-Ca-
Sr for its potential application as a vascular stent was investigated. Microstructure,
strength, and in vitro degradation behavior of this material and other extrusion
temperatures are reported in depth elsewhere. The microstructure exhibited large (50
µm x 10 µm) elongated grains interspersed with dynamically recrystallized grains
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around 1 µm in diameter, with a typical extrusion texture. The eutectic micro constituent
consisting of a combination of Mg, Mg2Ca, and Mg17Sr2 was broken up into strings of
particles in the extrusion direction. Mechanical properties of the material have been
investigated in the previous study. The alloy exhibited substantial mechanical properties
with yield strengths of 300 MPa and 250 MPa in tension and compression, respectively.
These mechanical properties are higher than that reported for other magnesium alloys
(WE43) as well as 316L annealed stainless steel which have been used in clinical
studies for vascular applications107,233. Additionally, the material exhibited a low in vitro
degradation rate in Hank’s solution, with hydrogen evolution ~0.022 mL H2/cm2/day.
Figure 5-1. Evaluation of cytotoxicity. Toxicity on HUVECs, expressed as a percentage of relative cytotoxicity, for three concentrations of alloy extract medium (10 %, 30 % & 50 %) in complete media. Data expressed as mean +/- SD, N=3. % Relative cytotoxicity of 10 % extract: 1.86 +/- 0.34 %, 30 % extract: 3.13 +/- 1.56 %, 50 % extract: 5.02 +/- 2.18 %, spontaneous TCP: 0 +/- 3.33 %. Data statistics based upon ordinary one-way analysis of variance (ANOVA) coupled with Dunnet’s multiple comparison test, N.S. indicates no significant difference (P > 0.05) and * indicates a statistically significant increase (P<0.05).
As a first step towards assessing the biocompatibility of this alloy cytotoxicity
response of the alloy was studied. Cellular response is one of the central factors
involved in selection, design and application of biomaterials. One of the rudimentary and
standardized techniques for assessing cellular response is by measuring in vitro
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cytotoxicity of the biomaterial, with appropriate cell types based on its eventual
application234. In this study the biocompatibility of the alloy was evaluated by assessing
cytotoxicity of alloy degradation products with human vascular endothelial cells. In this
study, vascular endothelial cells were chosen to closely mimic and model the cellular
environment that will be in direct contact with a vascular implant material. The cytotoxic
response elicited by the cells to the alloy was assessed via LDH release. HUVECs
exposed to 10, 30 and 50% diluted alloy degradation extracts were assessed for LDH
released into the media over a period of 24 h. The results as presented in Figure 5-1
revealed that the alloy extracts showed very low relative percent cytotoxicity (1.86+/-
0.34%, 3.13+/-1.56%, & 5.02 +/- 2.18 %) with all three extract concentrations (10, 30
and 50%) respectively. These results are consistent with previous reports on the
evaluation of cytotoxicity of this alloy composition with a pre osteoblast cell line (MC 3T3
E1) 222. The data further confirms the low cytotoxic response induced by the alloy on
human primary vascular endothelial cells which is an essential requirement of a
vascular stent material. Further the hemocompatibility of the alloy was assessed via
hemolysis, characterization of whole blood clotting behavior, platelet adhesion and
activation using circulating blood derived from domestic pigs. Hemolysis evaluation is
one of the most commonly used methods to determine preliminary hemocompatibility
and biocompatibility of a biomaterial235. The interaction of the biomaterial implant with
circulating blood can result in the release of hemoglobin due to RBC lysis. This process
can be influenced by chemical interaction (release of ions) of RBC with the degradation
products. The intactness of RBCs is important for maintenance of microcirculation. The
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lysed RBCs also release factors that can have detrimental effects such as platelet
activation and subsequent vasoconstriction236.
Figure 5-2. Hemolysis assessment. Percentage hemolysis of RBCs following incubation of diluted whole blood with Mg-Ca-Sr: 0.70 +/- 0.17 %, S.S: 0.02 +/- 0.05 %, PLLA: 0.16 +/- 0.28 %. Blood samples exposed to complete hemolysis with water (Max lysis: 100.00 +/- 9.23 %) and Blood samples with nothing added (No lysis: 0 +/- 0.11 %) were used to normalize the data (not shown in figure). Data expressed as mean +/- SD, N=9. Data statistics based upon ordinary one-way analysis of variance (ANOVA) coupled with Dunnet’s multiple comparison test, N.S. indicates no significant difference (P > 0.05) and * indicates a statistically significant increase (P<0.05). (C) Representative digital images of hemolysis by sample and controls. Hemolysis percentage correlated with absorbance of the supernatant.
Therefore, it is essential to determine the hemolytic percentage of the alloy for
vascular applications where the material is in direct contact with circulating blood. The
percentage hemolysis of the alloy as well S.S and PLLA controls as shown in Figure 5-2
indicate that the alloy caused very low and negligible (<1 %) hemolysis when incubated
with diluted blood for 2h. ISO 10993-4:2002 requires that the hemolysis rate of
hemocompatible materials be below 5 % 237. Previous reports assessing the
hemocompatibility of pure Mg, as well as, Mg-based alloys indicated that pure Mg as
well as most Mg based alloys were severely hemolytic. There are very few alloys such
A B
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as Mg–6Zn, Mg–1Si and WE43 that have been reported to show hemolytic rates less
than 5 % 219,233,238. High degradation rate of alloys have been known to result in drastic
pH variations which causes high hemolytic rates, therefore the sustained degradation
characteristics as exhibited by the alloy is beneficial to prevent excessive hemolysis233.
The results show that the Mg-Ca-Sr alloy causes very low hemolysis in comparison with
existing alloy systems thereby confirming the applicability of this alloy for vascular
applications.
Other than preliminary biocompatibility, the behavior of a potential stent material
with blood in terms of effect on clotting cascade is also an important criterion for
vascular application. Stent thrombosis caused by clot formation post implantation, leads
to catastrophic consequences that can result in death. Despite various anticoagulation
therapies after implantation, late stent thrombosis still remains a major challenge.108,208
Hence, one of the essential design considerations while choosing biomaterials for blood
contacting applications is the ability to resist thrombus formation. the thromboresistance
of this alloy along with other reference materials when in contact with activated porcine
whole blood was evaluated 229–231. The data (Figure 5-3) indicated that at 30 min, Mg-
Ca-Sr showed the lowest percentage clotting (P<0.05) when compared to S.S., PLLA
and Glass. At 60 min, even though there was some clot formation on the Mg-Ca-Sr
samples, the percentage clotting was still significantly lower than that initiated by S.S as
well as glass (P<0.05). The percentage clotting on PLLA samples however was
statistically similar to the alloy (P>0.05) at 60 min. Collectively, the data from the whole
blood clotting studies demonstrates that the Mg-Ca-Sr alloy exhibits significant
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thromboresistance which is essential for optimal performance of implants used for
vascular applications 71,109.
Figure 5-3. Whole Blood Clotting assessment. Percentage clotting of whole blood expressed as percentage decrease in absorbance calculated with respect to initial absorbance (A) at 30 min (Mg-Ca-Sr: 4.12 +/- 4.39 %, S.S: 34.08 +/- 5.31 %, PLLA : 25.22 +/- 10.78 % & Glass: 69.25 +/- 3.87 % ) (B) at 60 min (Mg-Ca-Sr: 43.46 +/- 11.79 %, S.S: 83.82 +/- 1.39 %,PLLA: 48.87 +/- 3.61 % & Glass: 84.82 +/- 1.84 % ). Data statistics based upon ordinary one-way analysis of variance (ANOVA) coupled with Dunnet’s multiple comparison test, N.S. indicates no significant difference (P > 0.05) and * indicates a statistically significant increase (P<0.05). (C) Representative digital images of clot formation on sample and control surfaces at 30 min and 60 min.
The ability to resist/delay clot formation is important to prevent thrombosis which
leads to fatal complications and subsequent failure of implant 84,109,216,229. It is important
A B
C
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to note that both Ca2+ ions and Sr2+ ions have been known to cause surface clot
formation, platelet adsorption and activation when exposed to blood plasma 225,239,240.
However, the Mg-Ca-Sr alloy shows desirable resistance to thrombus formation. This
effect is possibly due to the well-known anti-thrombotic effect of Mg ions 114,225,227,241.
Previous studies have shown that administration of Mg produces an inhibition of acute
stent thrombosis under high-shear flow conditions without any hemostatic or significant
hemodynamic complications113. Interestingly, many studies have also shown the ability
of Mg ions to act as natural Ca blockers and this may explain the improved blood
compatibility of the alloy despite the presence of clotting activating ions such as Ca 2+
and Sr 2+ ions 224,242. Further investigation is required to clearly understand the
mechanism of thromboresistance and synergistic interaction between the Mg, Ca and
Sr ions released by the alloy.
After confirming the thromboresistive behavior of the alloy, to further characterize
the interaction of this alloy with blood components, adhesion and activation platelets on
sample surface was quantified. Both platelet aggregation and activation on biomaterial
surfaces are major biomarkers of thrombogenicity of a biomaterial 75,243,244. Additionally,
platelets also play an important role in inflammation and immune responses. Platelets
release many factors that regulate inflammation which do not have any role in
hemostasis 245. Therefore, assessing the interaction of platelets on biomaterial surfaces
is vital for assessing its hemocompatibility as well as immune response 246. The
adhesion of platelets on the surface of the Mg-Ca-Sr alloy was quantified via release of
LDH by lysed platelets. The results as shown in Figure 5-4 A indicate that the number of
platelets adhered on the alloy surface was statistically similar to PLLA and S.S (P>
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0.05). However, the number of platelets adhered to the alloy was significantly lower than
collagen-coated plastic which was included as a control for positive thrombogenic
surface (P< 0.05). Although some surfaces may inhibit platelet adhesion, they can still
cause platelet activation 75
Figure 5-4. Platelet adhesion and activation. Data expressed as mean +/- SD, N=3. (A) Number of platelets adhered on Mg-Ca-Sr: 1.88E+06 +/- 4.93E+05, S.S: 2.30E+06 +/- 2.46E+05, PLLA: 9.69E+05 +/- 3.32E+05, Glass: 3.54E+06 +/- 4.46E+05 & Collagen coated plastic: 2.99E+06 +/- 8.34E+04 (B) Platelet activation by PF4 release in ng/ml by Mg-Ca-Sr: 5.33 +/- 1.02, S.S: 6.21 +/- 1.93, PLLA: 6.82 +/- 2.14, Glass: 5.88 +/- 1.45 & Collagen coated plastic: 25.66 +/- 2.40). Data statistics based upon ordinary one-way analysis of variance (ANOVA) coupled with Dunnet’s multiple comparison test, N.S. indicates no significant difference (P > 0.05) and * indicates a statistically significant increase (P<0.05).
When platelet activation occurs the platelets change shape and morphology
causing release of many intracellular granular contents245. Platelet factor 4 (PF4) is a
well characterized platelet-specific marker released from platelet alpha granules upon
activation247,248. Therefore, release of PF4 to measure the activation of platelets
contained within PRP when in contact with the samples in comparison with S.S and
PLLA controls was quantified as presented in Figure 5-4 B. The results indicate that the
amount of PF4 released by platelets in contact with the alloy was not statistically
A B
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different from PLLA or S.S (P> 0.05). However, the PF4 release caused by collagen-
coated plastic which was used a positive control for platelet activation in this study was
5-fold higher than that caused by the alloy (P<0.05). Some studies have reported the
increased platelet activation caused by metallic stents due to release of metal ions in
the blood plasma 104,249–251. Therefore, it was important to determine the effect of
platelet interaction with the alloy. The quantification of platelet adhesion and activation
indicated that the Mg-Ca-Sr alloy does not promote significant platelet adhesion or
activation, which are markers for both thrombus formation as well as inflammatory
response. This effect could potentially explained by the ability of Mg to reduce platelet
activation by inhibiting production of factors such as thromboxane A2252. Although
platelet aggregation and activation leads to inflammation, it is not the only mechanism of
inflammatory response. Therefore an assessment of inflammatory response was made
by assessing some markers of inflammation was made,since eliciting minimal
inflammatory response is also an important characteristic of a biomaterial 16,75. This
inflammatory response is mainly governed by interaction with monocytes and
macrophages 245,253. Monocytes have been known to mediate long-term chronic
inflammatory response. Peripheral blood monocytes when activated release many
cytokines and chemokines that play an important role in modulating inflammation
caused by the biomaterials254–256. Hence, to fully characterize the response to the Mg-
Ca-Sr alloy to evaluate its suitability as a vascular stent material, the release of pro-
inflammatory cytokines (IL-1β, IL-6 & TNF-α) and chemokines (IL-8) by THP-1 cells
which are widely studied to mimic monocytic response in cell culture models was
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quantified.253 The release of these factors was quantified via a custom designed ELISA
kit as described above (Figure 5-5).
