PEGDA Pouches to Prevent Adhesions with in vivo Bioreactor- based Vascular

Graft Strategies

By

Kranthi Vuppuluri Bachelor of Technology in Biotechnology GITAM University, India, May 2014

A thesis submitted to of the College of Engineering at Florida Institute of Technology in Partial Fulfillment of the Requirements for the Degree of

Master of Science in Biomedical Engineering

Florida Institute of Technology Melbourne, Florida May, 2017

© Copyright 2017 Kranthi Vuppuluri All rights reserved

The author grants permission to make single copies ______

We the undersigned committee hereby approve the attached thesis “PEGDA Pouches to Prevent Adhesions with in vivo Bioreactor- based Vascular Graft Strategies” By Kranthi Vuppuluri

C. Bashur, Ph.D. L.K. Moore, Ph.D. Assistant Professor, Research Professor, Biomedical Engineering Biological Sciences Committee Chairperson Committee Member

V. Kishore, Ph.D. Alessandra Carriero, Ph.D. Assistant Professor, Assistant Professor, Chemical Engineering, Biomedical Engineering Committee Member Committee Member

Ted Conway, Ph.D. Department Head of Biomedical Engineering

Abstract TITLE

PEGDA Pouches to Prevent Adhesions with in vivo Bioreactor- based Vascular

Graft Strategies

Author Kranthi Vuppuluri

Principle Advisor

Christopher A. Bashur, PhD

Different strategies have been investigated for the fabrication of vascular grafts.

Using the peritoneal cavity of the patient as an “in vivo bioreactor” to recruit autologous cells to the implanted vascular conduit is one of the promising options.

One of the main drawbacks with this strategy is the potential to form adhesions in the peritoneal cavity. In this project we are trying to address this potential side- effect by using polyethylene (glycol) diacrylate (PEGDA) to produce a hydrogel pouch into which the electrospun conduits are placed, with the goal of reducing the potential to form peritoneal adhesion after implantation. PEG is hydrophilic

iii material that shows low cell attachment and resistance to protein adhesion making it a suitable material. PEG pouches will be fabricated by crosslinking PEGDA with

Irgacure 2959. A pilot study has been performed to determine the response in a rat model. Mechanical properties of these pouches will be studied at different concentration and loading rates to find the best concentration of PEGDA that gives desired mechanical properties and strength to prevent the pouch from breaking when implanted inside the peritoneal cavity.

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Contents

Abstract ...... iii List of Figures ...... vii LIST OF TABLES ...... x Attribution ...... xi Acknowledgement ...... xii CHAPTER 1 INTRODUCTION ...... 1 1.1 Introduction ...... 1 1.2 Cardiovascular diseases ...... 2 1.2.1 Percutaneous coronary intervention (PCI) ...... 3 1.2.2 Coronary artery bypass grafting ...... 3 1.3 Tissue engineered vascular grafts ...... 4 1.4 Bioreactor...... 6 1.4.1 Bioreactors in tissue engineering...... 6 1.4.2 Types of bioreactors ...... 7 1.4.3 Peritoneal Cavity as a Bioreactor ...... 9 1.5 Host response to implanted biomaterials ...... 10 1.5.1Beginning of the inflammatory response ...... 11 1.5.2 Macrophages ...... 13 1.5.3 Inflammatory response in peritoneal cavity ...... 15 1.5.3.1 Formation of peritoneal adhesions ...... 15 ...... 17 1.5.4 Prevention of peritoneal adhesions ...... 17 1.5.4.1 General principles during surgery ...... 18 1.5.4.2 Mechanical barriers ...... 18

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1.5.4.3 Chemical agents ...... 19 1.6 Hydrogels ...... 20 1.6.1 Hydrogels in tissue engineering ...... 20 1.6.2 Biocompatibility ...... 21 1.6.3 PEGDA ...... 22 1.6.4 Mechanics of hydrogels ...... 23 1.7 Strategy ...... 23 CHAPTER 2 MATERIALS AND METHODS ...... 25 2.1 Materials ...... 25 2.2 Production of conduits ...... 25 2.3 Producing PEG pouches ...... 26 2.4 Implantation of PEGDA pouches in peritoneal cavity...... 28 2.5 Cryosection ...... 28 2.6 Hematoxylin and Eosin staining ...... 28 2.7 Immunofluorescent Imaging ...... 29 2.8 Mechanical testing ...... 30 2.9 Statistics ...... 31 CHAPTER 3 RESULTS AND DISCUSSION ...... 32 3.1 Hydrogel characterization ...... 32 3.2 Pilot Implantation Study ...... 33 3.3 H&E Staining ...... 35 3.4 Peritoneal Fluid ...... 37 ...... 39 3.3 Mechanical Testing ...... 39 CHAPTER 4 CONCLUSION AND FUTURE WORKS ...... 44 4.1 Conclusion ...... 44 4.2 Future Work ...... 45 REFERENCE ...... 47

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List of Figures Figure 1. This graph shows a classical response to biomaterials. The timeline shows the inflammation stages. Permitted by Robbins Basic Pathology 9th edition.

...... 12

Figure 2. Shows adhesions in peritoneal cavity. Fig (A) shows minor adhesions.

Fig (B) shows severe adhesions. Permitted by [119] ...... 17

Figure 3. Shows the schematic of production of PEG pouches. The pouch is produced using PDMS molds and conduits are placed in it. The pouch is then enclosed using more PEG ...... 27

Figure 4. Shown is a PEG pouch enclosed with electrospun conduits and silicone tubes placed for support ...... 32

Figure 5. Shown are two different pouches implanted in the peritoneal cavity for 4 weeks. (A) shows PTFE pouch with dense adhesions formed. (B) Shows the limited adhesions caused due to PEG pouch. The arrow marks show the adhesions that are formed. The adhesions are noticeable where the pores are present...... 33

Figure 6. Shows PEG hydrogel in peritoneal cavity post implantation. (A)

Shows a strip of PEG hydrogel with negligible adhesions due to absence of pores.

(B) Shows fractured part of hydrogel believed to be caused due to forces inside the body...... 34

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Figure 7. Shows H&E staining of sections of conduits from 2 different conditions. (A) Shows conduit from 3x2 pouch. (B) Shows conduit from 4x5 pouch.

