TOWARDS DEVELOPMENT OF AFFINITY POLYMER-BASED ADHESION
BARRIERS FOR SURGICAL MESH DEVICES
by
GREG DANIEL LEARN
Submitted in partial fulfillment of the requirements for the degree of
Doctor of Philosophy
Dissertation Advisor: Horst A. von Recum, PhD
Department of Biomedical Engineering
CASE WESTERN RESERVE UNIVERSITY
May 2021 CASE WESTERN RESERVE UNIVERSITY
SCHOOL OF GRADUATE STUDIES
We hereby approve the thesis/dissertation of
Greg Daniel Learn
Candidate for the degree of Doctor of Philosophy *.
Committee Chair: Jeffrey Capadona, Ph.D.
Committee Member: Horst von Recum, Ph.D.
Committee Member: Kathleen Derwin, Ph.D.
Committee Member: Guang Zhou, Ph.D.
Committee Member: Michael Rosen, M.D.
Date of Defense: Monday 14 December 2020
*We also certify that written approval has been obtained
for any proprietary material contained therein.
2
This dissertation is dedicated to the memory of Jim Simpson. Though not able to see this,
as a scientist, engineer, runner, and family friend I looked up to from a young age, he furthered my inspiration and resolve to pursue and complete my degree. Thank you, Jim.
3 TABLE OF CONTENTS
TABLE OF CONTENTS ...... 4
LIST OF TABLES ...... 12
LIST OF FIGURES ...... 13
PREFACE ...... 16
ACKNOWLEDGEMENTS ...... 17
LIST OF ABBREVIATIONS ...... 20
ABSTRACT ...... 23
1. CHAPTER 1: PATHOGENESIS AND PREVENTION OF ABDOMINAL
ADHESIONS, AND THE ROLE OF SURGICAL MESH BIOMATERIALS ...... 25
1.1. Abstract ...... 25
1.2. Introduction ...... 26
1.2.1. Prevalence of Post-Surgical Adhesions ...... 28
1.2.2. The Relation of Peritoneal Adhesions to Surgical Meshes ...... 30
1.2.3. Regulatory Environment for Mesh Adhesion Prophylaxis ...... 32
1.2.4. Purpose ...... 33
1.3. Clinical Significance of Peritoneal Adhesions ...... 34
1.3.1. Complications ...... 34
1.3.2. Socioeconomic Burden ...... 37
1.3.3. Insufficient Awareness ...... 38
1.4. Clinical Treatment of Peritoneal Adhesions ...... 39
1.4.1. Surgical Adhesiolysis...... 39
1.4.2. Physiotherapy ...... 39
4 1.5. Pathogenesis of Peritoneal Adhesions and Relation to Mesh Biomaterials ...... 40
1.5.1. The Healthy Peritoneum ...... 40
1.5.2. Adhesion-Free Peritoneal Healing ...... 42
1.5.3. Abnormal Peritoneal Healing and Adhesion Formation ...... 44
1.5.3.1. Tissue Disturbance ...... 44
1.5.3.2. Fibrin Deposition ...... 45
1.5.3.3. Proximity of Surfaces ...... 45
1.5.3.4. Time for Adhesion Initiation and Maturation ...... 46
1.5.3.5. Inflammation ...... 47
1.5.3.6. Ischemia and Hypoxia ...... 48
1.5.4. The Role of Surgical Meshes in Peritoneal Adhesion Formation ...... 48
1.5.4.1. Effects of Material Properties ...... 49
1.5.4.2. Effects of Structural Characteristics ...... 52
1.6. Prophylaxis of Peritoneal Adhesions ...... 56
1.6.1. General Preventive Strategies ...... 56
1.6.1.1. Surgery-Related Factors - Techniques and Instrumentation ...... 56
1.6.1.2. Pharmacologic Strategies ...... 59
1.6.1.3. Barrier Materials - Membranes and Gels ...... 63
1.6.2. Mesh-Specific Preventive Strategies ...... 68
1.7. Emerging Opportunities ...... 73
1.8. Conclusions and Future Outlook ...... 78
1.9. Acknowledgments ...... 80
5 2. CHAPTER 2: NONTHERMAL PLASMA TREATMENT IMPROVES