Figure 5-5. Release of inflammatory cytokines/chemokines by THP-1 cells reported as absorbance at 450 nm. Data expressed as mean +/- SD, N=3. Blank absorbance value was not subtracted from the data sets for clarity. (A) Release of IL-1β, Mg-Ca-Sr: 0.03 +/- 0.01, S.S: 0.03 +/- 0.01, PLLA: 0.04 +/- 0.01, Untreated cells: 0.02 +/- 0.00 & LPS induced positive: 1.75+/- 0.16. (B) Release of IL-6, Mg-Ca-Sr: 0.02 +/- 0.01, S.S: 0.02 +/- 0.01, PLLA: 0.02 +/- 0.00, Untreated cells: 0.02 +/- 0.01 & LPS induced positive: 0.07+/- 0.00. (C) Release of IL-8, Mg-Ca-Sr: 0.47 +/- 0.06, S.S: 2.39 +/- 0.37, PLLA: 2.28 +/- 0.50, Untreated cells: 1.29 +/- 0.09 & LPS induced positive: 3.88+/- 0.06. (D) Release of TNF-alpha, Mg-Ca-Sr: 0.07 +/- 0.01, S.S: 0.13 +/- 0.01, PLLA: 0.13 +/- 0.01, Untreated cells: 0.11 +/- 0.00 & LPS induced positive: 0.51+/- 0.10. Data statistics based upon ordinary one-way analysis of variance. (ANOVA) coupled with Dunnet’s multiple comparison test, N.S. indicates no significant difference (P > 0.05) and * indicates a statistically significant increase (P<0.05).
THP-1 cells incubated with 10 ng/mL of Lipopolysaccharide (LPS) derived from
E. coli was included in these studies to induce an increased inflammatory response
while cells grown in culture media with no sample added was considered a negative
A B
C D
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control for inflammation 75,257. The data revealed that the alloy caused significantly lower
induction of inflammatory factors when compared to LPS induced controls.
The results also showed that the release of IL-1β and IL-6 (Figure 5-5A & B) caused by
the alloy was statistically similar to that induced by S.S, PLLA as well the negative
control (P>0.05). Interestingly, the quantification of release of IL-8 and TNF-alpha
(Figure 5-5 C & D) by alloy showed significantly lower amounts when compared to S.S,
PLLA and negative control (P<0.05). This data corroborates the findings obtained from
platelet studies and confirms that the alloy elicits very low inflammatory response.
Although a more thorough investigation of alternate and synergistic inflammatory
pathways is required to completely understand the inflammatory response to the alloy,
the current data suggests preliminarily confirms the minimal inflammatory response
invoked by the Mg-Ca-Sr alloy.
Summary
Structural and architectural cues presented to cell systems via biomaterials elicit
important cellular responses in them including alteration of cell shape, alignment,
differentiation, growth and tissue formation14,258. The response between a biomaterial
and cell system can be varied across multiple length scales. The capability of cell
systems to sense microarchitectural, physical as well mechanical properties of the
substrate underneath falls under the meso scale (10 to over 100 µm). This capability is
made use of while designing biomaterials to modulate specific cellular response for a
variety of applications. Cardiovascular biomaterials can be designed to alleviate
complications arising from different cardiovascular diseases, such as atherosclerosis,
myocardial infarction, peripheral artery disease etc. Chapters 4 and 5 covered two
important classes of cardiovascular biomaterial applications, namely, small diameter
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vascular grafts and vascular stents, respectively. This work was focused on studying the
interaction of these biomaterials with in vitro models of the vascular environment to
understand and develop strategies to repair damaged vasculature and enable
regeneration.
The first part of the study was focused on development of a novel bilayered,
construct for application as a vascular graft. The graft was designed to mimic the
biological and architectural aspects of native blood vessels. The design consisted of two
layers, a solid luminal layer made of elastomeric, biodegradable POC and a fibrous
medial layer consisting of a mixture of collagen, POC and elastin. The rationale behind
the design is to provide a solid lumen for the formation of a monolayer of endothelial
cells and fibrous medial layer enriched with ECM proteins to enable formation of an
interconnected network of SMCs. The individual layers of the graft were separately
fabricated and assessed for cell attachment and proliferation. The constructs
mechanical properties were also assessed. Overall the results showed that the graft
was biocompatible and was able to support attachment and growth of vascular cells.
More interestingly, the mechanical properties evaluation revealed that the compliance of
the fabricated graft matched native artery values. This property is very important to
prevent implant failure due to mismatch in mechanical properties195,203,204. This graft is
currently being developed to recruit circulating stem cells to promote integration into
vasculature.
The second part of the study evaluated the hemocompatibility of an Mg-Ca-Sr
alloy to determine its utility as a candidate material for vascular applications. Overall,
this study showed that the alloy is biocompatible and invokes very low cytotoxic
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response when cultured with a mature vascular endothelial cell population. The alloy
sample was hemocompatible and exhibited excellent thromboresistance as indicated by
low hemolytic percentage as well as delayed and reduced clotting when in contact with
whole blood. The alloy also did not cause significant platelet adhesion and activation.
Finally, it also does not induce significant release of pro-inflammatory factors confirming
minimal inflammatory response. This data collectively confirms the suitability of the Mg-
Ca-Sr alloy for potential application as a vascular stent material. The studies combined
investigated the effect of biomaterial structural cues on the meso scale on their ability to
be used in vascular environments. The studies involved evaluating some important
characteristics of these materials in terms of their biocompatibility, hemocompatibility,
ability to support vascular cell types, mechanical and biochemical behavior. These
factors are essential in preventing vascular damage due to complications such as
thrombosis, toxicity as well as, fostering healthy remodeling and regeneration of
vasculature. Combined, these strategies are useful for developing suitable biomaterials
for vascular engineering.
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CHAPTER 6 MACROSCALE SYSTEMIC CUES- PART 1- EFFECT OF SIMULATED
MICROGRAVITY ON DIFFERENTIATION AND DOWNSTREAM FUNCTION OF PORCINE BLOOD DERIVED ENDOTHELIAL PROGENITOR CELLS
Introduction
Mammalian cells, especially stem cells have been known to be highly sensitive to
environmental cues that are provided to them128,259,260. These cues can be a variety of
different factors such as pH, temperature, pressure, nutrient supply and waste removal,
mechanical, electrical and magnetic stimuli, etc. There is a significant amount of
evidence in literature suggesting that physical stimuli may affect gene expression and
significantly increase the biosynthetic activity in a range of different cells176,261,262. The
ability of physical stimuli to modulate favorable cell response has caused the
development of functional simulations systems to create models for tissue engineering
in vitro. Besides providing tissue engineering solutions, these studies are also important
to understand the effect of mechanotransduction on cellular systems.
This is important especially in vascular tissue engineering since the endothelium
has been shown to respond to physiological mechanical cues, such as shear stress,
cyclic stretch, and alterations to gravity which cause them to undergo significant
morphological and functional changes129,143,260,263,264. One such specialized macro
environmental model is microgravity in space flight, which provides a unique opportunity
to study the effect of gravitational changes on cells.
Within the past 50 years of human space flight missions, several physiological
threats to astronauts have been identified. These include: losses in bone mineral
density, muscle atrophy, shifts in the distribution of body fluids, and cardiovascular
deconditioning.135,265,266 In general, these clinical manifestations are a result of what is
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termed as cardiovascular deconditioning or the inability of the cardiovascular system to
function properly. There has been a significant amount of research focused on studying
the negative effects of cardiovascular degeneration and the development of possible
strategies and countermeasures to address them267,268.
Proper function of the cardiovascular system is largely dependent on the
maintenance of healthy functioning vascular cells, namely endothelial cells that line the
vasculature and are in direct contact with circulating blood169,269. The endothelial cells
(ECs) that line the vasculature form a thin layer called the endothelium, and this layer of
the vasculature plays a critical role in regulating blood flow, vascular permeability, and
maintaining an actively anti-thrombogenic surface. The ability of the endothelium to
perform these functions is largely governed by their highly regulated secretory
capabilities270,271.
EC dysfunction has been implicated as an important factor in weightlessness-
induced cardiovascular deconditioning143. In addition to resident ECs, circulating
progenitor cells recruited from the bone marrow into circulation are also involved in
maintaining vascular homeostasis272,273 . While great attention has been placed on
understanding the effects of weightlessness on the endothelium, much less work has
been done to study the effect of microgravity on the “repair cells”, circulating endothelial
progenitor cells (EPCs).These EPCs reside in the bone marrow and are released into
circulation upon stimulation effected by specific triggers270,273–275. Once in circulation
they can integrate into the vasculature, differentiate into endothelial cells, and
participate in the repair of damaged vascular tissues. Recent work has shown that the
amount and function of EPCs are both significantly impaired due to the incidence of
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cardiovascular diseases. Interestingly, the number as well as the health of circulating
EPCs have been identified to be indicators of the onset as well as the progression of
cardiovascular diseases276–278. Additionally, due to their natural ability to act as ‘repair’
cells for the cardiovascular system, several in vitro studies and subsequent animal
studies have exploited their therapeutic potential279,280. Recently these cells were also
used in clinical studies with human subjects to alleviate different cardiovascular
conditions with promising results281.
A common feature of the effects of altered gravitational conditions on different
cellular systems is that the affected tissues rely on tissue-specific stem cells for
regeneration and repair. Understanding and replicating the behavior in such a condition
such as altered microgravity can lead to development of stem cell based therapy for
cardiovascular degeneration on earth. A recent study by Chiu et al. has shown that
human cord blood stem cells when cultured under simulated microgravity proliferated
and trans-differentiated into 3D structures resembling vascular tubules while exhibiting
vascular endothelial phenotype expressions140. Many such studies on the effect of stem
cells under alternate culture environments especially simulated microgravity is currently
being studied to develop regenerative strategies. Current research is also studying the
hypothesis that cardiovascular deconditioning due to space travel may potentially trigger
faster maturation of stem cells and therefore enhancing their regenerative potential.
The objective of this study is to fill the gap in knowledge regarding the effects of
microgravity on the downstream function of EPCs to understand their role in
cardiovascular repair during space travel and to propose a stem cell based therapeutic
strategy to counteract the negative effects and regenerate vascular tissue. More
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specifically this work is targeted towards studying the effect of altered gravity conditions
on the differentiation of EPCs to ECs and the downstream functionality of the ECs
themselves.
Towards this goal, this study was conducted in two parts, the first part focused
on understanding the effect of simulated microgravity on the differentiation and function
of EPCs after being pre-exposed to simulated microgravity while the second part
involved investigating of differentiated EPCs when cultured under simulated
microgravity conditions. These two parts will be covered in Chapters 6 and 7
respectively. With these goals in mind, the following specific aim was established and
studied in this work.
Specific Aim 3
Study and understand the effect of simulated microgravity on the function of
vascular stem cells to assess their potential to be used as a therapeutic strategy for
vascular repair and regeneration.
Before the methodology and results of the ensuing study can be discussed, it is
important to establish a definition and understanding of simulated microgravity and the
bioreactor culture set up that was used to generate this condition. Microgravity refers to
the condition where the effect of gravitational pull on objects seems to be diminished in
space265. Although the golden standard to understand effect of this condition on cell
systems is by studying their behavior in real space travel, spaceflight missions are often
very rare and expensive. Therefore, for preliminary studies, researchers have
developed ground-based models that can simulate this condition on Earth to prepare
their future space missions. Many strategies and devices with considerable success in
simulating microgravity have been developed for tissue engineering purposed119.
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In this study a NASA designed, rotating wall vessel (RWV) bioreactor that mimics
certain functional aspects of microgravity on earth (Synthecon, Houston, TX) was
utilized. The RWV was designed and developed at NASA’s Johnson Space Center to
simulate the effects of microgravity on cells in a ground-based culture system134,282.The
bioreactor set-up consists of a cylindrical vessel/chamber that rotates around its
horizontal axis with a coaxial tubular silicon membrane for gas exchange. (as shown in
figure below) 119,126,127,129
Figure 6-1. RWV bioreactor set-up (A) Digital image of Synthecon Slow Turning Lateral Vessel with rotor base, (B) RWV bioreactor detachable chamber.
To study the effect of this simulated microgravity on EPCs a porcine model was
chosen since it is widely preferred as a model system for vascular research due to its
physiological and hemodynamic similarities to the human vascular system283.