Both the conduits show recruitment of cells ...... 35

Figure 8. Shows cross sections of 2 different pouches. (A) Shows cross-section of PTFE pouch. The tissue is evident and the cells are well spread along with the presence of connective tissue. The arrow mark indicates the pouch material (PTFE)

(B) shows the cross section of PEG pouch. The cells are not well spread and there is negligible connective tissue present...... 36

Figure 9. Shows cells from peritoneal fluid after pouch implantation. (A) Shows peritoneal fluid from the condition 3x2 pores. (B) Shows peritoneal fluid from the condition 4x5. Many cells appear to be round...... 37

Figure 10. Shows representative images for contractile markers calponin and α smooth muscle actin and M1 macrophage marker CD80 for the conditions 3x2 and

4x5 number of pores in the pouches. The nuclei are stained with DAPI and shown in blue...... 38

Figure 11. Shows the representative images for the cell proliferation marker, thrombospondin and endothelial cell marker vWF. Nuclei are stained with DAPI and are shown in blue ...... 39

Figure 12. Representative stress strain curve of the compression of PEG hydrogels at different concentrations. The compression was done at a constant rate of 25 min-1. The arrows indicate the peak where the first fracture was formed...... 40

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Figure 13. Shows percent compression of different concentration of PEGDA at a constant compression rate of 25min-1 ...... 41

Figure 14 Shows ultimate compression stress for different concentrations of

PEGDA for n=3 for each group ...... 42

Figure 15. Shows tangential moduli of different concentration of PEGDA. The data had significant differences represented by * from 10%, # from 15% and @ from

20% PEGDA (all p values < 0.0165) ...... 43

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LIST OF TABLES Table 1 Grades of adhesions ...... 16

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Attribution

Several committee members and colleagues aided in the research and writing of this thesis. A brief description of their background and their contributions are included below.

Chris A. Bashur - Ph.D. (Department of Biomedical Engineering/ Florida

Institute of Technology) is the primary Advisor and Committee Chair. He provided extensive guidance for the research and guided me with writing of this document.

Mozhgan Shojaee – M.S (Department of Biomedical Engineering/ Florida

Institute of Technology). A significant amount of work has been done with her help.

She collaborated equally in this study. She assisted in implantation surgeries, and helped prepare the electrospun conduits used in implantation. She also helped in the analysis of the data.

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Acknowledgement

Firstly, I would like to sincerely thank my adviser Dr. Bashur. It has been a great pleasure to work under his guidance and support. He has encouraged me and dedicated his valuable time in teaching us technical evaluation and writing.

To all my fellow lab mates and buddies at Florida Tech: Aditya, Susan,

Tabby, Mozhgan, Likhitha, Kenyatta, Mahyar, Andrea. Thank you so much guys for keeping me happy all the time and providing me our support all throughout my master’s journey. You guys made this experience a memorable one

A special thanks to my parents and Divya for the unconditional love and support you have given me. There was never a dull moment.

Last but not the least thank you anna and Prithvi, it would not have been possible without you guys. You have been a constant support and strength in my life. Love you and thank you for everything.

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Dedicated to my grandparents

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CHAPTER 1 INTRODUCTION

1.1 Introduction

This research study focuses on producing a functional pouch which helps in aiding the peritoneal cavity to act as an in vivo bioreactor. We aim at developing tissue engineered vascular conduits to produce a graft for small diameter arteries with the proliferation of autologous cells, which also possess biodegradable properties [1].

The autologous cells infiltrate into the conduits subsequently after implanting them in the peritoneal cavity [2]. The peritoneal cavity acts as a bioreactor to the electrospun conduits by providing the required environment in terms of nutrients, temperature and growth factors [2]. The PTFE pouches used in our preceding studies served the purpose of enclosing the conduits [3,4]. However, adhesions are developed in the peritoneal cavity as an inflammatory response [5] and as a result, deposition of fatty tissue and the adhesions proved to be a hindrance to further procedures making it difficult to find the pouch. The adhesions can be classified based on their scoring [6] and analysis of the intensity of the situation. The hydrophilic properties of Polyethylene Glycol (PEG) [7] makes it a promising alternative to address the problems faced due to attachment of proteins to the biomaterial. These challenges include the peritoneal adhesion caused in the peritoneal cavity. It also aids in the peritoneal preimplantation strategies by acting as a carrier vessel in the peritoneal cavity. The overall goal of this project is to produce

1 a PEG pouch which acts a carrier to the enclosed electrospun conduits when implanted in the peritoneal cavity. This helps the tissue engineered vascular graft

(TEVG) in recruiting autologous cells to it prior to aortic implantation.

1.2 Cardiovascular diseases

Cardiovascular disease is one of the leading causes of death in the USA and around the world. It has been reported that cardiovascular disease is the cause of death of 1 in 3 people [8]. Coronary artery disease is one of the cardiovascular diseases caused due to atherosclerosis of the coronary arteries. The atherosclerotic plaques that build up along the walls of the coronary arteries constrict the diameter of the arteries resulting in interruption of the oxygen- rich blood flow, thrombus formation, or completely blocking the blood flow i.e., occlusion. It is estimated that the number of deaths due to cardiovascular disease is increasing steadily. So, treatment of cardiovascular diseases remains as a priority of concern [9,10] .

A number of attempts have been made to address the current issue, depending on the different conditions of the patients, for example, the extent of blockage at the lesion site. These include percutaneous coronary intervention using stents, carotid endarterectomy and coronary artery bypass grafting.

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1.2.1 Percutaneous coronary intervention (PCI)

Percutaneous coronary intervention, also known as angioplasty is a minimally invasive procedure that uses a catheter and a balloon along with a stent made of wire mesh [11,12]. The balloon catheter is directed towards the clot and the balloon is inflated. This helps in driving the plaque towards the walls of the vessel. The stent is introduced at the site to ensure that the flattened plaque remains in the same position and does not impede the blood flow. The advantage of this procedure is that it is fast and being a minimally invasive surgical procedure, it does not leave a scar behind.

The limitation of this procedure is, it cannot be performed if multiple blood vessels are to be treated [13]. There is also a risk of recurrence of plaque and that stent thrombosis might also occur [14].

1.2.2 Coronary artery bypass grafting

Though PCI is a minimally invasive surgical procedure and can provide a speedy recovery, coronary artery bypassing has been the gold standard [15,16]. The coronary artery that has been occluded can be bypassed by using natural or synthetic grafts.

The natural grafts can be obtained from the autologous veins, or mammary artery, among which the saphenous vein mostly is used [17,18]. These veins can be used to bypass the flow by suturing the vein on both the sides of occlusion so that the blood does not have to pass through it and the flow is maintained. However, the option of saphenous vein is not available for more than 33% of the patients [18]. In these cases

3 synthetic grafts like Dacron or ePTFE have been used [22]. Dacron is woven from polyethylene terephthalate and ePTFE is expanded polytetrafluoroethylene. They are successful in bypassing the large diameter arteries [19]. The graft patency is compromised when these synthetic graft are used for small diameter arteries. The large diameter arteries are associated with high velocity of blood flow whereas the small diameter which have reduced blood flow velocity cause vessel thrombosis.

Thrombosis occurs either due to the inherent thrombogenicity of the material or intrusion of the intimal hyperplasia [21] due to continued migration and proliferation of smooth muscle cells into the lumen where the artery and the synthetic material come into contact [22]. Thrombosis might occur due to the lowered velocity in the small diameter arteries. The tissue engineered, self- repairing vascular grafts might be an alternative option to overcome some of the limitations with the other options

[23].