UNIFORMITY AND ADHERENCE OF CYCLODEXTRIN-BASED COATINGS ON
HYDROPHOBIC POLYMER SUBSTRATES ...... 81
2.1. Abstract ...... 81
2.2. Introduction ...... 82
2.3. Materials and Methods ...... 86
2.3.1. Materials ...... 86
2.3.2. Plasma Cleaning and Activation of PP Substrate Surfaces ...... 87
2.3.3. Effects of Plasma Treatment on PP Substrates ...... 88
2.3.3.1. Wettability - Contact Angle Goniometry ...... 88
2.3.3.2. Surface Chemistry - XPS ...... 88
2.3.4. pCD Synthesis and Coating onto Surfaces ...... 90
2.3.5. Effects of Plasma Treatment on pCD Coatings ...... 90
2.3.5.1. Qualitative Uniformity - Direct Visualization ...... 90
2.3.5.2. Qualitative Uniformity - SEM ...... 91
2.3.5.3. Semi-Quantitative Uniformity - Direct Visualization ...... 92
2.3.5.4. Adherence - Lap-Shear Testing ...... 92
2.3.5.5. Interfacial Covalent Bonding - XPS ...... 93
2.3.6. Statistical Analysis ...... 94
2.4. Results ...... 95
2.4.1. Effects of Plasma on PP Substrate Wettability and Surface Chemistry ..... 95
2.4.2. Effects of Plasma on pCD Coating Uniformity, Adherence, and Interfacial
Covalent Bonding ...... 99
6 2.5. Discussion ...... 105
2.6. Acknowledgments ...... 108
3. CHAPTER 3: NONTHERMAL PLASMA TREATMENT OF POLYMERS
MODULATES BIOLOGICAL FOULING BUT CAN CAUSE MATERIAL
EMBRITTLEMENT ...... 109
3.1. Abstract ...... 109
3.2. Introduction ...... 110
3.3. Materials and Methods ...... 113
3.3.1. Materials ...... 113
3.3.2. Plasma Treatment ...... 114
3.3.3. Surface Characterization ...... 115
3.3.3.1. X-ray Photoelectron Spectroscopy ...... 115
3.3.3.2. Fibrinogen Adsorption ...... 116
3.3.3.3. Bacterial Attachment ...... 116
3.3.3.4. Fibroblast Attachment ...... 117
3.3.4. Bulk Characterization ...... 118
3.3.4.1. Uniaxial Tension Testing of PP Monofilaments ...... 118
3.3.4.2. Uniaxial Tension Testing of PP Meshes...... 119
3.3.4.3. Suture Retention Testing of PP Meshes ...... 120
3.3.4.4. Tear Resistance Testing of PP Meshes ...... 121
3.3.4.5. Ball Burst Testing of PP Meshes ...... 121
3.3.5. Statistical Analysis ...... 122
3.4. Results ...... 123
7 3.4.1. Surface Characterization ...... 123
3.4.1.1. X-ray Photoelectron Spectroscopy of PP Meshes ...... 123
3.4.1.2. Fibrinogen Adsorption ...... 126
3.4.1.3. Bacterial Attachment ...... 127
3.4.1.4. Mammalian Fibroblast Attachment ...... 128
3.4.2. Bulk Characterization ...... 129
3.4.2.1. PP Monofilament Uniaxial Tension Testing ...... 130
3.4.2.2. PP Mesh Uniaxial Tension Testing ...... 131
3.4.2.3. Mesh Suture Retention Testing ...... 134
3.4.2.4. Mesh Tear-Resistance Testing ...... 135
3.4.2.5. Mesh Ball Burst Testing ...... 137
3.5. Discussion ...... 138
3.6. Conclusions ...... 143
3.7. Acknowledgments ...... 144
4. CHAPTER 4: CYCLODEXTRIN POLYMER COATINGS RESIST FOULING BY
PROTEINS, MAMMALIAN CELLS, AND BACTERIA...... 145
4.1. Abstract ...... 145
4.2. Introduction ...... 146
4.3. Materials and Methods ...... 151
4.3.1. Materials ...... 151
4.3.2. Plasma Cleaning and Activation of PP Substrate Surfaces ...... 152
4.3.3. pCD Synthesis and Coating onto Surfaces ...... 153
4.3.4. Effects of HDI-Crosslinking on pCD Physicochemical Properties ...... 154