Additionally, a reproducible and consistent method to induce differentiation of porcine
blood derived EPCs has been reported85 . Lastly, the effect of altered gravity culture
conditions on the function of mature porcine endothelial cells has been studied in the
past284. However, the effects of short term hypogravity or microgravity exposure, on
circulating progenitor cells have not been reported before. This chapter describes the
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effect of simulated microgravity on the differentiation of circulating progenitor cells
derived from pig blood along an endothelial pathway. This work also describes the
effect of short term exposure to simulated microgravity, on the downstream functionality
of the differentiated endothelial cells such as release of soluble factors, prostaglandin
E2 (PGE2) and nitric oxide (NO), proliferation, migration and ability to regulate clotting
mechanisms. This study is the first step towards understanding the functionality of
EPCs in altered condition to assess their ability to be used therapeutically to repair and
regenerate vasculature.
Methods and Materials
Isolation of Peripheral Blood Mononuclear Cells from Porcine Blood
About 50 ml of circulating peripheral blood was collected from the exterior jugular
vein of domestic pigs (Department of Animal Sciences, University of Florida) in the
presence of acid citrate dextrose anticoagulant. All blood collection procedures were
performed in accordance with the rules and regulations of the Institutional Animal Care
and Use committee at the University of Florida. Within 4h of blood collection, the
peripheral blood mononuclear cells (PBMNCs) were isolated via density gradient
centrifugation as previously described 85,270. An average density of 140-160 million
mononuclear cells obtained from the isolation process were suspended in endothelial
cell growth media (EGM-2) (Lonza, Walkersville, MD) supplemented with 20% fetal
bovine serum. Porcine aortic endothelial cells (PAECs) (Cell Applications, SanDiego,
CA) between passages 4-6 cultured in the EGM-2 served as a positive control for
mature EC function and response.
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Rotating Wall Vessel Bioreactor System
The culture of cell systems in actual microgravity can be achieved only through
parabolic flights, rockets, space crafts, or laboratories available on the International
Space Station (ISS). However, the ability to culture various cell systems in real
microgravity is limited by the cost involved in performing these experiments on space
missions. Therefore a variety of devices have been designed to mimic some aspects of
real microgravity on earth 119. In this study, some aspects of simulated microgravity (µg)
conditions were established using a RWV bioreactor consisting of a rotor base and
detachable cell culture chambers (Figure 6-1 A). The chamber consists of a silicone
membrane for gas exchange as shown in Figure 6-1 B. Using this set up, the time-
averaged gravitational vector acting on these cellular assemblies is reduced to about
10−2 g. For simplicity, this condition is referred to as simulated microgravity. The validity
of the RWV model to simulate certain aspects of weightlessness has been verified and
reported extensively in literature 119,127,136. Additionally, the use of RWV to simulate
aspects of microgravity has been confirmed and validated by experiments in real
microgravity in the past 133,135,285,286.
Microgravity Culture of PBMNCs
The bioreactor set up as described above was placed in a cell culture incubator
maintained at 37°C and 5% carbon dioxide. The experimental set-up included 1)
PBMNCs suspended (18x106 cells in 50ml of EGM2) inside the bioreactor chamber for
microgravity exposure; and 2) PBMNCs seeded at a density of 2-3x105 cells/cm2 onto
fibronectin coated tissue culture polystyrene (TCP) multiwell plates (BD Biosciences,
Franklin Lakes, NJ) to be cultured statically with no simulated microgravity exposure. To
maintain conditions of microgravity the bioreactor chambers containing PBMNCs were
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rotated at a minimum speed of 5rpm. Following 6, 12 or 18h-µg exposure, the PBMNCs
were collected, then seeded onto fibronectin coated TCP plates and cultured statically
for 7-10 days.
Differentiation, Colony Formation and Immunocytochemistry
The statically cultured cells, both those pre-exposed to microgravity and those in
normal gravity for the entirety, were visually assessed to observe colony formation
every day for 8 days using an inverted phase-contrast microscope (Nikon TE2000U,
Melville, NY). The first appearance of densely packed, isolated, cobblestone
morphology, endothelial-like cells in each field of view was recorded as indicative of EC
differentiation. The number of colonies and the rate of new colony formation was
recorded.
PE-like cells were examined using immunohistochemical analysis to confirm
endothelial cell phenotype. Briefly, cells were fixed in 4% paraformaldehyde,
permeabilized, if required, using 0.2% Tween-20 and then blocked using 3% bovine
serum albumin. Following which, the cells were incubated overnight at 4⁰C with a
primary antibody (diluted 1:100) against endothelial nitric oxide synthase (Abcam,
Cambridge, UK) and CD31 (diluted 1:100) (AbDSerotec, Raleigh, NC). Subsequently,
the probed cells were incubated with a fluorescein isothiocyanate tagged isotype-
matched secondary antibody (Santa Cruz Biotechnology, Santa Cruz, CA) and nuclei
were stained using Hoechst 33342 (ThermoScientific, Waltham, MA). Fluorescent
images were captured using an inverted fluorescent microscope (Nikon TE2000U).
Endothelial-Like Cell Migration and Proliferation
Cell migration was assessed using the established scratch assay 287. Briefly, PE-
like cells differentiated after 6, 12, and 18h of pre-exposure to simulated microgravity as
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well as those cells differentiated in normal gravity, were cultured to approximately 90%
confluence on multi-well TCP plates in the culture media as described above. A linear
scratch was made through the monolayer using a pipette tip to create a “cleared” area,
then the cells were incubated at 37° C for 18h to allow for migration. Digital images of
the scratch surface were taken before and after incubation and the percentage area
covered by the migrating cells was quantified using ImageJ software version 1.46.
Cell proliferation was assessed via changes in DNA concentration from cultured
cells over time, quantified using Quant-it picogreen dsDNA assay kit (ThermoScientific).
Briefly, PE-like cells differentiated after 6, 12, and 18h of pre-exposure to microgravity
as well as normal gravity cultured controls, were seeded on 12-well TCP plates at a
seeding density of 2500 cells/cm2. At 24hr increments over 4 days, the cells were lysed
to collect and quantify the intracellular DNA. The concentration of DNA contained within
the lysates were normalized by subtracting the initial seeded DNA concentration from
the concentration at each time point for each sample.
Nitric Oxide and Prostaglandin Secretion
NO and PGE2 secretion was evaluated in PE-like cells differentiated after 6, 12,
and 18h of pre-exposure to simulated microgravity as well as the cells differentiated
under normal gravity. The amount of total NO and PGE2 secreted by the cells into the
culture medium over a period of 24h for each cell population was determined using the
total NO assay kit and Prostaglandin E2 Assay kit (R&D systems, Minneapolis, MN)
following manufacturer protocols. Briefly, the concentration of NO measured is based
on the enzymatic conversion of nitrates to nitrites and the detection of total nitrite
concentration via the Greiss reaction. PGE2 concentration was assessed via enzyme
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linked immunosorbent assay. The concentration of both NO and PGE2 was normalized
to cell number and expressed in terms of concentration per cell.
Clotting Kinetics
Following previously reported methods, the kinetics of both plasma and whole
blood clot formation was assessed 85,270. Briefly, plasma and whole blood were exposed
to monolayers of PE-like cells differentiated after 6, 12, and 18h of pre-exposure to
simulated microgravity as well as cells differentiated under normal gravity, and the
kinetics of clot formation were reported. In this study, PAEC served as a positive control
for anti-thrombogenic endothelial cell function, and bare TCP serves as a negative
control representing the kinetics of uninhibited plasma and blood clotting.
Statistical Analysis
All data are reported as means +/- standard deviations. Numerical means were
compared using a student 2-sample t-test. Analysis of variance (ANOVA) with Newman-
Keuls multiple comparison test post hoc analysis was employed to identify significant
differences between 3 or more means. P-values <0.05 was reported as significant.
Results and Discussion
This study reports the effect of early exposure to simulated microgravity on the
differentiation and downstream function of porcine blood derived progenitor cells.
Although multiple studies have investigated and reported the effect of altered gravity
conditions on the behavior of adult endothelial cells, there are limited reports on its
effect on circulating progenitor cells143. Since these cells have been shown to be
involved in vascular repair mechanisms in the body, it is important to understand the
functionality of these cells when exposed to short term microgravity, to better
understand effects of cardiovascular repair in space and develop stem cell based
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therapies to repair, preserve and regenerate vasculature. Previous studies have
reported significant alterations to morphological characteristics and changes to gene
and protein expression in mature endothelial cells within a few minutes to a few days of
microgravity exposure143,288–290. Moreover, the changes in expression of immunological
cytokines in PBMNCs and the recovery of their functions on short term exposure to
simulated microgravity ranging from within a few hours to up to 3 days has been
reported before291.
In this study, circulating progenitor cells contained within the PBMNC fraction of
porcine whole blood were exposed to simulated microgravity, using a RWV bioreactor
designed originally by NASA to simulate reduction in earth’s gravity (1g) to micro-scale
(µg) on cultured cells. The time-averaged gravitational vector acting on these cells is
reduced to about 10-2 g using this set-up. After exposing the circulating progenitor cells
to 6, 12 and 18h of microgravity, the cells were then cultured on fibronectin coated TCP
under static conditions and their ability to differentiate into endothelial cells was
assessed by the formation of colonies with EC specific ‘cobble-stone’ morphology.
(Figure 6-2A) In all cases the observation for colonies was concluded after 8 days after
which it was difficult to discern discrete isolated colonies. Cells exposed to microgravity
for longer than 18h showed significantly reduced or no colony formation when cultured
on fibronectin coated TCP. This indicated the existence of a possible threshold time of
microgravity exposure beyond which the differentiation ability of these PBMNCs may be
permanently impaired. On assessing the kinetics of EC colony formation, it was
observed that the average rate of endothelial cell colony formation decreased with
increased time of exposure to microgravity between 3-6 days of culture (Figure 6-2 B).
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While the normal gravity showed the highest rate of colony formation as indicated by an
average of 10 colonies/day.
Figure 6-2. Differentiation and characterization of PE-like cells. Representative phase contrast image of PBMNC derived endothelial colony formation after 4 days of static culture (4x), Scale bar 200 µm. The region indicated by the dotted circle shows a newly differentiated PE-like cell colony (B) Graph showing the kinetics of colony formation of microgravity exposed PE-like cells vs cells maintained in normal gravity. (0h µg: 10 colonies/day, 6h and 12h µg: 8 and 6 colonies/day vs 18h µg: 2.67 colonies/day, respectively) (C) Representative phase contrast image of characteristic endothelial ‘cobble-stone’ morphology of PE-like cells (20X). Representative 20X fluorescent images of positive staining of both 12h µg exposed as well as normal gravity exposed PE-like cells for (D-E) endothelial nitric oxide synthase (F-G) CD31 shown in green, nuclei were counterstained with Hoechst 33342 shown in blue. Scale bar is 100 µm.
The control population of PBMNC, cultured under normal gravity conditions
formed colonies at the earliest period of 3 days compared to microgravity exposed
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populations which did not form any colonies as early as the third day of culture. The
endothelial phenotype of both the microgravity exposed and normal gravity cultured
cells were confirmed by their characteristic cobble stone morphology (Figure 6-2 C) and
positive staining for eNOS (Figure 6-2D-E) and CD31 (Figure 6-2 F-G).
In the past, the decrease/delay in colony formation of ECs from EPCs have been
associated to pre-existing cardiovascular complications292,293.This delay in differentiation
indicates a possible detrimental effect of simulated microgravity on the onset of
differentiation of EPCs to ECs that may affect the regulation of cardiovascular health by
these cells. Also, the ability of simulated microgravity to inhibit bone marrow
hematopoetic stem cell differentiation has been confirmed by Plett et al. (2004)294.
Further, the effect of this altered condition on the downstream functional abilities
of the differentiated endothelial-like cells that were pre-exposed to microgravity was
investigated. The differentiated cells were assessed for their proliferation and migratory
ability. Migration was assessed by observing and quantifying the migration of cells into
an area cleared of cells (Figure 6-3 A). The quantification of percentage area of the
scratch covered by the migrating cells (Figure 6-3 B), showed that the 6h-µg exposed
and normal gravity cultured PE-like cells migrated the most followed by 12h-µg exposed
cells (P<0.05). Finally, the 18h-µg exposed PE-like cells migrated the least (P<0.05),
indicating a direct decrease in migration ability of the EC with increased time of
exposure of the progenitor cells under simulated microgravity.