1.3 Tissue engineered vascular grafts

Tissue engineering can be defined as “application of principles of engineering and life sciences to develop biological substitutes that restore, maintain or improve tissue function” [24]. Synthetic or natural materials can be used to fabricate small diameter grafts on which cells can be either seeded or recruited to the surface [25]. This method can be very promising as it bypasses the problems like limited supply of

4 veins, donor site morbidity and has a potential to raise complications due to thrombosis.

For the production of any tissue engineered product it is important to understand the native tissue. The coronary artery consists of three layers namely tunica intima, tunica media and tunica adventitia. Tunica intima is the inner most layer which consists of endothelial cells. It is the layer which is in direct contact with the blood flow. The endothelial cells which are oriented in the direction of the blood flow prevent the escape of blood cells and the blood proteins to the outer layers by providing a charged layer [26]. It also promotes anti-thrombogenicity by inhibition of platelet adhesion [27].

The middle layer, tunica media primarily is comprised of vascular smooth muscle cells (V-SMC’s) [28] and elastic fibers [29]. The V-SMC’s present produce collagen and elastic fibers help in retaining the structural stability and strength of the vessel [30]. The outer layer, tunica adventitia is comprised mostly of collagen and glycosaminoglycan produced by the fibroblasts [31]. It helps in preserving the structural integrity of the vessel during high arterial pressure [32] and anchors the blood vessel to the connecting tissue.

The structure of the arterial vessels described above provide the functional characteristics and the structural stability required for withstanding the various forces exerted on the arterial walls. Tissue engineered strategies can be used to produce a

5 vascular graft with properties mimicking the native tissue to provide long-term graft patency. As the mechanical response and cellular response run hand in hand, the mechanical cues are supposed to be maintained. The cell types observed mostly in vascular tissue engineering are endothelial cells and V-SMC’s. The strategy used in this research is to produce a graft with an autologous cell seeded scaffold with biodegradable properties. This is done by using an in vivo bioreactor to recruit the autologous cells to the graft prior to grafting in the blood vessel.

1.4 Bioreactor

Bioreactors are devices that are used to influence a biological process by providing controlled conditions [33]. Traditionally, bioreactors are used in various industrial applications like fermentation, waste water treatment, production of pharmaceuticals, recombinant proteins like antibodies, antibiotics and vaccines and

[34]. Bioreactors in tissue engineering are designed typically to provide the scaffolds with the similar environment of the native tissue in terms of nutrition and the mechanical stimulation which leads to cellular activity and differentiation [35].

1.4.1 Bioreactors in tissue engineering

Bioreactors have been used in tissue engineering for various applications like cell proliferation in both small and large scale, fabrication of 3D tissues in vitro and direct organ support devices [36]. The bioreactors help in the development of tissue by providing the required biochemical and physiological signals to the cells to

6 differentiate and allows for tissue production and maturation [37] before using it as a graft.

1.4.2 Types of bioreactors

The bioreactors that are used for tissue engineering must be designed differently for various purposes and take into consideration the 3D characterization, complexity of the tissue and physiological environment required [38,39]. In other words the bioreactor must be able to provide required physical and mechanical cues, nutrients and oxygen in a sterile manner. A few in vitro bioreactors include spinner flask, rotating wall, strain, hydrostatic and flow perfusion bioreactors. Spinner flask bioreactors are one of the most basic bioreactors in which a spinner generates the mixing of nutrients and oxygen throughout the media thereby reducing the boundary layer at the surface [40]. The scaffolds are suspended in the media using needles and the media is stirred using a magnetic stirrer [40]. Rotating wall bioreactors consists of cylindrical chamber in which both the exterior and interior wall rotate at a constant angular speed. The scaffolds are in a vessel where they can move freely. The vessel is rotated at a speed where a balance is acquired between the gravitational force and the upward dynamic force [41]. The scaffolds are thus suspended in the media due to the opposing forces acting on them simultaneously [40].

Strain bioreactors have been used in tissue generation including ligament, cardiovascular, bone and cartilage tissues [41]. The scaffolds are clamped to apply

7 the tensile force. The scaffolds can be loaded and uniaxial tension is applied. Further, the clamps can be adjusted by stretching the distance between them. Hydrostatic pressure bioreactors can be used to apply mechanical stimulation to the cell seeded scaffolds [41]. The scaffolds are cultured earlier and moved into the hydrostatic chamber. The chamber consists of a piston and an actuator where the pressure is applied using an impermeable membrane [42].

Flow perfusion bioreactors are used to provide more homogenous cell distribution through the scaffolds [43]. They consist of a pump and a scaffold chamber which are joined together using tubing [41]. The pump is used to force the media through the scaffolds which are present in the chamber placed across the flow path. Media perfuses through the scaffolds thus improving the fluid transport [44].

The bioreactors play a major role in production of vascular grafts by providing the scaffolds with required mechanical stimulation and the environment for the growth of the cells. The bioreactors that are popular to produce TEVG’s are strain, hydrostatic pressure and perfusion bioreactors. The body can be viewed as an in vivo bioreactor to recruit autologous cells hence avoiding the tissue rejection. The usage of the peritoneal cavity as in vivo bioreactor was started by Campbell et.al for the production of vascular graft [2]. This usage of the peritoneal cavity as an in vivo bioreactor has advantages like recruitment of autologous cells and reduction in thrombosis.

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1.4.3 Peritoneal Cavity as a Bioreactor

The peritoneal cavity is a space between the parietal and visceral peritoneum that has been used as a bioreactor. It is comprised of a minor amount of serous fluid. This fluid helps in the lubrication of the visceral peritoneum along the walls of the peritoneum [45]. It is comprised of water, proteins, electrolytes and other solutes derived from interstitial fluid in the adjacent tissues and from the plasma. The peritoneal cavity is the largest cavity in the whole human body [45]. It is composed of a thick layer of mesothelium and is supported by a thinner layer of connective tissue. The mesothelial cells being very fragile and loosely attached to the surface can be easily detached even during a small trauma.

The peritoneal cavity has been used as an in vivo bioreactor which can provide nutrients, cells and growth factors in a temperature controlled and sterile environment [46]. When a foreign object is placed in the peritoneal cavity, as a part of the inflammatory response a capsule forms around it [47]. The capsule consists of myofibroblasts which are derived from the bone marrow and it is covered by a continuous layer of mesothelial cells [45,48], which are derived from the outer lining of the [49]peritoneal cavity and have similar antithrombotic characteristics to endothelial cells [48].

The conduits can be implanted in the peritoneal cavity to recruit cells to the conduits prior to aortic implantation[48]. These grafts are attached with nonthrombogenic

9 mesothelial cells as well as they avoid an adaptive response being generated by the same patient. The phenotype can be maintained as the cells can be directly recruited to the graft; whereas it might pose a problem if the cells are cultured using an in vitro like expansion of the primary cells, which then may not mimic the inflammatory response. The peritoneal cavity can be used to circumvent these difficulties to produce a functional graft. The biomaterial that is used to produce a vascular graft must be biocompatible, provide the mechanical strength and support cell infiltration.