8 4.3.4.1. Determination of pCD Swelling Ratio ...... 154
4.3.4.2. Unconfined Compression Testing of pCD ...... 155
4.3.4.3. Contact Angle Goniometry ...... 156
4.3.4.4. Attenuated Total Reflectance FTIR Spectroscopy ...... 156
4.3.5. Effects of HDI-Crosslinking on pCD Anti-Biofouling Performance ...... 157
4.3.5.1. Evaluation of Protein Adsorption ...... 157
4.3.5.2. Investigation of Mammalian Cell Attachment and Viability ...... 159
4.3.5.3. Measurement of S. aureus Attachment ...... 159
4.3.5.4. Assessment of E. coli Attachment ...... 161
4.3.6. Statistical Analysis ...... 162
4.4. Results ...... 163
4.4.1. Effects of HDI-Crosslinking on Physicochemical Properties of pCD ...... 163
4.4.2. Effects of HDI-Crosslinking on Protein Adsorption to pCD ...... 167
4.4.3. Effects of HDI-Crosslinking on Fibroblast Adhesion to pCD ...... 169
4.4.4. Effects of HDI-Crosslinking on Bacterial Attachment to pCD ...... 172
4.5. Discussion ...... 174
4.6. Conclusions ...... 178
4.7. Acknowledgments ...... 179
5. CHAPTER 5: PNEUMOPERITONEAL COMPUTED TOMOGRAPHY AS A PRE-
CLINICAL APPROACH FOR VISUALIZING INTRA-ABDOMINAL ADHESIONS TO
PROSTHETIC MESH ...... 180
5.1. Abstract ...... 180
5.2. Introduction ...... 181
9 5.3. Materials and Methods ...... 184
5.3.1. Materials ...... 184
5.3.2. pCD Preparation...... 184
5.3.3. Mesh Preparation and Sterilization ...... 184
5.3.4. Animals ...... 186
5.3.5. Mesh Implantation and Animal Recovery ...... 187
5.3.6. Animal Sacrifice and Pneumoperitoneal CT ...... 187
5.3.7. Animal Dissection and Adhesion Grading ...... 189
5.3.8. Three-Dimensional Image Segmentation and Analysis ...... 190
5.4. Results ...... 192
5.5. Discussion and Conclusions ...... 198
5.6. Acknowledgements ...... 202
6. CHAPTER 6: CONCLUSIONS AND FUTURE DIRECTIONS ...... 203
6.1. Conclusions ...... 203
6.1.1. Utility of Plasma for Improving pCD Coatings on PP Substrates ...... 203
6.1.2. Versatility of CD-Based Polymers...... 204
6.1.3. Tradeoffs with pCD HDI Crosslinking: Biofouling Resistance versus
Mechanical Robustness ...... 205
6.1.4. Tradeoffs with Plasma Treatment: PP Surface versus Bulk Properties .... 205
6.1.5. Ability of Plasma Treatment to Directly Modulate Protein Adsorption,
Fibroblast Attachment, and Bacterial Attachment to PP Substrates ...... 206
6.1.6. Value of pCD as a PP Mesh Adhesion Barrier, and of Pneumoperitoneal CT
as a Method for Visualizing Adhesions to Intra-Abdominal Mesh ...... 207
10 6.2. Future Directions ...... 208
6.2.1. Application of pCD Coatings to Other Polymeric Substrate Materials .... 208
6.2.2. Evaluation of Drug-Loading Impacts on pCD Biofouling Resistance ..... 209
6.2.3. Reformulation of pCD to Achieve Crosslinking-Independent Resistance to
Biofouling, and/or Greater Extensibility ...... 210
6.2.4. Investigation of Impacts of Plasma Exposure and Carrier Gas on Different
Mesh Materials ...... 216
6.2.5. Exploration of the Impacts of Plasma Treatment on Mesh-Related
Complications ...... 217
6.2.6. Further Studies on Adhesion Prophylaxis Using pCD Barriers, and
Refinement of Pneumoperitoneal CT ...... 218
APPENDIX ...... 219
Permission for Reprint of Chapter 2 ...... 219
Permission for Reprint of Chapter 3 ...... 219