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Figure 6-3. Proliferative and migratory functions of simulated microgravity pre-exposed PE-like cells (A) Representative phase contrast images showing migration of microgravity exposed PE-like cells over 18 h into a scratch made in the cell monolayer vs the migratory response of cells maintained in normal gravity. Representative results showing (B) Quantification of % area covered by migration of cells into the scratch. Data expressed as mean +/- SD, N=6; (6h µg exposed cells: ~91+/-8 %; 0h µg ~93+/-10%; 12h µg: ~78+/- 4%; 18h µg: ~38+/-3%) (C) Proliferation kinetics of microgravity exposed PE-like cells vs cells maintained in normal gravity as measured by DNA concentration in cell lysates. (Average initial rate of proliferation reported as rate change in concentration of DNA content (pg/day), 6h µg exposed cells: 2.64E+05 +/- 8.76E+03; 12 µg: 3.01E+05 +/- 2.74E+04, 18 µg: 4.98E+05 +/- 1.49E+03, 0h µg: 6.93E+05 +/-9.40E+02). Data expressed as mean +/- SD, N=3. Data statistics based upon one-way analysis of variance (ANOVA) coupled with Tukey’s multiple comparison test, N.S. indicates no significant difference (P > 0.05) and * indicates a statistically significant increase (P<0.05).
The effect of pre-exposure to µg on the proliferative capacity of the EC-like cells
was also assessed by quantifying the change in DNA over time in cultured cells. The
data show that 6 and 12h of pre-exposure to µg altered the initial rate of proliferation
most significantly. (Figure 6-3C). In contrast, the PE-like cells pre-exposed to
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microgravity for 18h proliferated faster than both (Figure 6-4C, P<0.05). As expected,
the PE like cells differentiated under normal gravity showed no inhibition in proliferation,
but rather showed the greatest amount of proliferation (P<0.05).
These data collectively show that both the migratory ability and proliferative
ability of ECs are impaired by pre-exposure of the EPCs to microgravity. While
increased exposure time to µg impaired the migration of the microgravity exposed cells,
the proliferation of these cells seems to be enhanced by increased time of microgravity
pre-exposure. This indicates a possible recovery of the proliferative ability of the PE-like
cells on longer times of exposure of the EPCs to simulated microgravity. A similar
recovery of immune functions of human PBMNCs after their short term exposure to
simulated microgravity has been reported291. Further investigation of this recovery
mechanism may result in the development of possible countermeasures for treating
cardiovascular deconditioning to enhance recovery291,295,296.
A healthy endothelium can control and regulate thrombogenic properties and
clotting cascades in the vasculature by releasing anti-thrombogenic factors into the
blood stream. The ability of porcine EPC derived ECs to inhibit clotting mechanisms has
been shown before85,270,271,297. Hence it is important to assess the ability of ECs derived
from microgravity pre-exposed progenitors to regulate clotting mechanisms. To evaluate
the anti-thrombogenicity of these cells, their ability to release soluble factors such as
NO and PGE2 into the surrounding medium was first investigated. Prostacyclin and
Prostaglandin E2 have been known to be released from the endothelium after
stimulation by vasoactive agents and thrombin243. Prostacyclin is predominantly
released by the endothelium of large diameter blood vessels, while the ECs from
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smaller diameter blood vessels release predominantly prostaglandin E2298. Both of
these factors contribute to platelet inactivation and regulate clotting mechanisms243.
Another important molecule that regulates clotting mechanisms in the body is nitric
oxide, which is produced by endothelial nitric oxide synthase (eNOS)243,299. Both these
molecules synergistically regulate anti thrombogenicity in the body. The results show a
decrease in secretion of soluble PGE2 by 6h-µg exposed cells (P<0.05) while there
was no significant difference in the levels of PGE2 between the other microgravity
exposed cells as well as the control cells cultured under normal gravity (Figure 6-4 A).
Figure 6-4. Representative results of quantification of release of soluble factors. (A) Prostaglandin E2 over 24 h by microgravity exposed PE-like cells vs cells maintained in normal gravity (0h µg). (Data shown are Mean +/- SD; n=4). (PGE2 concentration in pg/ml/cell; 6h µg: 4.06E-06 +/- 1.71E-06, 12h µg: 1.86E-05 +/- 4.8E-06; 18h µg: 1.84E-05 +/- 4.84E-06; 0h µg: 1.34E-05 +/- 3.33E-06) (B) Nitric Oxide over 24 h by microgravity exposed PE-like cells vs cells maintained in normal gravity (0h µg). (NO concentration; in µmol/L/cell, 6h µg: 1.66E-06+/- 7.12E-07, 12h µg: 3.135E-05 +/- 3.29E-06; 18h µg: 2.88E-05 +/- 2.89E-06; 0h µg: 2.75E-05 +/- 6.35E-06). Data shown are Mean +/- SD; n=3. Data statistics based upon one-way analysis of variance (ANOVA) coupled with Tukey’s multiple comparison test, N.S. indicates no significant difference (P > 0.05) and * indicates a statistically significant increase (P<0.05).
The assessment of release of soluble NO indicated a similar trend where the
amount of NO released by 6h-µg exposed cells was significantly lower (P<0.05) when
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compared to other microgravity exposed cells as well as normal gravity cultured cells
(Figure 6-4B).
This data shows the ability of short term microgravity exposure of EPCs to inhibit
the regulation of release anti thrombogenic molecules into the surrounding medium.
However, with increased time of microgravity exposure, a potential recovery effect is
also observed where the 12h and 18h-µg exposed cells secrete NO and PGE2 in levels
comparable to the cells cultured under normal gravity. Previous reports have shown an
increase in the release of both the factors by mature ECs when cultured under
microgravity conditions 143. However, the effect of short term microgravity exposure of
progenitor cells, to alter the release of soluble factors by ECs derived from them have
not been reported before. This data is in coherence with whole blood clotting and
recalcified plasma clotting studies subsequently performed on these cells. Plasma
recalcification clotting profiles indicated that as expected the TCP surface induced
highest rate of plasma clotting within the first 90 seconds followed by 6h µg exposed
PE-like cells (Figure 6-5 A, P<0.05). Following the clotting rate for 6h-µg exposed cells
are the clotting times for the 12h and 18h-µg exposed cells. Normal gravity cultured
cells clotted the next while the last to clot was the plasma exposed to the PAEC control
cells (P<0.05). Similarly, the whole blood clotting kinetics showed that after 20 min the
average rate of whole blood clot formation was the highest for TCP surface (Figure 6-5
B, P<0.05) followed by whole blood exposed to 6h, 12h and 18h-µg exposed cells,
whose average clotting rates were lower than TCP control (Figure 6-B). While the whole
blood exposed to PAEC and cells cultured under normal gravity showed the greatest
inhibition of clotting.
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Figure 6-5. Representative result showing quantification of blood clotting studies. (A) recalcified plasma clotting, rate of plasma clotting expressed as change in absorbance/min, TCP: 0.537 +/- 1.26E-03, 6h µg: 0.510 +/- 1.25E-03, 12h µg: 0.368 +/- 2.5E-03, 18h µg: 0.3635 +/- 1.5E-03, 0h µg: 0.335 +/- 1.02E-03, PAEC: 0.1567 +/- 2.14E-03 (B) whole blood clot formed over time on monolayers of microgravity exposed PE-like cells vs cells maintained in normal gravity, PAECs and TCP. Error bars have been omitted for visual clarity. Average rate of clotting in g/min on TCP: 2.733E-03 +/- 4.42E-04; 6h µg cells: 2.08E-03 +/- 1.71E-04; 12h µg: 1.89E-03 +/- 4.32E-04; 18h µg: 1.633E-03 +/- 3.93E-05; PAEC: 1.175E-03 +/- 1.37E-04 and 0h µg: 1.288E-03 +/- 2.92E-04. Data statistics based upon one-way analysis of variance (ANOVA) coupled with Tukey’s multiple comparison test, N.S. indicates no significant difference (P > 0.05) and * indicates a statistically significant increase (P<0.05).
This data collectively shows that exposure to microgravity in the early progenitor
stages negatively influences the anti-thrombogenic function of PE-like cell relative to
those differentiated under normal gravity. The data shows that within the 3 time points
of µg exposure, the longer exposure times showed greater inhibition of clotting, which
may be attributed to the increased ability to release soluble anti-thrombogenic factors,
NO and PGE2. Although previous reports have shown the ability of simulated
microgravity to directly affect platelet functions 144, to the knowledge, the effect of
simulated microgravity on anti-thrombogenic properties of ECs, such as inhibition of
blood and plasma clotting has not been reported so far. This work provides an insight
into the microgravity mediated alterations to regulation of clotting mechanisms by
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circulating stem cell derived endothelial cells which is important to assess their role in
cardiovascular regeneration during space travel. As an extension of this study, Chapter
7 deals with the functionality of blood derived endothelial cells when cultured under
simulated microgravity after differentiation.
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CHAPTER 7 MACROSCALE SYSTEMIC CUES- PART 2- EFFECT OF SIMULATED
MICROGRAVITY ON FUNCTION OF HUMAN BLOOD DERIVED ENDOTHELIAL CELLS
Introduction
Within the past 50 years of human space flight missions, several physiological
threats to astronauts have been identified. These include: losses in bone mineral
density, muscle atrophy, shifts in the distribution of body fluids, and cardiovascular
deconditioning135,265,266. There has been a significant amount of research focused on
studying the negative effects of cardiovascular degeneration and the development of
possible strategies and countermeasures to address them267,268. Proper function of the
cardiovascular system is largely dependent on the maintenance of healthy functioning
vascular cells, namely endothelial cells that line the vasculature and are in direct contact
with circulating blood169,269. The endothelial cells (ECs) that line the vasculature form a
thin layer called the endothelium, and this layer of the vasculature plays a critical role in
regulating blood flow, vascular permeability, and maintaining an actively anti-
thrombogenic surface. The ability of the endothelium to perform these functions is
largely governed by their highly regulated secretory capabilities270,271. The endothelium
has also been shown to respond to physiological mechanical cues, such as shear
stress, cyclic stretch, and alterations to gravity which cause them to undergo significant
morphological and functional changes129,143,260,263,264.
Supplementary to ECs, circulating progenitor cells (CPCs) also play a role in
regulating the normal functioning of the vasculature. These CPCs reside in the bone
marrow and are released into circulation upon stimulation270,273–275. Once in circulation
they can integrate into the vasculature, differentiate into endothelial cells, and
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participate in the repair of damaged vascular tissues. The long-term goal of this study is
to fill the gap in knowledge regarding the effects of microgravity on the downstream
function of CPCs after they have differentiated into ECs. These stem cells serve as a
source of healthy vascular cells and play a very important role in the maintenance of
vascular homeostasis271,272,277,279. While the effect of weightlessness on mature ECs
has been studied extensively, there is a dearth of information regarding the effect of
microgravity on ECs derived from the circulating stem/progenitor cell population. Hence,
it is important to study the effect of simulated microgravity on the functions of these cells
to understand their implications during space flight and their possible role in the onset of
cardiovascular dysfunction.
As a first step towards establishing an understanding of the effect of simulated
microgravity on circulating progenitor derived ECs, modeled microgravity was studied
using a ground based rotating wall vessel (RWV) bioreactor. The RWV was designed
and developed at NASA’s Johnson Space Center to simulate the effects of microgravity
on cells in a ground-based culture system134,282. The bioreactor set-up used in this study
consisted of a cylindrical vessel/chamber that rotates around its horizontal axis with a
coaxial tubular silicon membrane for gas exchange 119,126,127,129. The results obtained
using simulated microgravity conditions utilizing the RWV have been confirmed and
validated by experimental models in space flights to outer space133–135,285.
This study reports the effect of simulated microgravity on the downstream
functions of human blood derived ECs such as: proliferation, release of soluble factors
nitric oxide (NO), interleukin-6 (IL-6), tissue plasminogen activator (tPA), gene
expression of HSPA4 (gene encoding heat shock protein 70 kDa A) and anti-
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thrombogenic potential. More specifically this work is focused on the ability of these
cells to adapt to hypergravity conditions and modulate these functions after short term
culture and expansion under simulated microgravity conditions. To the knowledge, this
is the first report that provides valuable insight into the functionality of human blood
progenitor cell derived ECs following exposure to simulated microgravity. This work
provides a foundation for future work to elucidate vascular repair mechanisms taking
place in the body in outer space.