1.5 Host response to implanted biomaterials

A surgical procedure is always accompanied by an injury and disruption of the tissue at the surgical site. Details of the pathways of inflammatory response have been determined, but there are still important details that continue to be discovered. The response usually follows a set of programmed events which result in the formation of a dense fibrous connective tissue at the site of injury [49–52]. Regeneration of the tissue is carried out by the bone marrow, liver, epithelium and the epidermis of the skin. The host reactions can include blood-material interaction, matrix formation, acute inflammation, chronic inflammation, granulation tissue formation and fibrous tissue formation. With proper modulation of the inflammatory response, some of the later steps that lead to graft rejection can be avoided.

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1.5.1Beginning of the inflammatory response

Neutrophils are responsible for the acute inflammatory response. The degranulation process produces interleukin-4 (IL-4) and interleukin 13 (IL-13) which determine the extent of inflammation [53] (Fig. 1). The proteins from the blood and the interstitial fluid are adsorbed by the biomaterial surface which plays an important role in determining the extent of coagulation. Thrombin, which is released by factor XII, is responsible for the formation of fibrin from fibrinogen, forming a primary fibrous mesh around the biomaterial [54,55]. The complement system is activated following the contact with a biomaterial. Absorbed proteins such as albumin, fibrinogen, complement, fibronectin, fibronectin, γ globulin on the surface extra cellular matrix can adsorb onto the surface of the biomaterial [56–58].

The next step is chronic inflammation, which is associated with the presence of activated macrophages. The process of macrophage accumulation takes place in one week depending upon the biomaterial implanted and the type of proteins adsorbed

[59]. After the meshwork of ECM is deposited around the material, the presence of macrophages and the formation of granulation tissue follows which leads to the formation of dense layer of collagenous connective tissue layer [55].

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Figure 1. This graph shows a classical response to biomaterials. The timeline shows the inflammation stages. Permitted by Robbins Basic Pathology 9th edition. For the progression of these events successfully, there must be a migration of macrophages to the site of implantation. The chemokines are responsible for the movement of the monocytes. Integrins serve as surface receptors that mediate the intercellular interactions [60]. Chronic inflammation progresses into a granulation tissue phase where a new ECM is formed and growth of vasculature is evident. The foreign body reaction involves macrophages and often, multinucleated foreign body giant cells which are formed through the fusion of macrophages. A number of factors like composition of the material surface topography contribute to the potential foreign body giant cell formation [61]. Depending on type of signal stimulated and

12 the immunological environment, macrophages may undergo fusion and form foreign body giant cells. However, recent studies have shown that chronic inflammation does not always lead to the formation of foreign body giant cells [62].

1.5.2 Macrophages

It is assumed that the phenotype of the macrophage population stimulated in the host response that follows a grafting surgery may be dependent on the properties of a biomaterial [63]. Macrophages are derived from the myeloid progenitor cells through a cascade which includes such as cytokines granulocyte-macrophage colony stimulating factor (GM-CSF), granulocyte colony stimulating factor (G-CSF), [64] and macrophage stimulating factor (M-CSF). These signals are responsible for the differentiation of myeloid progenitor cells into monoblasts, pro-monocytes and monocytes [65,66]. Mature monocytes enter the blood stream a few days prior to entering tissues to become macrophages as a part of the inflammatory response.

Preceding the entrance of tissues and differentiation into macrophages, they are of two types namely, inflammatory and resident respectively. While the macrophages are characterized by their ability to migrate to the site of injury to proliferate with chronic inflammatory diseases, the resident macrophages are responsible to guard the tissue, populate normal tissues and act as regulators of the inflammatory response.

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Macrophages are similar to monocytes in possessing different markers and functions that depend on signals that are generated. Upon leaving the blood stream, macrophages are activated increasing the production of cytokines, chemokines, and other inflammatory molecules [67–69]. Polarized macrophages are typically either

M1 or M2 [70,71]. M1 are the pro-inflammatory macrophages that are induced by interferon- γ (IFN-γ) alone or in combination with either LPS or tumor necrosis factor

(TNF) [72–75]. The M1 classical response includes Th1 type responses, secretion of

TNF, and killing of the intracellular pathogens [76–78]. M2 are alternatively activated which can be further be divided into M2a, 2b, M2c subpopulations namely alternative, type II, and deactivated respectively. The Th2 responses include allergy, immunoregulation, killing and encapsulation of parasites, matrix deposition and remodeling and tumor promotion [79–83]. It has been shown that th3e macrophages can switch their phenotype from M1 to M2 and M2 to M1. The M1 type is an inflammatory response to destroy the pathogens and remove the dead cells from wound site while the M2 is associated with tissue remodeling and wound healing.

M1 polarization might result in scar formation. So, a timely switch between the polarization from M1 to M2 is a key component to result in positive wound healing response [62]. A TEVG should be designed to avoid this innate immune response, which would lead to graft failure.

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1.5.3 Inflammatory response in peritoneal cavity

Peritoneal adhesions are a side effect of peritoneal TEVG strategy. Peritoneal adhesions are thick bands of tissue that are formed as a result of inflammatory response in the peritoneal cavity which typically occur between 2-3 weeks of abdominal surgery [5,84]. Adhesions are the scar tissue that is formed during the healing process which loops itself along the lining of peritoneal cavity as well as around the organs present in the cavity like intestines, liver, spleen, gall bladder and fallopian tubes. Usually these organs have free movement but adhesions can cause immobility and may cause an obstruction of normal function [85].

Peritoneal adhesions are unavoidable in approximately all the abdominal surgeries.

In a study it has been reported that 90% of the surgeries resulted in peritoneal adhesions [86]. Peritoneal adhesions are one of the significant consequences faced with abdominal surgeries. Complications of more severe peritoneal adhesions include small and large bowl obstruction, nausea, vomiting, chronic abdominal pain and sterility in females and also make it more complicated for later surgeries as the organs adhere to each other or twist together [87].

1.5.3.1 Formation of peritoneal adhesions

Any kind of surgery or implantation of a biomaterial causes a foreign body reaction and leads to cellular reactions of wound healing to restore homeostasis. These reactions usually take place 2-3 weeks post-surgery or implantation [86]. In the

15 peritoneal cavity the injury causes the activation of mesothelium cells along the lining of peritoneal cavity. This causes the release of inflammatory cytokines tumor necrosis factor-α, interlukin-6 and successive recruitment of neutrophils, macrophages, eosinophils and fibrous discharge into the peritoneum [88]. These reactions lead to the formation of fibrin- based adhesions. Peritoneal regeneration occurs if the adhesions undergo fibrinolysis thus degrading fibrinous proto-adhesions by proteases like tissue plasminogen activator (tPA) [89]. However, in the presence of ischemia or sepsis, the proteolytic pathways are broken down and the fibrinous tissue develops by the up-regulation of chemokines such as transforming growth factor β, invasion of fibroblasts, macrophages and deposition of collagen to the site of inflammation. The fibroblasts secrete ECM at the site of inflammation as they proliferate thus resulting in the formation of strong fibrous adhesions [90,91].