Permission for Adaptation of Table 5-1 ...... 220
REFERENCES ...... 221
11 LIST OF TABLES
Table 1-1...... 57
Table 1-2...... 58
Table 1-3...... 60
Table 1-4...... 61
Table 1-5...... 65
Table 1-6...... 66
Table 1-7...... 67
Table 1-8...... 70
Table 1-9...... 71
Table 2-1...... 96
Table 2-2...... 101
Table 2-3...... 102
Table 2-4...... 104
Table 3-1...... 130
Table 3-2...... 133
Table 3-3...... 133
Table 3-4...... 136
Table 3-5...... 138
Table 4-1...... 154
Table 5-1...... 190
Table 5-2...... 197
Table 5-3...... 197
12 LIST OF FIGURES
Figure 1-1...... 27
Figure 1-2...... 27
Figure 1-3...... 29
Figure 1-4...... 32
Figure 1-5...... 75
Figure 2-1...... 86
Figure 2-2...... 97
Figure 2-3...... 98
Figure 2-4...... 100
Figure 2-5...... 100
Figure 2-6...... 102
Figure 2-7...... 103
Figure 2-8...... 104
Figure 3-1...... 113
Figure 3-2...... 124
Figure 3-3...... 124
Figure 3-4...... 125
Figure 3-5...... 126
Figure 3-6...... 127
Figure 3-7...... 128
Figure 3-8...... 129
Figure 3-9...... 130
13 Figure 3-10...... 132
Figure 3-11...... 132
Figure 3-12...... 134
Figure 3-13...... 136
Figure 3-14...... 137
Figure 4-1...... 151
Figure 4-2...... 165
Figure 4-3...... 167
Figure 4-4...... 168
Figure 4-5...... 171
Figure 4-6...... 173
Figure 5-1...... 186
Figure 5-2...... 186
Figure 5-3...... 188
Figure 5-4...... 192
Figure 5-5...... 193
Figure 5-6...... 194
Figure 5-7...... 194
Figure 5-8...... 195
Figure 5-9...... 196
Figure 5-10...... 196
Figure 6-1...... 211
Figure 6-2...... 211
14 Figure 6-3...... 212
Figure 6-4...... 213
Figure 6-5...... 215
Figure 6-6...... 215
Figure 6-7...... 216
15 PREFACE
This dissertation is organized as follows: Chapter 1 is a comprehensive review on the subject of peritoneal adhesions, and their formation, prevention, and relation to surgical mesh devices. Chapter 2 represents a published study on the use of nonthermal plasma as a strategy to improve the physical and chemical connection between polymerized cyclodextrin, the experimental material being developed herein as a potential adhesion barrier for surgical meshes, and polypropylene, the most common surgical mesh polymer.
Chapter 3 represents published research on the effects of nonthermal plasma exposure on the surface properties, biofouling resistance, and mechanical properties of polypropylene mesh substrates. Chapter 4 is an in vitro investigation on the cytocompatibility and resistance of polymerized cyclodextrin materials and polypropylene controls to protein adsorption, mammalian fibroblast attachment, and bacterial attachment. These events are involved in mesh adhesions, mesh migration, and mesh infection, respectively. Chapter 5 is a preliminary study on the effectiveness of polymerized cyclodextrin as an experimental adhesion barrier for polypropylene meshes, and also explores a novel method for non- invasively imaging mesh adhesions. Finally, Chapter 6 discusses overarching conclusions and future directions from across this body of work. With the exception of Chapter 6, all chapters are written in the form of stand-alone papers with broad relevance, such that each can tell a complete story independently of the others.