Materials and Methods
Isolation of Peripheral Blood Mononuclear Cells from Human Blood
About 45 ml of circulating peripheral blood was collected from healthy donors by
venipuncture in the presence of heparin (158 USP) anticoagulant in vacutainer tubes
(Thermo Fisher Scientific, Waltham, MA). For clotting studies alone, blood was collected
in vacutainer tubes containing Acid Citrate Dextrose (ACD) solution (Thermo Fisher
Scientific) as the anti-coagulant. All blood collection procedures were performed in
accordance with the rules and regulations of the Institutional Review Board at the
University of Florida. Within 2h of blood collection, the peripheral blood mononuclear
cells (PBMNCs) were isolated using density gradient centrifugation as described in
detail elsewhere85,270. Approximately 80 million PBMNCs were obtained from the
isolation process, then subsequently suspended in endothelial cell basal media-2 (EBM-
2) (Lonza, Walkersville, MD) supplemented with 20% fetal bovine serum, human
fibroblast growth factor (hFGF), vascular endothelial growth factor (VEGF), human
epidermal growth factor (hEGF), human recombinant insulin-like growth factor (R3-IGF-
1), ascorbic acid, gentamicin, amphotericin, and heparin. Henceforth, this fully
supplemented EBM-2 is referred to as endothelial cell growth media-2 (EGM-2).
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Culture and Differentiation of Endothelial Progenitor Cells
The PBMNCs were plated at a density of 2-3X105 cells/cm2 onto fibronectin-
coated tissue culture polystyrene plates (BD Biosciences, Franklin Lakes, NJ) in
complete cell culture media as described above. The culture was maintained for 2-6
weeks and the cells were visually assessed to observe colony formation every day
using an inverted phase-contrast microscope (Nikon TE2000U, Melville, NY). The first
appearance of tightly packed, cobblestone morphology, human endothelial-like (HE-like)
cells in a given field of view was recorded as indicative of EC differentiation. The
differentiated cells were sub cultured and utilized within passages 1-3. Human umbilical
vein endothelial cells (HUVECs) (Lonza) were used as a positive control for mature EC
functionality. HUVECs were cultured in EGM-2 (with only 2% FBS) and utilized within
passages 3-5.
Immunocytochemical Assessment of Differentiated Endothelial Cells
The endothelial cell phenotype of the differentiated cells was confirmed using
immunocytochemical analysis (herein after referred to as human endothelial-like cells
(HE-like cells)). Briefly, HE-like cells were fixed in 4% paraformaldehyde, and then
blocked using 3% bovine serum albumin for 1h. The cells were incubated overnight at
4⁰C with a primary antibody specific to von Willebrand factor (vWF) (diluted 1:100)
(Abcam, Cambridge, UK) and CD31 (diluted 1:100) (Abcam). The cells were
subsequently incubated with a fluorescein isothiocyanate tagged isotype-matched
secondary antibody (Santa Cruz Biotechnology, Santa Cruz, CA) and the nuclei were
counter-stained using Hoechst 33342 (Thermo Scientific). Fluorescent images were
obtained using an inverted fluorescent microscope (Nikon TE2000U) (Nikon, Melville,
NY).
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Simulated Microgravity Culture of HE-Like Cells
Cytodex 1 beads (Sigma Aldrich, St. Louis, MO) were swollen and sterilized
using manufacturer’s instructions.127,282,288 The beads were coated with fibronectin
(25ng/ml) (Sigma Aldrich) to promote cell adhesion. HE-like cells obtained using the
above-mentioned procedure, were seeded on the beads at a density of 3 X 106 cells per
250 mg of beads in ultra-low attachment cell culture plates. The culture was maintained
for up to four days until complete confluence of cells on beads was obtained. The cell
coated beads were then transferred into the bioreactor chambers (Synthecon, Houston,
TX) as shown in Figure 7-1 A&B and cultured under simulated microgravity conditions.
Figure 7-1. Bioreactor set-up and operation. (A) Synthecon RWV bioreactor set-up (B) RWV bioreactor detachable chamber (C) Representative 20X phase contrast image of HE-like cells cultured on cytodex-1 microcarrier beads. Scale bar is 100 µm. In panel C, the arrows indicate cell covered beads.
Briefly, the bioreactor was set to rotate at an initial speed of 8.5 rpm, to suspend the
cells under simulated microgravity with minimum shear stress in a cell culture incubator
with 5% carbon dioxide (CO2). For all experiments, as a static control, the cell coated
beads were placed into the bioreactor chambers in a cell culture incubator, however, the
chambers were not rotating and thus were maintained under normal gravity.
Subsequently, the cultures were maintained for up to 6 days. About half of the cell
culture media in the chamber was replaced every other day.
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Cell Proliferation Assessment
Cell proliferation was assessed via changes in DNA concentration of HE-like
cells cultured both under simulated microgravity and normal gravity over time. The
dsDNA was quantified using Quant-it picogreen dsDNA assay kit (Thermo Scientific)
following manufacturer instructions. Briefly, HE-like cells on beads, exposed to
simulated microgravity were collected at 1, 2, 4, and 6 days of culture from the
bioreactor chamber. Correspondingly, samples were collected from normal gravity
controls at the same respective time points. Then, the cells were lysed using 0.1 %
triton-X to collect and quantify the intracellular DNA. The DNA concentrations of lysates
were quantified and plotted versus time.
Gene Expression
To determine the effect of culture conditions on gene expression of hsp70 by the
HE-like cells, quantitative reverse transcriptase polymerase chain reaction (qRT-PCR)
was conducted. After 2 days and 4 days of culture, total RNA was isolated from the HE-
like cells cultured under normal and reduced gravity conditions, using the RNeasy Mini
Kit (Qiagen, Valencia, CA) according to manufacturer instructions. Subsequently iScript
cDNA Synthesis Kit (Bio Rad, Hercules, CA) was used to generate first-strand cDNA.
The RNA sequences of human glyceraldehyde 3-phosphate dehydrogenase (GAPDH)
and heat shock protein 70 kDa A 4 (HSPA4) were identified and selected from the
National Center for Biotechnology Information (NCBI) gene database, and
corresponding primers for each of the genes were obtained from Integrated DNA
Technologies ( IDT, Coralville, IA). qRT-PCR was run on the Bio Rad CFX Connect (Bio
Rad, Hercules, CA) using Maxima SYBR Green (Thermo Scientific, Waltham, MA) to
quantify the expression of each target gene. Melt curve analysis was run to ensure
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amplification of non-specific PCR products did not affect the quantification results.
Results for gene expression were normalized using GAPDH. The data obtained was
derived from cells from a single donor population.
Endothelial Cell Secretory Function
The amount of nitric oxide (NO), interleukin-6 (IL-6), and tissue plasminogen
activator (tPA) secreted into the culture medium by the HE-like cells was assessed after
2 and 4 days of culture in both normal and simulated microgravity. Quantification was
obtained utilizing the total NO assay kit (Cayman Chemical, Ann Arbor, MI), Human IL-
6, and Human t-Plasminogen Activator/tPA Quantikine ELISA Kits (R and D systems,
Minneapolis, MN) respectively, following manufacturer protocols. Briefly, the
concentration of NO released into the medium was quantified based on the enzymatic
conversion of nitrates to nitrites and the photometric detection of total nitrite
concentration via the Greiss reaction. The concentration of NO was normalized to cell
density using the dsDNA concentration of the cells and expressed in terms of nitrite
concentration (µM) per pg of dsDNA, to draw comparisons between times in simulated
microgravity.
Plasma Clotting Kinetics
The plasma clotting kinetics of the HE-like cells was assessed using a modified
protocol previously reported230,270. Briefly, human allogenic platelet poor plasma was
obtained via centrifugation from whole blood. HE-like cells on beads cultured for 2 and 4
days, both under simulated microgravity as well as under normal gravity were exposed
to the 250 µl of the plasma recalcified by the addition of 0.025 M calcium chloride
(CaCl2). Following which, the absorbance at 405 nm was recorded using a Synergy H1
microplate reader (Biotek, Winooski, VT). As the plasma clots it becomes more opaque,
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thus an increase in absorbance with time is indicative of clot formation. HUVEC coated
microcarrier beads served as a positive control for anti-thrombogenic function of mature
ECs. The full kinetic profile of clot formation as well as the time taken to reach maximum
absorbance, indicative of full clotting, was recorded and reported for each
sample/control.
Statistical Analysis
All data are reported as mean +/- standard deviations. A student 2-sample t-test
was employed to compare two numerical means. Analysis of variance (ANOVA) with
Tukey’s multiple comparison post hoc analysis was used to analyze and identify
significant differences between 3 or more means. P-values <0.05 are reported as
significant.
Results and Discussion
The physiological threats faced by astronauts as a direct consequence of human
space flight have been identified and studied extensively in the recent past135,139,265. The
cardiovascular system of astronauts has been known to be affected by orbital space
flight with the changes manifesting as cardiovascular deconditioning135,266,267. From a
clinical perspective, deconditioning is observed as a: (1) decrease of circulating blood
and interstitial fluid volumes, (2) decrease of arterial blood diastolic pressure, (3)
decrease of ventricular stroke volume, (4) decrease of estimated left ventricular mass
and (5) a resetting of the carotid baroreceptors. The deconditioning effects of
microgravity manifest itself mostly upon the reentry to Earth268. Maintenance of vascular
homeostasis lies in the proper function of the vascular cells, specifically, the ECs that
line the vessel wall. This work is focused on understanding the effects of short term
simulated microgravity on the downstream functions of ECs derived from whole blood.
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Previous work has shown the exposure of mature ECs to simulated microgravity for as
little as a few minutes to a few days results in structural and morphological changes as
well as modifications to both gene and protein expression 143,263,285,288–290,300. These
results have been observed and validated in both space flight experiments as well as
ground based microgravity simulators as was used here. However, there is very limited
literature on the effect of altered gravity conditions on blood-derived circulating
endothelial progenitor cell (EPC) derived ECs. EPCs are a small population of
progenitor cells that reside in bone marrow and, upon stimulation, are released into
circulating peripheral blood. EPCs along with the endothelium are critical to the
maintenance of vascular homeostasis, which includes regulating the hemodynamics of
blood flow, vessel wall tone, and the secretion of important signaling molecules270,274,279.
Previous studies have shown the ability of circulating progenitor derived ECs to function
in a similar manner to mature ECs including their ability to regulate clotting mechanisms
in vitro85,230,243,270,299.85,230,243,270,299. Despite the conflicting reports on the suitability of
using EPCs for the study of cardiovascular diseases276, this data shows the promising
potential of EPCs to act as a repair cell population in cardiovascular regeneration.
Therefore, this work is an important stepping stone towards understanding the role
played by EPCs for the onset of cardiovascular deconditioning during space travel and
to develop possible cell based therapies. As a foundation for the work, the group has
previously shown the effect of simulated microgravity on the differentiation and
downstream anti-thrombogenic functions of porcine endothelial progenitor cells230.
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In this report, circulating progenitor cells contained within the PBMNC fraction of
whole blood were isolated then cultured in the presence of vascular endothelial growth
factor (VEGF) to induce differentiation along the endothelial lineage84,85,230,270.
Differentiation into endothelial phenotype was confirmed by the formation of colonies
with EC specific ‘cobble-stone’ morphology (HE-like cells) as shown in Figure 7-2 A.
Figure 7-2. Isolation of PBMNCs and differentiation into ECs: (A) Representative phase contrast image of PBMNC derived colony formation of HE-like cells, Scale bar 200 µm. (B) Representative phase contrast image of characteristic endothelial ‘cobble-stone’ morphology of HE-like cells in subsequent culture (20X). Representative 20X fluorescent images of positive immunocytochemical staining of HE-like cells for (C) CD31 and (D) von Willebrand factor, both shown in green, nuclei were counterstained with Hoechst 33342 shown in blue. Panel B-D Scale bar is 50 µm.
Colony formation was observed between 2-6 weeks of culture. In all cases the
observation for colonies was concluded after 50 days. The process was repeated for
three populations; each from a different donor. The data reported are representative of
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these three different populations. The endothelial phenotype of these cells were
confirmed and validated by the maintenance of the characteristic ‘cobble-stone’
morphology in subsequent culture as shown in Figure 7-2 B and positive
immunocytochemical staining for CD31 as shown in Figure 7-2 C, and vWF, as shown
in Figure 7-2 D both of which are specific for EC phenotype.
To simulate microgravity culture conditions, HE-like cells were seeded onto
fibronectin coated microcarrier beads. A RWV bioreactor system, designed originally by
NASA was used to simulate a reduction in time-averaged gravitational vector acting on
these cells, reducing it to about 10-2 g. For simplicity, this reduced gravity condition is
referred to as microgravity133,136,263. Although RWV systems do not mimic actual
microgravity as experienced by astronauts, however, it is a useful tool to provide insight
into the behavior of various cell systems in space127,282.The culture was maintained by
adjusting the rotational speed just enough to maintain the vessel wall and the medium
containing cells bound to microcarrier beads rotating at the same speed.