The adhesions can be graded. They are graded from a scale of 0-4 as shown in Table

1. The grading of adhesions is done according to the criteria of Mazuji classification

[92].

Table 1 Showing grades of adhesions from 0-4 [93]. Grade Adhesions 0 No adhesion 1 Very small, irregular adhesions 2 Easily separable, medium intensity adhesions 3 Intense, not easily separable regular adhesions 4 Very intense, not easily separable, homogenous adhesions

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A B

Figure 2. Shows adhesions in peritoneal cavity. Fig (A) shows minor adhesions. Fig (B) shows severe adhesions. Permitted by [119]

1.5.4 Prevention of peritoneal adhesions

The prevention of peritoneal adhesions is important to reduce the complications due to the formation of adhesions. Adhesion formation can be reduced by a few steps like activation of fibrinolysis, obstructing coagulation, weakening the inflammation, prevention of collagen deposition or by creating a barrier between the wounded surface and the proximate surroundings. These can be summarized under the concept of general precautions during surgery, usage of mechanical barriers and introduction of chemical agents.

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1.5.4.1 General principles during surgery

Procedures like laparoscopy can be followed that require only minimal contact the peritoneal cavity. The advantages of this method include reduced incision of the parietal peritoneal cavity. This method includes reduced usage of foreign object like gauze, hair lint and sutures. Having a humidified environment prevents tissue from drying out and leaves less trauma at the site of injury. It involves less manipulation of the post-operative structure. However, this method by itself does not prevent adhesions and adhesion reformation has been reported in many cases [94].

1.5.4.2 Mechanical barriers

Liquid and solid barriers can be introduced to separate the layers in the peritoneal cavity during the period of peritoneal regeneration and prevent post-surgical adhesions. Gynecare intergel® is an FDA approved product which is a sterile solution of sodium hyaluronate crosslinked with ferric ions and has been suggested in patients undergoing abdominal surgeries for the prevention of post-surgical adhesions.

However, it has been reported that there was late outbreak of pain and tissue adherence in the peritoneal cavity [90]. Seprafilm® is one of the other products which is composed of hyaluronic acid with carboxymethyl cellulose which in turn develops into a hydrophilic gel after 24hrs of placing it in the body [95].

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1.5.4.3 Chemical agents

The balance between the fibrin production and peritoneal repair by subsequent fibrinolysis is the breaking point of adhesion formation. The fibrinolysis leads to peritoneal regeneration. Chemical agents can be used to propagate fibrinolysis thus inhibiting the proliferation of fibroblasts. Many chemical agents like corticosteroids, antibiotics, antioxidants, fibrinolytic agents and anticoagulants can be used to inhibit the fibroblast proliferation. Actilyse®, Arixtra®, Xigris® are three recombinant tissue- type plasminogen activator (tPA) which affect the coagulation process to prevent adhesions [96,97]. Though all the three products have been successfully tested there have been reports where the peritoneal adhesions were not completely eliminated

[96].

The above mentioned methods reduce the complications due to the adhesions forming in the peritoneal cavity. However, the above described procedures cannot not be used to enclose electrospun conduits to aid the peritoneal cavity to act as an in vivo bioreactor. However, these current methods may possibly be useful if applied after graft removal and aortic implantation to prevent the formation of adhesion in the long term.

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1.6 Hydrogels

Our method to prevent adhesions is a non-degradable hydrogel. Hydrogels consist of three-dimensional hydrophilic polymers that are formed by crosslinking of polymer chains either by chemical or physical crosslinking. Hydrogels can retain more water, up to 500 times of their own weight [98] [99]. This characteristic can be attributed of the hydrophilic components like carboxyl, amide, amino, and hydroxyl groups in the polymer chains [100]. Due to the crosslinking between the chains they do not dissolve in the solution and thus retain the three dimensional structure. This property to retain water in the biological conditions make them more advantageous over other biomaterials for a wide range of applications ranging from food industry to clinical and biomedical applications [100] [101].

1.6.1 Hydrogels in tissue engineering

The hydrogels have a significant variety of applications due to their distinctive properties. Their properties like increased water absorption, molecular stability and crosslink network porous structure provides advantages for drug delivery applications and biomedical implants [102–106]. Hydrogels have been used in tissue engineering for applications like production of 3D scaffolds because of their modifiable compositions, porosity, surface morphology and the ability to adjust their shape according to the need. The composition of the hydrogels can be modified in

20 terms of mechanical strength and be used as bioreactor which bring in stimuli like hydrostatic pressure, compression, shear stress.

1.6.2 Biocompatibility

For any synthetic hydrogels when used in tissue engineering, it is essential that the material is biocompatible. Biocompatibility can be defined as the ability of the material in contact with the body organs not to cause any kind of damage to the surrounding tissue and without eliciting any detrimental response[107].

Biocompatibility is the ability of a biomaterial to accomplish its desired function with respect to a treatment, without eliciting any undesirable local or systemic effects in the recipient or beneficiary of that therapy, but generating the most appropriate beneficial cellular or tissue response in that particular condition, and optimizing the clinically relevant performance of that therapy[107]. The natural polymers that are available are known for their biocompatibility, as they do not trigger any undesirable response while, the synthetic hydrogels can be modified in terms of their chemical structure to make them biocompatible [108]. The highly hydrophilic nature of the hydrogels leads to low adherence of the proteins to the surface because of the low interfacial free energy when they come in contact with the body fluids [109]. Any toxicity that is produced due to the synthetic hydrogels may be due to the unreacted monomers or oligomers that leak out during polymerization.

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1.6.3 PEGDA

Poly ethylene (glycol) hydrogel is widely used in tissue engineering for its properties like low toxicity and high water absorption [110]. Photo polymerization is the most commonly used method in the fabrication of PEGDA gels. 1-[4-(2-hydroxyethoxy)- phenyl]-2-hydroxy-2-methyl-1-propanone (or Irgacure 2959) is a photo initiator typically used in the preparation of gels. Irgacure is a light sensitive, non-toxic compound that is usually used at low concentrations of 0.01% to 0.05% (w/v). The higher concentrations of Irgacure (>0.05%) were reported to be cytotoxic. The concentration of PEG materials can be modified to make them biocompatible and support the cells due to their properties [111,112]. Studies have shown that as the density of the polymer increases the cell viability decreases due to the restriction of

ECM production as the cells cannot spread out and does not allow any king of cell attachment [113–115]. PEGDA gels have been investigated for their property to prevent protein adsorption for endothelial cells which is attributed to the uncharged hydrophilic groups forming hydrogen bonds with the surrounding molecules [116].