16 ACKNOWLEDGEMENTS
I would first like to extend my most genuine appreciation to my advisor Dr. Horst von Recum for welcoming me into his lab during a time of uncertainty in summer 2017.
Horst demonstrated an unparalleled enthusiasm for prioritizing the development of his mentees. He was tremendously supportive of my goals, and never hesitated to inspire me to place my professional interests first (even before needs of the lab), such as during the six months when I gained valuable career experience on co-op at BD. Horst taught me a great deal, whether it was key theoretical concepts related to research, or effective communication skills broadly transferable to any area of life. His mentorship style promoted a truly positive and inclusive culture, so my time in the von Recum lab was refreshing. I only wish I had worked with Horst for longer, for the uplifting work environment, and the many interesting research ideas we never had time to attempt. I know he has urged me to try modifying surfaces with ammonia plasma for a while now.
I am also grateful to my committee members: Drs. Jeffrey Capadona, Kathleen
Derwin, Guang Zhou, and Michael Rosen. They committed their time and effort to guide my development as a scientist and engineer, encouraging me to believe in myself and strive toward my fullest potential. As leaders in their fields, they shared their valuable wisdom and expertise with me, and I’ve learned a lot from all of them.
Next, I want to recognize my student mentees: Emerson Lai, Emily Wilson, Ashley
Djuhadi, and Katherine Yan. They devoted countless hours to working at my side, helping to perform experiments that directly made this (and other) research projects possible. They persevered even when studies didn’t work out as planned or result in a publication, and their involvement made the work infinitely more enjoyable. I deeply appreciate their
17 dedication and efforts, and I will miss their enthusiasm and amusing antics. Such as that time we caught a cockroach in Wickenden building, and secretly housed “Carl” in the drawer of a temporarily vacant desk in our office. RIP Carl.
A sincere thank you, from the bottom of my heart, goes to my friends I met at Case, for the all of the fun times and fond memories we shared, and helping me troubleshoot day- to-day problems, and pulling me through my roughest patches during my PhD. I will never forget it. Phillip McClellan and Hyungjin Jung were always willing to celebrate or commiserate with me, whether at the Jolly Scholar, or during experiments that ran until
3AM. In trying moments, they offered valuable perspective and reassurance that all challenges eventually subside. Erika Cyphert, Kathleen Young, and Sunny Lu were a tremendous moral support system, and provided a sense of camaraderie that made the lab feel like home. I valued our “Bibibop Thursday” tradition and our adventures across states and continents. Sunny had a beneficial and unique ability to make me stop taking myself too seriously. Kathleen helped me remember the importance of taking necessary mental breaks, such as socializing in lab get-togethers or game nights, and appreciating pet videos.
Erika ran with me on possibly every “Emerald Necklace” trail, shared laughs with me over our surprisingly frequent mishaps, and was always a good sport about letting me persuade her to try new foods or activities. Hopefully she can forgive me for that time her foot broke on one unlucky run I had picked.
I want to bestow my appreciation also to my manager during my time at BD,
Jonathan Trexler, and my colleagues Keith Greenawalt and Jessica Powell. They all helped me to realize my potential and grow in my career, provided me invaluable opportunities to gain experience in the medical device industry, and showed me the outstanding teamwork
18 that goes into advancing the world of health. I’ll give another shout-out to Ashley Djuhadi, for encouraging me to apply for the BD co-op position.
Last, I must express my heartfelt thanks to my loved ones and family. My partner
Katie Chapin remained steadfast at my side throughout this long journey, with all its ups and downs, and watched out vigilantly for my well-being. My parents provided a cohesive family and home throughout my life, instilled in me the drive to always go the extra mile, and have continually offered guidance as I carve out my life and career paths. I could not have completed or even begun any of this work without their unconditional support.