One of the primary and most important functions of the HE-like cells is their
ability to adapt to changing conditions and to continually proliferate. The effect of
simulated microgravity (µg) on differentiated HE-like cells and compared to the kinetics
of proliferation under normal gravity (1g) conditions. The proliferative capacity of the
HE-like cells was assessed by quantifying the change in DNA over 1, 2, 4 and 6 days of
culture. The results indicated the concentration of DNA in the normal gravity (1g)
cultured cells initially was higher than the simulated microgravity (µg) up to 2 days of
culture (p<0.05) as shown in, Figure 7-3. [dsDNA concentration (µg/ml) after 1 day: 1g
0.098 +/- 0.002, µg 0.073+/- 0.013 p<0.05; after 2 days: 1g 0.110+/-0.005, µg 0.093+/-
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0.007 p<0.05]. However, the cells cultured under simulated microgravity caught up by
day 4 and showed no significant difference in DNA content. [dsDNA concentration
(µg/ml) after 4 days: 1g 0.116 +/- 0.005; µg 0.113 +/- 0.003 not significant]. After 4 days,
the HE-like cells under simulated microgravity increased there rate of proliferation as
indicated by a higher concentration of DNA when compared to the normal gravity
control at day 6. [dsDNA concentration (µg/ml) after 6 days: 1g 0.151+/- 0.010, µg
0.173+/- 0.003 p<0.05].
Figure 7-3. Proliferation kinetics: Representative proliferation kinetics curve HE-like cells cultured under microgravity (µg) vs cells maintained in normal gravity (1g) as measured by DNA concentration in cell lysates. Data shown are Mean +/- SD; n=3.
The results showed an initial decrease in cell growth which was reversed over
increased time in culture. A similar trend in proliferation has been shown for a mature
EC population upon exposure to simulated microgravity285,301. This data indicates that
HE-like cells function similar to mature EC populations when exposed to simulated
microgravity, thus enhancing their ability to proliferate. Additionally, this trend also
indicates a potential recovery effect in the proliferative capacity of these ECs and HE-
like cells. Such a proliferation effect has been reported in PBMNCs 291 and other mature
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EC populations136,263,285,302 .This trend also indicates the potential ability of these ECs to
sense the change in gravitational conditions as stress and their ability to recover from
this stress by enhancing proliferation and cell growth302.
Further the ability HE-like cells to sense simulated microgravity as a stressful
condition and adapt to it by regulating the levels of genes encoding stress proteins such
as hsp70 was investigated. The regulation of gene expression of HSPA4 via RT PCR.
The HSPA4 gene belongs to the family of heat shock protein 70 was quantified. Heat
shock proteins (hsp) have been known to play an important role in cell protection and
repair upon stress as well as under certain non-stressful conditions302–304. More
specifically, hsp70 has been identified as a molecular chaperone which can inhibit
apoptosis. Mechanistically hsp70 inhibits apoptosis in cells by preventing the
constitution of the apoptosome and by protecting against the apoptogenic effects of
apoptosis inducing factors (AIF) targeted to the extramitochondrial compartment304,305.
The results showed that simulated microgravity-cultured cells up-regulated gene
expression of HSPA4 which belongs to the family of genes encoding hsp70 protein,
after 2 days of culture as shown in Figure 7-4. [After 2 days: 1g 0.496+/- 0.129, µg
1.0+/- 0.306, p<0.05]. However, after 4 days in culture, there was no statistical
difference in the level of HSPA4 between simulated microgravity and normal gravity.
[After 4 days: 1g 0.346 +/- 0.154; µg 0.726+/- 0.101, not significant]. The data generated
was from a representative population of HE-like cells isolated from a single donor.
Previous studies have shown the ability of mature EC populations to sense simulated
microgravity as a stressful condition and upregulate proteins such as hsp70 and
maintain their ability to proliferate133,263,301,302.
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Figure 7-4. Gene expression. Representative relative gene expression of HSPA4 by microgravity (µg) cultured HE-like cells vs cells maintained in normal gravity (1g) over 2 and 4 days of culture. Data shown are Mean +/- SD; n=3. N.S. indicates no statistically significant difference (p > 0.05) and * indicates a statistically significant increase (p<0.05).
Consistent with this trend, the results indicated that culture under simulated
microgravity conditions led to an initial upregulation of gene expression of HSPA4 over
2 days of culture but by day 4 the levels were similar between the normal and simulated
microgravity conditions. This data provides mechanistic insight into the ability of the HE-
like cells to survive and proliferate under “stressful” conditions such as simulated
microgravity by regulating stress proteins such as hsp70.
Another important molecule that regulates EC function is nitric oxide (NO). NO is
an important regulator of various important functions of the endothelium such as: cell
proliferation, angiogenesis, differentiation, and anti-thrombogenicity 299,306.Results
indicate that HE-like cells cultured under simulated microgravity for 2 days released
higher amounts of soluble NO when compared to the respective normal gravity controls.
[NO concentration (µM/pg of dsDNA): 1g 0.64+/- 0.308, µg 1.77+/- 0.202 p<0.05].
However, after 4 days of culture, the level of NO release was statistically similar
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between both groups. [NO concentration (µM/pg of dsDNA): 1g 2.49 +/- 0.050; µg 2.51
+/- 0.079 not significant] as shown in Figure 7-5 A.
Figure 7-5. Secretory functions of cells. Representative results of quantification of release of (A) Nitric oxide and (B) Interleukin-6 by microgravity (µg) cultured HE-like cells vs cells maintained in normal gravity (1g) over 2 and 4 days of culture. (Data shown are Mean +/- SD; n=3). N.S. indicates no statistically significant difference (p > 0.05) and * indicates a statistically significant increase (p<0.05).
Previous studies have reported a simulated microgravity-mediated increase in
endothelial nitric oxide synthase protein levels as well increased NO production at 48 h
and 72 h of simulated microgravity exposure respectively by mature ECs300. Also, a
study by Spisni et al. (2006) 307 reported no change in endothelial nitric oxide synthase
(eNOS) or inducible nitric oxide synthase (iNOS) protein levels yet an increase in NO
production at 24 h of simulated microgravity exposure. This increased level of NO may
also indicate another possible recovery mechanism of the cells to adapt to the changing
gravitational forces 285,301,302,308.Also more importantly, an increase in expression of NO,
a known vasodilator, leads to increased relaxation and lowering of blood pressure which
is known to disrupt sympathetic flow, leading to orthostatic intolerance306,308.Hence, this
may also provide a possible insight into cardiac dysfunction caused by space travel.
The relationship between endothelial dysfunction and inflammatory response
leading to cardiovascular complications such as atherosclerosis has been widely
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studied in the recent past246,299,309. Immune dysfunction in astronauts in space
environment has also been studied extensively288,310,311. A potential anti-inflammatory
response to short term culture under simulated microgravity by the down regulation of
pro-inflammatory cytokines has been reported by multiple studies133,300. These findings
have also been implicated to be mediated by changes in cytoskeletal arrangement in
simulated microgravity300. Assessment of the release of soluble IL-6, a major pro-
inflammatory cytokine, show that culture under simulated microgravity conditions for 2
and 4 days led to significant decrease in secretion of soluble IL-6 in the cell culture
medium when compared to normal gravity cultured controls. [IL-6 concentration (pg/ml)
after 2 days: 1g 46.3+/-1.0, µg 11.71+/-0.88 p<0.05; after 4 days: 1g 97.013 +/- 0.592;
µg 53.697 +/- 9.196 p<0.05] as shown in Figure 7-5 B.
Consistent with the literature reports for mature ECs, the results showed a
decrease in expression of IL-6, an important pro-inflammatory and anti-proliferative
cytokine 312. It is noteworthy that the inverse relationship between NO and IL-6 that was
seen in this study has previously been reported for ECs313,314. Hence the
downregulation of IL-6 may play an important role in the response of the HE-like cells to
simulated microgravity, since low amounts of IL-6 might contribute to cell growth as well
as to the enhancement of NO synthesis. Similarly the ability of hsp70 to inhibit the
synthesis of pro-inflammatory cytokines in ECs has also been reported315–317. This
altered immune function observed for the HE-like cells not only has implications in their
regulatory role regarding the overall immune response of the body during space travel,
but also additional implications on the effect of space travel on overall cardiovascular
health.
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The anti-thrombogenic behavior of ECs is important to maintain a healthy
vascular system and to prevent cardiovascular complications such as thrombosis269.
Healthy endothelium secretes various factors that are important in regulating
coagulation and fibrinolysis 243,299.Tissue plasminogen activator (tPA), an important
protein released by ECs, has been shown to be a main factor in the activation of the
fibrinolytic system, which is involved in dissolution of clots. Specifically, tPA activates
the conversion of clot bound plasminogen to plasmin which leads to clot lysis318,319. An
increase in expression of tPA has been directly correlated with increased bleeding while
a low expression of tPA has been implicated in thrombosis320. Quantification of the
release of soluble tissue plasminogen activator (tPA) into the surrounding cell culture
medium showed that tPA released by HE-like cells cultured under simulated
microgravity conditions for 2 and 4 days were significantly lower than their respective
normal gravity cultured controls [tPA concentration (ng/ml) after 2 days, 1g 4.572+/-
0.370, µg 2.01+/- 0.263 p<0.05; after 4 days: 1g 13.26 +/- 0.557; µg 9.867 +/- 0.528
p<0.05] as shown in Figure 7-5 C.
The data assessing the expression of tPA released by HE-like cells indicates
simulated microgravity mediated an increase in expression of tPA over 2 and 4 days of
culture. This indicates that exposure of HE-like cells to simulated microgravity can
potentially decrease the anti-thrombogenic potential of the cells. In order to more
directly study the anticoagulative functions of the HE-like cells, the ability of these cells
to delay clotting of human blood derived allogenic plasma was assessed using a
modified version of plasma clotting assay was performed. This assay provided a robust
and direct method to assess and compare the anti-thrombogenic potential of HE-like
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cells cultured under different gravity conditions. Once clotting was activated, progress to
full clot formation was assessed spectrophotometrically by continuous monitoring of the
change in absorbance over 30 min.
Figure 7-6. Plasma clotting kinetics. Recalcified plasma clotting kinetics for human plasma incubated with HE-like cells cultured in normal vs. simulated microgravity for (A) 2 days; and (B) 4 days. Data show absorbance plotted versus time. Arrows indicate the rightward shift in the normal gravity conditions, indicative of delayed clotting. SD error bars have been omitted for visual clarity. (Data shown are Mean +/- SD; n=3). (C) Representative result showing quantification of the time taken for recalcified plasma to clot (in minutes) when incubated with HE-like cells cultured for 2 and 4 day in normal vs simulated microgravity. Data shown are Mean +/- SD; n=6. N.S. indicates no statistically significant difference (p > 0.05) and * indicates a statistically significant increase (p<0.05).
The clotting profiles of recalcified plasma incubated with HE-like cells cultured in
both simulated microgravity as well as normal gravity for both 2 and 4 days is shown in
Figure 7-6. HE-like cells cultured in normal gravity exhibited a significant delay in the
onset of clotting as indicated by the rightward shift in the kinetic profile, indicating a
degree of anti-thrombogenic function. This contrasted with the kinetic profile of plasma
clotting when incubated with HE-like cells cultured under simulated microgravity, which
show an earlier onset to clot formation. These trends are seen at both 2 days and 4
days of culture as shown in Figure 7-6 A and B. As a control, CaCl2 was omitted from a
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subset of samples, and the plasma did not clot (data not shown). We also quantified
and reported the time taken to reach maximum absorbance indicating complete clotting
of plasma on the cell covered beads. The results indicate that recalcified plasma clots
faster when incubated with HE-like cells cultured under simulated microgravity as
compared to the longer time it takes plasma to clot when incubated with HE-like cells
cultured under normal gravity as shown in Figure 7-6 C [clotting time (min): with 2 day
cultured HE-like cells: 1g 14.31+/- 2.199, µg 10.001+/- 1.395 p<0.05; with 4 day
cultured HE-like cells: 1g 16.569 +/- 1.357; µg 12.456 +/- 1.457 p<0.05]. HUVEC coated
beads were utilized as the control population of mature ECs, and showed the most
delayed plasma clotting (19.3 +/- 1.9 min) (Data not shown). Similarly, naked beads
devoid of HE-like cells were also assessed, and they showed no delay in plasma clot
formation (data not shown). The clotting data corroborated the assessment of tPA
secretion, in that HE-like cells exposed to simulated microgravity were not able to delay
clot formation as well as the normal gravity cultured controls over 2 and 4 days of
culture. Also, both populations, normal and simulated microgravity show anti-
thrombogenic function, however, with exposure to simulated microgravity, this function
was hindered. This data collectively shows that exposure to simulated microgravity
negatively influences the anti-thrombogenic function of HE-like cells relative to those
cultured under normal gravity conditions. This inhibition of anti-thrombogenic
functionality of the cells potentially provides insight into thrombosis and other
coagulation processes in the body which have been known to directly cause
cardiovascular deconditioning. Although the ability of simulated microgravity to directly
affect platelet functions has been shown before,144 to the knowledge, the effect of
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simulated microgravity on direct anti-thrombogenic properties of EPC derived ECs, such
as inhibition of plasma clotting has not been reported thus far.