For this reason, PEGDA gels are very hydrophilic and have been used during a tissue injury or coating the tissue surface to prevent thrombosis. All these properties of

PEGDA make it a suitable scaffold for enclosing the conduits prior to the aortal implantation and additionally preventing the adhesions in the peritoneal cavity.

22

1.6.4 Mechanics of hydrogels

Apart from biocompatibility, the mechanical properties of the hydrogel are equally important and the gel must be designed to address the requirements before implanting them into the body. For our peritoneal strategy it is possible that the hydrogel can fracture in the body if it cannot hold the force exerted on it by the internal organs like the intestines liver. The gel must maintain its integrity and strength to withstand the forces inside the body exerted by the organs. The toughness of the hydrogel can be increased by adding monomers or incorporation of crosslinking agents [117].

However, higher degree of crosslinking might lead to brittleness of the gel and ultimately lead to fracture. So, it is believed that modification can be done by varying the concentration of PEG used to obtain the required mechanical strength.

Compression of the gels at similar force exerted in the body can be performed to determine the required degree of crosslinking. Although the forces inside the body at a particular time are uncertain, the experiments can be optimized with tests like three point bending and determine the right degree of crosslinking.

1.7 Strategy

The overall goal of this research study was to develop a new pouch allowing the peritoneal cavity to act as an in vivo bioreactor as well as reduce side effects caused due to adhesions. It was hypothesized that PEG being a hydrophilic polymer, would limit protein adsorption and cell attachment. The peritoneal cavity acts as a bioreactor

23 for the implanted conduits by providing the required temperature and recruiting autologous to conduits. The study aimed at determining the concentrations of

PEGDA hydrogels that are suitable to implant in the peritoneal cavity in terms of mechanical stability. A pilot study was done by implanting a PEG pouch containing electrospun conduits in a peritoneal cavity for four weeks. The conduits were then analyzed the response elicited. The conduits and the fluid from the peritoneal cavity were assessed for the type of cells. The conduits were produced by electrospinning

PCL as described in previous studies [4].

The approach of this study was varying the number of pores on the surface of the hydrogel before implanting it in the peritoneal cavity. Analysis was done after four weeks to investigate the properties of PEGDA in terms of cell attachment and the inflammation response after in the body. Mechanical properties of the hydrogel were also investigated to make it more stable. The strategy used in this technique was by varying the concentration of PEG and compressing them at a constant load rate. The specific mechanical properties were then analyzed.

24

CHAPTER 2 MATERIALS AND METHODS

2.1 Materials

All disposables, chemicals and biological supplies were purchased from

Fisher Scientific (Pittsburgh, PA) unless specified otherwise. Poly (ethylene) diacrylate (PEGDA) and Irgacure 2959 were purchased from ESIBio (Alameda,

CA). Polycaprolactone (PCL) with inherent viscosity 1.0-1.3 dL/g in chloroform was purchased from Lactel Absorbable Polymers (Pelham, AL). Polydimethylsiloxane

(PDMS) was purchased from Ellsworth adhesives (Germantown, WI). All antibodies were purchased from Abcam (Cambridge, MA).

2.2 Production of conduits

The vascular conduits were produced by electrospinning PCL. The solution concentration was maintained at 12% w/v PCL in 1, 1, 1, 3, 3, 3-hexafluoro-2- propanol (HFIP) in PCL. The electrospinning was performed using a 22 gauge needle with a voltage potential of 15 kV. The throw distance was maintained at 10 cm and the flow rate was maintained at 0.8 mL/h. The electrospinning was performed on a cylindrical rod with an outer diameter of 1.6 mm to produce small diameter conduits.

The consistency in the thickness of the conduits and the random orientation of the fibers was achieved using lateral movement and a slow axial rotation (<100 rpm) of

25 the cylindrical rod. The conduits were then carefully removed from the rod. The conduits were then cut at an optimal length and stored in a desiccator until use.

2.3 Producing PEG pouches

The pouch was prepared by the crosslinking of PEGDA solution which takes the shape of the mold during crosslinking (Fig. 3). The molds were prepared using

PDMS. PDMS molds were cut in the shape of hollow rectangle. A transparent glass piece of dimensions 1.9 x 1.4 x 0.3 cm was placed in between the three stacks to make enough hollow space in the hydrogel to put the contents inside. The whole set up was sealed with parafilm around the perimeter to prevent any leaks.

Hydrogels were prepared using PEGDA at 15% (w/v) in DI water at 37°C. To this mixture, Irgacure 2959 was added at 0.5% (w/v). The solution was protected from light. The mixture was agitated in a vortex until a completely dissolved solution was formed. All the apparatus was disinfected using 70% ethyl alcohol. The solution was filter sterilized and was added in to the PDMS mold and placed under a UV lamp

(365 nm) for 20 minutes until the gels were solidified. Figure 2 shows a schematic of production of PEG pouch.

26

SIDE VIEW

PEG solution Electrospun conduit Silicone tube Using U V glass rod

U V

PEG solution Figure 3. Shows the schematic of production of PEG pouches. The pouch is produced using PDMS molds and conduits are placed in it. The pouch is then enclosed using more PEG

Two kinds of pouches were made which differed in the number of pores on each surface. The pores were made using a sterile glass rod of 1.4 mm diameter. The two different conditions had 3x2 and 4x5 pores on each surface respectively. The pores are present for the conduits to be in contact with the peritoneal cavity for recruitment of cells. The glass piece was removed carefully and electrospun vascular conduits each separated by silicone tubes were placed in carefully placed in to the pouch. The

27 top surface was then closed using additional PEGDA solution. The pouches were soaked in DI water to prevent any shrinkage.

2.4 Implantation of PEGDA pouches in peritoneal cavity.

Pouches were implanted intraperitoneally into male Sprague-Dawley rats (Charles

River, Wilmington, MA) according to an IACUC-approved protocol 14-04. The surgical procedure involved laparotomy on the ventral side and implantation of the

PEGDA pouch. The pouches along with the enclosed conduits were removed after four weeks of implantation. The pouches and the conduits were subjected to analysis of recruited cells. Peritoneal fluid was also collected from the rat for analysis of inflammatory cells.

2.5 Cryosection

The conduits and the pouches were mounted in an optimal cutting temperature compound (OCT) (Tissue-Tek, Torrance, CA) and frozen at -80˚C. 10 μm thick cross-sections were prepared by cryosectioning.

2.6 Hematoxylin and Eosin staining

Sections were stained with hematoxylin and eosin (H&E). Histology was performed to analyze construct remodeling and cell infiltration within the scaffold and cellular

28 deposition on the pouch surface. After staining, tissue sections were cover-slipped using a mounting medium with limonene as the solvent (Electron Microscopy

Sciences, Hatfield, PA). Image pro Plus® (Media Cybernetics) was used to take images.