19 LIST OF ABBREVIATIONS
Abbreviation Meaning
(In alphabetical order)
2D Two-Dimensional
2-TPI 2-(Trifluoromethyl)phenyl Isocyanate
3D Three-Dimensional
ABF Anti-Biofouling
ASTM American Society for Testing and Materials
ATR Attenuated Total Reflectance
BD Becton, Dickinson and Company
CD Cyclodextrin
CFU Colony Forming Unit
CMC Carboxymethylcellulose
CT Computed Tomography
CWRU Case Western Reserve University
DICOM Digital Imaging and Communications in Medicine
DMEM Dulbecco’s Modified Eagle Medium
DMF N,N-Dimethylformamide
E. coli Escherichia coli
EDTA Ethylenediaminetetraacetic Acid
EGDGE Ethylene Glycol Diglycidyl Ether ePTFE Expanded Polytetrafluoroethylene
FBS Fetal Bovine Serum
20 FDA Food and Drug Administration
FITC Fluorescein Isothiocyanate
FTIR Fourier-Transform Infrared Spectroscopy
FWHM Full Width at Half Maximum
GRAS Generally Recognized as Safe
HDI Hexamethylene Diisocyanate
IACUC Institutional Animal Care and Use Committee
IPC International Plasma Corporation
LB Luria-Bertani
MRI Magnetic Resonance Imaging
NIH National Institutes of Health
NRSA National Research Service Award
O3FA Omega-3 Fatty Acid
OD600 Optical Density at 600 nm
ORC Oxidized Regenerated Cellulose
PAI Plasminogen Activator Inhibitor
PBS Phosphate Buffered Saline pCD Polymerized Cyclodextrin
PDMS Polydimethylsiloxane
PE Polyethylene
PEG Poly(ethylene glycol)
PET Positron Emission Tomography
21 PETE Poly(ethylene terephthalate)
PGA Poly(glycolic acid) pHEMA Poly(2-hydroxyethyl methacrylate)
PLGA Poly(lactic acid-co-glycolic acid)
PP Polypropylene
PTFE Polytetrafluoroethylene
PVDF Poly(vinylidene fluoride)
PS Polystyrene
RIF Rifampicin
S. aureus Staphylococcus aureus
SEM Scanning Electron Microscopy
SiC Silicon Carbide
TCPS Tissue Culture Polystyrene
TGF Transforming Growth Factor tPA Tissue Plasminogen Activator
US United States
XPS X-Ray Photoelectron Spectroscopy
22 Towards Development of Affinity Polymer-Based Adhesion Barriers
for Surgical Mesh Devices
ABSTRACT
by
GREG DANIEL LEARN
Post-surgical adhesions are internal scars that pathologically adhere together adjacent tissues/organs/biomaterials. They pose a tremendous but frequently underestimated burden across many surgical disciplines, being especially prevalent following abdominal surgery. Peritoneal adhesions can cause discomfort, intestinal obstructions, infertility, and increased morbidity/mortality of subsequent surgery. Once formed, treatments for adhesions tend to be risky and ineffective, so prophylactic strategies are desirable. Implantation of meshes, such as in hernia repair, often exacerbates peritoneal adhesions. Knitted polypropylene (PP) meshes are the most common hernioplasty devices, but are notoriously adhesiogenic owing to material and structural characteristics that promote incorporation, such as hydrophobicity and reticular construction. The ideal strategy to prevent mesh adhesions entails adhering a smooth, continuous, hydrophilic barrier material on the mesh visceral face to mitigate tissue attachment processes. Prior studies developed polymerized cyclodextrin (pCD) materials having unique capabilities for sustained, multi-window drug release, and suggested that these hydrophilic polymers
23 passively resist cell attachment. In several animal species, pCD could deliver antibiotics for weeks to successfully resolve mesh infection, another hernioplasty complication for which only suboptimal solutions exist. In the present work, pCD materials were explored toward application as novel adhesion barriers for PP surgical meshes. First, nonthermal plasma activation was assessed as a strategy to improve PP-pCD bonding, as PP is generally unreceptive to coatings. Plasma introduced hydroxyls onto PP, enhancing PP- pCD adherence. Second, protein adsorption, bacterial attachment, and fibroblast viability/attachment upon pCD-coated and bare PP materials were evaluated. These events play roles in mesh adhesion, infection, and biocompatibility. pCD decreased protein adsorption and bacterial attachment to PP, without fibroblast cytotoxicity. Third, effects of
PP plasma activation on protein adsorption, fibroblast/bacterial attachment, and mesh mechanical properties were investigated. Regardless of duration, plasma exposure of bare
PP reduced protein adsorption and bacterial attachment, and increased fibroblast attachment, but longer treatments progressively embrittled PP mesh. Fourth, preliminary studies in vivo explored effects of pCD barriers on adhesions to PP mesh. These animal experiments suggested that pCD-covered mesh surfaces resisted adhesions, while bare PP meshes did not. Altogether, pCD materials have potential as adhesion barriers that could uniquely combat both mesh adhesions and prosthetic infection.