It is important to note that decreased production of inflammatory cytokines such
as IL-6 as well as increased production of NO have been known to synergistically inhibit
clotting processes in the body85,299,306,309. However, in this study, it was observed that
although there was a significant reduction in the release of IL-6 and an increase in
release of NO under simulated microgravity, the cells were able to induce faster clotting
of plasma. It is also interesting to note that although culture under simulated
microgravity conditions affects various secretory and proliferative mechanisms of the
HE-like cells, these functions seem to not influence the downstream ability of the cells to
delay clotting. This indicates the need to study extensively the effect of simulated
microgravity on various alternate and synergistic mechanisms involved in both
inflammatory as well as thrombogenic processes. Also, the relatively small sample size
(n=3, donor EPC populations) limits the detection of small statistical differences
between samples. Nevertheless, this report provides a useful insight into the modulation
of clotting mechanisms as well as secretory and proliferative ability of blood derived
ECs when exposed to reduced gravity conditions. Since these cells are in direct contact
with circulating blood, this study will prove to be particularly important when assessing
their role in cardiovascular deconditioning during space travel, as well as to develop
countermeasures to ameliorate or possibly reverse this effect.
Summary
The disruption of normal functioning of the endothelium has been shown to
contribute significantly in weightlessness-induced cardiovascular deconditioning during
space flight267. The focus of this study was to bridge the gap in knowledge on the effect
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of altered conditions such as microgravity on circulating stem cells as they differentiate
into endothelial cells. Circulating stem cells have been known to contribute significantly
to vascular repair mechanisms in the body277,279. The motivation for this work was to
study and understand the behavior of these cells in a low shear stress environment
such as simulated microgravity to assess their therapeutic potential for vascular repair
and regeneration.
In this study a NASA designed, rotating wall vessel (RWV) bioreactor that mimics
certain functional aspects of microgravity on earth (Synthecon, Houston, TX) was
utilized. The RWV was designed and developed at NASA’s Johnson Space Center to
simulate the effects of microgravity on cells in a ground-based culture system134,282. The
results obtained using simulated microgravity conditions utilizing the RWV have been
confirmed and validated by experimental models in space flights to outer space133,302.
The first part of the above described work involved the effect of simulated
microgravity on the differentiation of EPCs pre-exposed to simulated microgravity. This
report also studied the downstream function of the differentiated cells in terms of their
migration and proliferative capabilities, release of soluble factors such as PGE2 and
NO, and ability to regulate clotting mechanisms in the body. While both the migration
and proliferation of cells were inhibited by microgravity exposure, prolonged exposure of
EPCs to microgravity resulted in decrease in the migration abilities of EPC derived EC
and the proliferative capacity of the ECs seemed to recover on longer time of exposure.
The initial decrease and subsequent increase in release of soluble factors such as NO
and PGE2 reaffirmed the possibility of recovery mechanisms being involved on
prolonged exposure to weightlessness. Finally the assessment of whole blood and
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plasma clotting, indicated that microgravity pre exposure at the progenitor stage caused
inhibition, in the ability of PE-like cells to delay blood or plasma clotting while this effect
was again nulled with increase in time of exposure of EPCs to microgravity. This data
collectively provides some insight into the role of short term exposure to low shear
stress environment due to simulated microgravity, on cardiovascular deconditioning
during space travel. Especially, the effect of microgravity on anti-thrombogenic
properties of the endothelium has not been reported before. Encouraged by the results,
the study was further extended to include human blood progenitor derived endothelial
cells. The goal was to investigate and draw comparisons between the effect of
simulated microgravity on vascular regeneration mechanism of EPCs and ECs when
exposed prior and post differentiation respectively.
This was demonstrated in Chapter 6. The results from the study show that the
populations of human blood EPC derived ECs when exposed to simulated microgravity
proliferate faster, release higher amounts of soluble NO and lower amounts of pro-
inflammatory cytokines such as IL-6, when compared to their normal gravity cultured
controls. This can be possibly attributed to the ability of the cells to sense simulated
microgravity as a “stressful” condition and adapt by upregulating gene expression of
hsp70. However, these secretory as well as proliferative functions had no effect on the
anti-thrombogenic functions of these cells. Their anti-coagulative and anti-thrombogenic
ability as assessed by both upregulation of tPA as well their ability to delay plasma
clotting, were impaired upon exposure to simulated microgravity.
It is important to note here that even though simulated microgravity is considered
a low shear environment, further investigation have shown that a small amount of shear
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stress is experienced by particles in rotation in simulated microgravity bioreactors282,321.
This shear stress varies with particle size and the velocity of rotation. The average
shear stress on a microcarrier bead as used in the study based on mathematical
calculations and estimations has been shown to be < 0.5 dynes/cm2 as described in
detail in the background section. The shear stress is estimated to be even lower on cells
without a carrier being rotated in the field. Although this level of shear is tremendously
lower that arterial shear stress or even venous shear stress values, its effect still needs
to be considered. The endothelium is very sensitive to variations in the characteristics of
flow that in turn cause changes in shear stresses; which manifest as pathologic
changes in the artery wall.
While much work has been done to study the response of ECs as well as EPCs
under physiological and higher shear stress conditions, effects of low shear stress ( <1
dynes/ cm2) has not been well characterized. Recent work by Mazzolai et al. (2011),
characterized human late outgrowth EPCs under low shear stress ( ~ 0.3 dynes/ cm2)
and physiological shear stress values (~ 6 dynes/ cm2)322 . Their results showed that
under low shear the expression of tissue factor mRNA was significantly upregulated.
While the expression of tissue-type plasminogen activator (tPA), urokinase (u-PA) and
monocyte chemotactic protein (MCP1) were downregulated. These results implicated a
possible shift of these cells towards a pro-coagulant cell type. This shift was also shown
to be dependent on the direction of shear. These results are comparable to the current
study which showed reduced anti-thrombogenic potential. However, Mazzolai et al. also
noted no significant effect on inflammatory response via regulation of IL-6 which was in
contrary to what was observed in simulated microgravity culture where inflammatory
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cytokine IL-6 was downregulated significantly. This trend has however been seen with
mature ECs in simulated as well as real microgravity previously. Moreover, the current
study also shows consistent initial delay and then recovery of proliferative potential of
both EPCs and EPC derived ECs which is not observed with application of shear stress
alone. This could possibly be the differences in shear experienced by cells under
laminar flow in comparison with cells under rotation in a RWV bioreactor. Since
simulated microgravity produces the unique effect of mimicking weightlessness, the
study of cellular behavior under this condition is not directly comparable to shear stress
induced cell response.
Overall this study serves as the first step towards understanding the ability of
these repair cells to regulate clotting mechanisms and cardiovascular health in
response to macroscale cues provided to them by altered growth conditions. Further
studies to assess alternate and synergistic mechanisms involved in regulation of
cardiovascular health as well as inflammatory dysfunction under microgravity are
required to develop possible cell based countermeasures to prevent cardiovascular
deconditioning due to space travel. This approach also has the potential to be applied to
develop strategies for vascular regeneration on earth using EPCs.
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CHAPTER 8 SUMMARY, CONCLUSIONS AND FUTURE WORK
Summary of Results
Cardiovascular diseases are the leading cause of death worldwide, with more
than 60% occurring within United States itself1. Momentous improvements in clinical
care for patients suffering from complications due to cardiovascular diseases have been
made through the years. Significant efforts have been focused on perfecting
pharmaceutical and transplant based therapies to treat these diseases. However, the
chronic scarcity of donor organs fit for transplantation has led to increased efforts
towards developing novel tissue engineering based strategies to develop functional
replacements for diseased or damaged parts of the vasculature6,26,323.
Tissue engineering strategies are based on the premise that the functional cell is
the true ‘tissue engineer’17. Therefore, these strategies make use of ways to modulate
cellular response by providing appropriate cues to the cells. Given the complexity of the
vascular system itself, the cues presented must be multi-dimensional and at different
length scales of interaction. While many such cues have been studied historically, there
are many unknown stimuli whose effect is yet to be explored and exploited. With this
motivation in mind, the goal of this dissertation was to study the underlying mechanism
of specific cell-material interactions at different scales (nano, meso and macroscales),
to develop novel strategies to repair, preserve and regenerate vascular tissue.
Starting at the nanoscale, the first part of this work was focused on the use of a
supramolecular aptamer assembly as a receptor agonist. This subcellular interaction
was characterized by studying the activation of VEGFR2 receptor in endothelial cells by
the aptamer assembly. Activation of the VEGFR2 receptor has been shown to facilitate
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various processes such as proliferation, migration, neoangiogenesis and differentiation
in endothelial cells and various stem cell populations39,324. These positive outcomes of
the activation event have been exploited in designing pro-angiogenic therapies via
targeted activation of VEGFR2 for vascular tissue engineering to cause repair and
regeneration of vascular tissue37,325.
The results presented in this study show the design and evaluation of a divalent
aptamer assembly consisting of two monomeric units tethered together using a PEG
chain. This novel strategy of using nucleic aptamers based divalent assembly to target
VEGFR2 was effective in triggering the activation cascade in these receptors in
endothelial cells. The assembly was also able to cause positive downstream angiogenic
response which proves that this is a promising strategy to modulate cell behavior
favorably for vascular tissue engineering and regenerative medicine, where controlling
cell processes is of paramount importance. The ability to modulate a functional
response such as cell growth, differentiation and angiogenesis is critical for success in
these areas. Therefore, this novel, specific and targeted strategy has tremendous
potential towards designing novel nanoscale biomaterial cues for vascular tissue
engineering.
Moving up on the scale of interactions, the next part of this work consisted of
studying cues designed at the meso scale. The biological mesoscale range has been
described to be inclusive of biological structures ranging from 10 to several 100 microns
in size14,26. This includes interaction of cells with surrounding biomaterial scaffolds that
house the cells. Biomaterial environments provide a variety of physical, mechanical,
chemical as well as structural cues that can be altered for cardiovascular applications15.
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This section covered two classes of cardiovascular biomaterial applications, namely,
small diameter vascular grafts and vascular stents. This work was focused on studying
the interaction of these biomaterials with in vitro models of the vascular environment to
understand and develop strategies to repair damaged vasculature and enable
regeneration.
First part of the study, was focused on development of a novel bilayered,
construct for application as a vascular graft. The graft was designed to mimic the
biological and architectural aspects of native blood vessels. The design consisted of two
layers, a solid luminal layer made of elastomeric, biodegradable POC and a fibrous
medial layer consisting of a mixture of Collagen, POC and elastin. The rationale behind
the design is to provide a solid lumen for the formation of a monolayer of endothelial
cells and fibrous medial layer enriched with ECM proteins to enable formation of an
interconnected network of SMCs. The individual layers of the graft were separately
fabricated and assessed for cell attachment and proliferation. The constructs
mechanical properties were also assessed. Overall the results showed that the graft
was biocompatible and was able to support attachment and growth of vascular cells.
More interestingly, the mechanical properties evaluation revealed that the compliance of
the fabricated graft matched native artery values. This property is very important to
prevent implant failure due to mismatch in mechanical properties. This graft is currently
being developed to recruit circulating stem cells to promote integration into vasculature.
The second part of the study evaluated the hemocompatibility of an Mg-Ca-Sr
alloy to determine its utility as a candidate material for vascular applications. Overall,
this study showed that the alloy is biocompatible and invokes very low cytotoxic
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response when cultured with a mature vascular endothelial cell population. The alloy
sample was hemocompatible and exhibited excellent thromboresistance as indicated by
low hemolytic percentage as well as delayed and reduced clotting when in contact with
whole blood. The alloy also did not cause significant platelet adhesion and activation.
Finally, it also does not induce significant release of pro-inflammatory factors confirming
minimal inflammatory response. This data collectively confirms the suitability of the Mg-
Ca-Sr alloy for potential application as a vascular stent material.
The studies combined investigated the effect of biomaterial structural cues on
the meso scale on their ability to be used in vascular environments. The studies
involved evaluating some important characteristics of these materials in terms of their
biocompatibility, hemocompatibility, ability to support vascular cell types, mechanical
and biochemical behavior. These factors are essential in preventing vascular damage
due to complications such as thrombosis, toxicity as well as, fostering healthy
remodeling and regeneration of vasculature. Combined, these strategies widen the
knowledge in biomaterial-cell interaction and are preliminary steps towards developing
suitable biomaterials for vascular tissue engineering.