2.7 Immunofluorescent Imaging

The peritoneal fluid obtained after four weeks of implantation was subjected to immunofluorescence staining. The samples were placed in 24 well plates and were fixed with 4% w/v EM-grade formaldehyde (Electron Microscopy Sciences,

Hatfield, PA) and washed thrice using phosphate buffered saline (PBS) for 5 min.

Samples were then permeabilized with 0.1% v/v Triton X-100 (VWR) for 5 min, washed using PBS thrice and then blocked with 5% v/v goat serum (Fischer

Scientific, Pittsburg, PA) for 30 minutes. The primary antibodies α-smooth muscle actin (rabbit monoclonal at 1:100 dilution) and calponin (rabbit monoclonal at 1:100 dilution) were used for the investigation of contractile SMCs. Thrombospondin

(mouse monoclonal at 1:100 dilution) was used to detect the cell proliferation. CD80

(mouse monoclonal at 1:100 dilution) was added to detect the presence of M1 macrophages and Von willebrand factor (rabbit monoclonal at 1:200 dilution) was used to detect the presence of endothelial cells. Alexa 633 conjugated secondary antibodies (Life Technologies) were used and they were either anti-mouse highly-

29 cross absorbed or anti-rabbit depending on the primary antibody. ImagePro Plus®

(Media Cybernetics) software with a custom macro was used.

2.8 Mechanical testing

Mechanical testing of the PEG hydrogels was performed to characterize and modify the toughness of the hydrogels. These tests were made using a mechanical testing system (Instron E3000, Norwood, MA). Hydrogels with different concentrations were placed between two parallel plates and subjected to unconstrained compression until failure at a constant compression rate of 25 min-1.

The hydrogels at different conditions were similar in dimensions with a diameter of 1.00 ± 0.05 cm and height of 1.23± 0.05 cm. They were soaked in PBS overnight prior to testing.

The load and extension values at different points were determined from wave matrix software. The stress and strain curves were plotted from the obtained results. The ultimate stress (compression) was calculated at the end point representing first fracture. The percent compression (strain) was calculated from the average strain at the first fracture. The sample size was n = 3 at different conditions.

The modulus at 10%, 20% and 30% strain was derived from the stress and strain curves.

30

2.9 Statistics

Results were presented as mean ± standard deviation for all the mechanical testing including hydrogel measurements, ultimate compression, percent compression and modulus. For peritoneal implantation, n = 1 was used for each condition. For the imaging of immunofluorescence and H&E staining studies, n = 3 (3 images/sample) were used for analysis. For mechanical testing, n=3 samples were used for each condition. Statistical analysis was performed using JMP software. The statistical significance was determined using one-way ANOVA with Turkey post-hoc and data falling under normal distribution with a p≤ 0.05 was noted to be significant.

31

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3.2 Pilot Implantation Study

Implantation was performed for 4 weeks and images were taken post peritoneal

implantation prior to aortal grafting of the conduits. The rats were normal and no

surgical complications were observed. Figure 5 A shows the ePTFE pouch 4 weeks

after implantation where the adhesions were formed through pouch. Figure 5B

contains PEG pouch which show that the peritoneal adhesions were present however,

they were significantly lower and not densely formed which could be easily removed

using cuetips. This made the grafting process easier and may also reduce the

adhesions in the peritoneal cavity. According to the grading scale of adhesions the

adhesions could be graded by visual analysis [92]. We graded the PTFE pouch could

be graded between 3-4 while the adhesions from the PEG pouches could be graded

between 0-1.

A B

Figure 5. Shown are two different pouches implanted in the peritoneal cavity for 4 weeks. (A) shows PTFE pouch with dense adhesions formed. (B) Shows the limited adhesions caused due to PEG pouch. The arrow marks show the adhesions that are formed. The adhesions are noticeable where the pores are present.

33

A B

Figure 6. Shows PEG hydrogel in peritoneal cavity post implantation. (A) Shows a strip of PEG hydrogel with negligible adhesions due to absence of pores. (B) Shows fractured part of hydrogel believed to be caused due to forces inside the body.

The adhesions that were caused due to the PEG pouch were caused only at the sites where pores were present. It can be seen in figure 6A where the PEG strip was implanted but adhesions were negligible. Although the PEG pouch showed favorable results with less adhesions, it was observed that the pouches were fractured during the period of implantation. The fracture of the hydrogel appeared to be a clean fracture (Fig 6B). The forces inside the body are believed to cause a fracture in the hydrogels. The stress values in the peritoneal cavity is not certain at this point. Thus it is important for us to characterize the mechanical properties of our hydrogels and develop tough hydrogels to withstand the forces inside the peritoneal cavity and produce a more functional pouch.

34

3.3 H&E Staining

The H&E staining of the cross-sections of the electrospun conduits show that the cells were infiltrated with the stain. Figure 7 A and B show the conduits from 3x2 and 4x5 pores on the surface of the pouch respectively. The grafts were open during the implantation and no tissue formation was found in the lumen. It is evident from the figures that the pouch having the less number of pores had more cellular infiltration when compared to the pouch having more number of pores. However, little can be discussed about this since both the pouches were fractured and the sample size was n=1. So, the implantation might have to be repeated with a larger sample size.

200 um 200um

A B

Figure 7. Shows H&E staining of sections of conduits from 2 different conditions. (A) Shows conduit from 3x2 pouch. (B) Shows conduit from 4x5 pouch. Both the .conduits show recruitment of cells

35

The H&E staining of cross-sections of the pouches show the inflammatory response to the pouches when implanted in the body. There is a significant difference in the types of cells that are observed with pouch material. Figure 8 A shows the cross- section of a PTFE pouch. It can be seen that the cells are elongated and spread along with connective tissue. Figure B shows the cross-section of a PEG pouch where the cells appear to be more dispersed without connective tissue.

The adhesions on the PEG pouch seem to be reduced. The very hydrophilic nature of PEG can be attributed to the non-adherence of the cells to the pouch [118].

A B

100 um 100 um

Figure 8. Shows cross sections of 2 different pouches. (A) Shows cross-section of PTFE pouch. The tissue is evident and the cells are well spread along with the presence of connective tissue. The arrow mark indicates the pouch material (PTFE) (B) shows the cross section of PEG pouch. The cells are not well spread and there is negligible connective tissue present.

36

3.4 Peritoneal Fluid

Images of the peritoneal fluid that was extracted 4 weeks after PEG pouch implantation were taken at a magnification of 20x (Figure 9). The cells appear to be mostly round.

A B

100um 100um

Figure 9. Shows cells from peritoneal fluid after pouch implantation. (A) Shows peritoneal fluid from the condition 3x2 pores. (B) Shows peritoneal fluid from the condition 4x5. Many cells appear to be round.