24 1. CHAPTER 1: PATHOGENESIS AND PREVENTION OF ABDOMINAL
ADHESIONS, AND THE ROLE OF SURGICAL MESH BIOMATERIALS
Authors: Greg D. Learn, Emerson J. Lai, Horst A. von Recum
1.1. Abstract
Tissue adhesions constitute a major healthcare burden, frequently underestimated and unrecognized. Adhesions are not entirely understood, despite an overwhelming amount of information that has been, and continues to be, published on their subject. Peritoneal adhesion formation is a near-universal consequence of abdominopelvic surgery, poses a legitimate threat to patient health and quality of life, has serious implications for clinicians and healthcare systems, and is often exacerbated following implantation of common surgical mesh prostheses. Despite the advent of commercial technologies for post-surgical adhesion prevention both generally and in association with implanted mesh, evidence regarding anti-adhesion product efficacy is often conflicting. Furthermore, detailed mechanisms by which biomaterials such as surgical meshes (or adhesion barriers) may stimulate (or mitigate) adhesions are not readily evident in literature, impeding progress and consistency in clinical adhesion prophylaxis. In light of this, the present review aims to convey insights concerning the detailed interactions of mesh biomaterials with adhesions, their pathogenesis, and their prevention. Thorough awareness and understanding of adhesions, and of the roles of surgical mesh biomaterials in their formation and prophylaxis, is crucial for biomaterials researchers, device makers, clinicians, and patients, to work together towards improved clinical outcomes.
25 1.2. Introduction
Adhesions are fibrous internal scar-like tissues that form pathologic connections in the body. They may consist of vessels, nerves, and adipose 1, and are considered part of the body’s natural healing response to repair tissue-level damage after injury or homeostatic disturbance. Adhesions tether adjacent structures (e.g. tissues, organs, or implanted materials) to one another, restricting their free movement, and directly or indirectly cause a range of adverse consequences to human health and hence society (Section 1.3).
While adhesions can form in humans in the absence of surgical history (e.g. after endometriosis, Crohn’s disease, infection, or radiation treatment), surgery greatly increases adhesion incidence 2 and is the most common cause of adhesions seen clinically. Post- surgical (or post-operative) adhesions represent the broader context of this review. Post- surgical adhesions have been recognized for over 250 years 3,4 and a staggering amount of information has been published on their subject. Despite this, treatment (Section 1.4), formation (Section 1.5), and prevention (Section 1.6) of post-surgical adhesions endure as prominent areas of intensive research effort. Importantly, biomaterials can play a major role in the formation or prevention of post-surgical adhesions (Figure 1-1, Figure 1-2).
26
Figure 1-1. Post-surgical adhesions frequently cause morbidity following abdominal surgery, especially hernia repair. Intra-abdominal prosthetic mesh can worsen adhesions. Yet effects of implanted mesh biomaterials on adhesions are complex, poorly understood, and insufficiently studied. This review synthesizes detailed mechanisms and novel insights concerning mesh interactions with adhesion formation and prevention, to guide design of future mesh materials/products for improved clinical outcomes. Figure created with BioRender.com 5.