The final part of this dissertation was focused on macroscale systemic cues that
provide a unique environment for modulating response in vascular cells. The dynamic
nature of the cardiovascular system and the role of its constituent cell systems has been
studied extensively. The endothelium which lines the inner layer of blood vessels is
especially critical to its normal functioning269. It is this layer that actively and dynamically
controls repair and regeneration processes occurring within the vascular network. This
layer has been shown to respond to physiological mechanical cues, such as shear
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stress, cyclic stretch, and alterations to gravity which cause them to undergo significant
morphological and functional changes260,264.
One such specialized macro environmental model is microgravity in space flight,
which provides a unique opportunity to study the effect of gravitational changes on cells.
The negative effects of space travel on cardiovascular health has been well
documented in the past. A common feature of the effects of altered gravitational
conditions on different cellular systems is that the affected tissues rely on tissue-specific
stem cells for regeneration and repair. Understanding and replicating the behavior in
such a condition can lead to development of stem cell based therapy for cardiovascular
degeneration on earth. A recent study by Chiu et al., has shown that human cord blood
stem cells when cultured under simulated microgravity proliferated and trans-
differentiated into 3D structures resembling vascular tubules while exhibiting vascular
endothelial phenotype expressions140. Many such studies on the effect of stem cells
under alternate culture environments especially simulated microgravity is currently
being studied to develop regenerative strategies. Current research is also studying the
hypothesis that cardiovascular deconditioning due to space travel may potentially trigger
faster maturation of stem cells and therefore enhancing their regenerative potential. The
focus of this study was to bridge the gap in knowledge on the effect of altered conditions
such as microgravity on circulating stem cells as they differentiate into endothelial cells.
Circulating stem cells have been known to contribute significantly to vascular repair
mechanisms in the body270,326. The motivation for this work was to study and
understand the behavior of these cells in a low shear stress environment such as
176
simulated microgravity to assess their therapeutic potential for vascular repair and
regeneration.
The first part of the above described work involved the effect of simulated
microgravity on the differentiation of EPCs pre-exposed to simulated microgravity. This
report also studied the downstream function of the differentiated cells in terms of their
migration and proliferative capabilities, release of soluble factors such as PGE2 and
NO, and ability to regulate clotting mechanisms in the body. While both the migration
and proliferation of cells were inhibited by microgravity exposure, prolonged exposure of
EPCs to microgravity resulted in decrease in the migration abilities of EPC derived EC
and the proliferative capacity of the ECs seemed to recover on longer time of exposure.
The initial decrease and subsequent increase in release of soluble factors such as NO
and PGE2 reaffirmed the possibility of recovery mechanisms being involved on
prolonged exposure to weightlessness. Finally, the assessment of whole blood and
plasma clotting, indicated that microgravity pre exposure at the progenitor stage caused
inhibition, in the ability of PE-like cells to delay blood or plasma clotting while this effect
was again nulled with increase in time of exposure of EPCs to microgravity. This data
collectively provides some insight into the role of short term exposure to low shear
stress environment due to simulated microgravity, on cardiovascular deconditioning
during space travel. Especially, the effect of microgravity on anti-thrombogenic
properties of the endothelium has not been reported before. Encouraged by the results,
the study was further extended to include human blood progenitor derived endothelial
cells. The goal was to investigate and draw comparisons between the effect of
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simulated microgravity on vascular regeneration mechanism of EPCs and ECs when
exposed prior and post differentiation respectively.
This was demonstrated in Chapter 7. The results from the study show that the
populations of human blood EPC derived ECs when exposed to simulated microgravity
proliferate faster, release higher amounts of soluble NO and lower amounts of pro-
inflammatory cytokines such as IL-6, when compared to their normal gravity cultured
controls. This can be possibly attributed to the ability of the cells to sense simulated
microgravity as a “stressful” condition and adapt by upregulating gene expression of
hsp70. However, these secretory as well as proliferative functions had no effect on the
anti-thrombogenic functions of these cells. Their anti-coagulative and anti-thrombogenic
ability as assessed by both upregulation of tPA as well their ability to delay plasma
clotting, were impaired upon exposure to simulated microgravity.
Overall this study serves as the first step towards understanding the ability of
these repair cells to regulate clotting mechanisms and cardiovascular health in
response to macroscale cues provided to them by altered growth conditions. Further
studies to assess alternate and synergistic mechanisms involved in regulation of
cardiovascular health as well as inflammatory dysfunction under microgravity are
required to develop possible cell based countermeasures to prevent cardiovascular
deconditioning due to space travel. This approach also has the potential to be applied to
develop strategies for vascular regeneration on earth using EPCs.
Conclusions
As Biomaterial Engineers, the most important toolkit at our disposal is the
structure property relationship of materials. It is this characteristic of materials that
enables flexibility in design and development of novel biomaterials that can elicit
178
appropriate responses from cell systems. Given the complexity of living tissues,
techniques for tissue engineering require the use of novel external cues to aid the
process. The overall aim of this dissertation work was to study multiscale cues that can
be presented to cells and the response elicited by them and their relevance in a
vascular setting. The underlying theme of studying these interactions is to repair,
preserve and regenerate vascular tissue to alleviate conditions affecting the health of
the vascular network. The nano, meso and macro scale stimuli discussed in the above
chapters cause a unique response in a simulated vascular environment. Although this
work is not a comprehensive, all-encompassing study of all cell-material interaction
scenarios, these efforts contribute towards widening the knowledge base in this field.
This work also opens potential avenues to use these systems either separately or in
combination to develop a novel strategy for vascular tissue engineering. Therefore, this
work is a valuable addition to the field of cardiovascular tissue engineering and
regenerative therapy.
Future Work
Each chapter describes a specific and novel system of interactions at cell-
biomaterial interface. These studies represent preliminary work on characterizing the
initial response of cellular systems to cues at different scales. These results can be
further expanded to in vivo settings through animal studies and applied to organotypic
models. For example, the development of an agonist aptamer assembly for VEGFR2
receptors in endothelial cells as described in Chapter 3 opens up avenues to further
extend this work to other receptors and cell types for applications in tissue engineering,
stem cell engineering, cancer therapy etc.
179
Further the information learnt from these strategies can be combined to present
multiple cues at the same time. Such strategies have been used while designing
bioactive scaffolds for vascular tissue engineering in the past46,58,98,99. One possible
avenue in this direction would be the use of novel bioactive scaffold systems based on
the Collagen POC Elastin system decorated with the VEGFR2 aptamer assembly to
capture and recruit endothelial progenitor cells from circulating blood.
Therefore, future work based on the results presented in this dissertation will
potentially advance vascular tissue engineering and help further clinical care for these
complications.
180
APPENDIX PROTEIN CROSSLINKING USING CINNAMALDEHYDE VAPORS
Chapter 4 described the negative effects of glutaraldehyde crosslinking and its
potential toxicity on cells. The goal of this study is to investigate the use of
cinnamaldehyde (CA) as a relatively non-toxic cross linker to stabilize protein scaffolds.
Tissue engineering scaffolds made of proteins from animal and plant sources have
been used effectively for a variety of applications185. However, these scaffolds must be
crosslinked to stabilize the proteins before they can be used in in vitro applications.
Protein scaffolds, without proper crosslinking have inferior resistance to degradation
and mechanical strength. Therefore, to improve their physical properties, many
crosslinking strategies have been utilized such as chemical cross linkers, plasticizers as
well as ionizing radiations. Currently, the most robust and popular cross linking agents
employed are primarily aldehydes such as formaldehyde and glutaraldehyde which can
cause significant toxicity to cells unless appropriate precautions are taken186–188.
Therefore, there is a constant search for alternate natural cross-linking agents which do
not cause significant toxicity to tissues.
Cinnamaldehyde occurs naturally as an aromatic α,β-unsaturated aldehyde. It is
derived from cinnamon which is used popularly as a spice. It also has medicinal
properties and has been used as a home remedy for ailments for common cold and
digestive disorders. Cinnamon oil has also been used time and again as a food
preservative because of its excellent anti-microbial properties327. More recently it has
been identified as a crosslinking agent for protein films in the food packaging industry.
Another recent report also explores its use to cross link protein hydrogels for wound
healing applications327,328.
181
This preliminary study investigates the use of cinnamaldehyde vapors for cross
linking proteins. In this work, the crosslinking action of cinnamaldehyde liquid is
compared with that of cinnamaldehyde vapors. The use of the vapor form of
cinnamaldehyde for crosslinking applications has not been reported before.
Electrospinning Gelatin Nanofibers
12.5 wt. % solution of gelatin in Tetra fluoro ethylene was prepared for
electrospinning. The mixture was vortexed for 15 min and allowed to dissolve overnight
to thoroughly mix the constituents. The gelatin solution was fed through a 25G needle at
0.8 mL/ h with a syringe pump and a voltage potential of 22-25 kV was used; the
working distance was set at 25 cm.
Cinnamaldehyde Crosslinking and Preliminary Stabilization Assessment
The gelatin scaffolds in their as spun state were soluble in water. The scaffolds
were then cross linked using the conditions mentioned in the table below. Further the
stability of the scaffolds was preliminarily tested by assessing their ability to resist
dissolution in water.
Table A-1. CA cross-linking conditions and resistance to degradation
Conditions Water stability
Un-crosslinked control Dissolved in water
Liquid CA crosslinking for 24 h at RT Stable in water
Vapor CA crosslinking for 4 days at 45⁰ C Dissolved in water
Vapor CA crosslinking for 7days at 45⁰ C Stable in water
Thermal crosslinking control for 7 days at 45⁰ C Dissolved in water
From preliminary testing, it was shown that only two conditions stabilized the
gelatin fibers, viz, the fibers crosslinked with liquid CA overnight and the CA vapor
crosslinked fibers for 1 week.
182
FTIR Characterization
The crosslinking of scaffolds was confirmed by FTIR. Spectrochemical analysis
of the electrospun scaffolds crosslinked at various conditions was performed using
attenuated total reflectance Fourier transformation infrared spectroscopy (ATR-FTIR).
ATR-FTIR spectra was obtained for each of the samples using the Nicolet 6700 FTIR
Spectrometer (Thermo Scientific) and a diamond tip window. The spectra were read
over a range of 600-4000 cm-1 for each of the spectra and a total of 32 scans were used
with a resolution of 4 cm-1.
Figure A-1. FTIR spectra of cross linked fibers
FTIR analysis as shown above confirmed the crosslinking of scaffolds as
indicated by a characteristic Amide II (NH) bending vibrations peak observed at around
1550 cm-1. Additionally, Amide I (C=O stretching) and Amide A (N-H stretching)
signatures were also seen respectively at 1632- 1664 and 3320-3340 cm-1for all
scaffolds tested.
183
SEM Characterization
The stable samples were washed and frozen at -80˚C for 18 h before being
lyophilized. The dried scaffolds were then mounted for SEM analysis and the structure
of the scaffolds were observed using a table top SEM (PhenomWorld ProX, Eindhoven,
Netherlands).
Figure A-2. SEM characterization of CA crosslinking. (A) Un crosslinked gelatin fibers (B) Liquid CA crosslinked fibers after 24 h at room temperature (C) CA vapor
crosslinked fibers after 7 days at 45⁰C. All images are at a magnification of
5000 X.
The preliminary results from this study indicates CA vapors are effective in cross
linking protein fibers. FTIR spectra confirm this hypothesis. The SEM images
interestingly indicate that there is coalescing of fibers and the morphology changes on
cross linking. Further comprehensive evaluation of the crosslinked fibers as well as their
structure is required to further develop strategy to create functional scaffolds.
A B C
184
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BIOGRAPHICAL SKETCH
Vidhya was born in the beautiful metropolitan city of Chennai, home to the
world’s longest natural urban beach. She spent the major part of her life growing up in
this culturally rich city with her family. She graduated high school from Padma Seshadri
Bala Bhavan and then moved to Tiruchirapalli in south India for her bachelor’s degree.
Her alma mater was the National Institute of Technology where she earned a Bachelor
of Technology degree in Metallurgical and Materials Engineering in 2013. She then
moved to Florida to attend graduate school at University of Florida in fall 2013. She
started as a master’s student and transitioned to the doctoral program under the
guidance of Dr. Josephine Allen in fall 2014. She received her Master of Science
degree in 2016 and her doctoral degree in 2017, in materials Science and engineering
with a concentration in biomaterials. She has also served as the secretary for Society of
Biomaterials UF chapter during 2016-2017 and as an executive member of the student
safety council at the Department of Materials Science and Engineering.