Peritoneal fluid cell expression for contractile markers Calponin, α Smooth muscle actin, and a marker for M1 macrophages( Cd80) and thrombospondin, marker for cell proliferation and Von willebrand factor, marker for endothelial cells were done after 4 weeks of implantation with endothelial cells as control (Figure 10 and 11).

37

The cells expressed the contractile markers (Cnn1 and αSMA) for both experimental conditions and the expression for the pore size 3x2 seemed to be higher.

However, the there was no expression found for the M1 macrophage marker CD80, endothelial cell marker or cell proliferation marker. The expression for the vWF was expected for the presence of mesothelial cells as the adhesion formation is expected to cause the activation of mesothelial cells in the peritoneal cavity [87].

Cnn1 CD80 αSMA

3x2

100 um 100 um 100 um

4x5

100 um 100 um 100 um

Figure 10. Shows representative images for contractile markers calponin and α smooth muscle actin and M1 macrophage marker CD80 for the conditions 3x2 and 4x5 number of pores in the pouches. The nuclei are stained with DAPI and shown in blue.

38

Thbs2 vWF

3x2

4x5

Figure 11. Shows the representative images for the cell proliferation marker, thrombospondin and endothelial cell marker vWF. Nuclei are stained with DAPI and are shown in blue

3.3 Mechanical Testing

Compressive testing was performed to characterize the strength of PEG hydrogels.

The hydrogels at different concentrations (n=3) were tested by unconstrained compression. Median curves were determined for the resultant stress and strain.

39

Figure 12 shows the stress strain curves. The curves show the multiple breaks; however, the first breaking point was used for the quantification of results.

stress vs strain 1

0.9

0.8

0.7

0.6

0.5 10%

0.4 15% 20% 0.3 Stress (MPa) Stress 0.2 0.1

0 0 0.2 0.4 0.6 0.8 1 1.2 Strain (fraction)

Figure 12. Representative stress strain curve of the compression of PEG hydrogels at different concentrations. The compression was done at a constant rate of 25 min-1. The arrows indicate the peak where the first fracture was formed.

The percent compression of the hydrogels does not show any significant variation for the three concentrations present (fig. 13). The average percent compression for all the conditions was found to be 45%.

40

compression percent

70

60

50

40

30

20

Percent compression 10

0 10% 15% 20% Concentration of PEGDA

Figure 13. Shows percent compression of different concentration of PEGDA at a constant compression rate of 25min-1

The ultimate compression results are shown in figure 14. The compression rate increases with increase in concentration of PEG. An increase in the PEG concentration increases the density and stiffness of the material increasing overall strength. However, there was no statistical difference observed for this small sample size.

41

ultimate stress 350 300 250 200 150

stress (KPa)stress 100

50

0 10% 15% 20% concentration of PEGDA

Figure 14 Shows ultimate compression stress for different concentrations of PEGDA for n=3 for each group

The tangential moduli were calculated at 10%, 20% and 30% strain rate. There was a significant difference between the percent compositions measured. Results show that significant difference is due to the variability. The modulus increases as concentration increases. When the ultimate stress and the percent compression of the data are compared, it is evident that the compression percent has not showed any trend; however in spite of lack of significant difference, the ultimate compression follows a trend with an increase in the concentration of PEG. As density increases with PEG concentration, this is not an unexpected finding.

42

tangential modulus 1000 *# 900 800 700 *# 600 500 *@ *# 10% strain 400 20% strain *@ modulus(KPa) 300 30% strain #@ *@ Figu 200 #@ re 16 *@ #@ show 100 #@ *@ s 0 #@ *@ tange 10% 15% 20% #@ percentage*@ of ntialPEGDA #@ modu *@ li of Figure 15. Shows#@ tangential *@moduli ofdiffer different concentration of PEGDA. The data had significant#@ differences represented by * from 10%, # from 15% *@ ent and @ from 20% PEGDA (all p values < 0.0165) #@ *@ conc #@ entra *@ tions #@ *@ of #@ *@ PEG

#@ *@ DA. #@ The *@ data #@ *@ had #@ signif

icant differ ences repre sente d by * from 10%, # from 15% 43 and

@ from 20% PEG DA (all p value

CHAPTER 4 CONCLUSION AND FUTURE WORKS

4.1 Conclusion

In this research study it was demonstrated that PEG hydrogels can be used to aid the peritoneal cavity in serving as a bioreactor. It was also evident that the peritoneal adhesions were reduced when compared to using ePTFE pouches. Analysis using the adhesion grading scale demonstrated the adhesions scaled down from a grade of 3-4 to 0-1 according to Mazuji classification. As stated in the previous sections, the range of 0-2 is acceptable. A higher grade might be a clinical problem. The main aim of the project was to meet the needs of a pouch to carry the electrospun conduits in the peritoneal cavity. It was evident that this was achieved as the results indicate the infiltration of the cells on the conduits for the pouches implanted in the pilot study.

There was variability in the extent of cell infiltration between pouches and a detailed study will be needed to determine the impact of different pores on construct generation and prevention of peritoneal adhesions. When the peritoneal fluid was analyzed, it was found that the isolated cells marked for the contractile markers but there were no traces of the expression for macrophages as well as endothelial cells.

The mechanical testing analysis was performed to study the fracture of hydrogels with different concentrations of PEG at higher loading rates when compared to a previous study. It was found that the compressive strength increased with the

44 increase in concentration of PEG. This could be due to the strength attained from increased density, as well as a higher amount of crosslinking density. Although there was a difference that was observed with increase in concentration of PEG, there was not any significant difference noted in the percent compression due to the low sample size. Overall, the results were favorable in terms of the inflammatory response. The results also demonstrate that important considerations such as mechanical strength and indicates need to be considered in further research.

4.2 Future Work

The future work is focused on improving the toughness of the hydrogels and increasing their longevity in the peritoneal cavity. It is important to optimize their characteristics suitable to the conditions that are existing in the body. This study focused on producing a pouch to carry the conduits in the peritoneal cavity; however, it was found that the pouch was fractured during the period of implantation. For future work, we will be focusing on improving the functionality of the PEG pouches by modifying the composition of PEG hydrogel such that it can withstand the forces inside the peritoneal cavity. The forces inside the body are not characterized for this condition. However, the hydrogels can be tested using three point bending to characterize the properties that may have been related to fracture of hydrogels in this study. The mechanical testing will be performed at different rates of compression to analyze which concentration of PEG hydrogel can resist higher rates of load. Three

45 point bending test might be one of the most relevant tests as the intestines might be applying pressure at a particular point of the pouch and lead to bending fracture of the gel.

It was found that the adhesions on the pouch were formed at the place where the pores were present. The future work in in this area would include CAD analysis for of the geometrical analysis of the pouch and study the effects of the number of pores. Finite element analysis can be done to assess the nutrient diffusion through the pores and the effect of size of pores on the rate of inflammation with regard to the adhesion formation as well as the overall effect of it serving as a pouch for the peritoneal cavity as a bioreactor.

46

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