Figure 1-2. Tenacious peritoneal adhesions in a male Sprague-Dawley rat, conjoining a loop of intestine (white arrow) with an implanted bare polypropylene surgical mesh (Prolene, dashed yellow outline). A) As found, 1 month after implantation. B) After resecting the less challenging adhesions. Even with applied traction and meticulous dissection, it may not be possible to cleanly separate the bowel and prosthesis. Such adhesions elevate risks for obstruction, or inadvertent bowel perforation, mesh explantation, and sepsis upon subsequent surgery.
27 1.2.1. Prevalence of Post-Surgical Adhesions
Post-operative adhesions are prevalent across virtually all disciplines of surgery.
For example, peritendinous adhesions occur after ~4% of flexor tendon repairs 6, limiting dexterity and function of hands and fingers. Intrauterine adhesions form in ~19% of women following surgical dilatation and curettage procedures for abortion 7 or miscarriage 8, potentiating pelvic discomfort and menstrual/fertility disorders (i.e. Asherman’s syndrome). Significant epidural adhesions affect adult spine surgery patients at rates exceeding 84% 9, producing chronic pain. Pericardial and retrosternal adhesions are considered to be a universal finding following cardiac surgery 10. These cause impairment in ventricular function and cardiac efficiency 11, and increase the risk for lethal hemorrhage in subsequent surgery 12. The list continues 13,14.
As the scope of this review, peritoneal adhesions (also known as intra-peritoneal, abdominal, intra-abdominal, intestinal, or visceral adhesions) represent the most frequent subject of adhesions-related literature (Figure 1-3). Peritoneal adhesions are regarded as the most common complication after abdominal and pelvic surgery 15–17, occurring in up to 93% 18 and 55-100% of patients 19, respectively. One explanation for this ubiquitous incidence is that peritoneal adhesions reflect a defense mechanism that evolved to ensure survival; after exposure of vulnerable organs to potential attack, healing speed takes precedence over long-term tissue organization and function.
28
Figure 1-3. Total number of published articles having a title pertaining to peritoneal, uterine, tendinous, pleural, or dural adhesions. Data was collected using Web of Science 20, searching across all databases between the start of 1864 (earliest year in the database) and the end of 2020. Title search criteria for each adhesions subject were as follows. Peritoneal: ["*abdominal adhesio*" or "intestinal adhesio*" or "*peritoneal adhesio*" or "visceral adhesio*"]. Uterine: ["*adnexal adhesio*" or "Asherman's syndrome" or "fimbriolysis" or "hysterolysis" or "*ovarian adhesio*" or "*ovariolysis" or "pelvic adhesio*" or "*salpingolysis" or "*uterine adhesio*"]. Tendinous: ["*tendinous adhesio*" or "tendon adhesio*" or "tenolysis"]. Pleural: ["pleural adhesio*" or "*pulmonary adhesio*" or "*thoracic adhesio*"]. Cardiac: ["*cardiac adhesio*" or "*cardial adhesio*" or "mediastinal adhesio*" or "retrosternal adhesio*"]. Dural: ["*dural adhesio*" or "spinal adhesio*"].
29 1.2.2. The Relation of Peritoneal Adhesions to Surgical Meshes
Mesh implants are commonly used in abdominal and pelvic surgeries. Hernia repair and prophylaxis (i.e. prevention) represent the most common family of procedures (e.g. for ventral, incisional, inguinal, and parastomal hernias) involving surgical meshes. In the
United States (US), an estimated 350k ventral hernia repairs 21, >700k inguinal hernia repairs 22–25, and >4k parastomal hernia repairs 26 are performed each year. Worldwide incidence has been estimated at 20M per year for inguinal hernia repair 24,25,27. Synthetic meshes have been used for hernia repair in humans in Europe and the US since ~1948 28 and ~1957 29,30, respectively. It was only in the early 2000s, that several clinical trials showed lower hernia recurrence rates with mesh-based repair (i.e. hernioplasty) compared to primary closure (i.e. non-mesh repair using sutures only, also known as herniorrhaphy)
27,31–36, given that the mesh acts a mechanical seal and enables a “tension-free” closure 37,38.
For this reason, mesh-based closure has since become the predominant hernia repair strategy 39, with reports suggesting mesh use occurs in >65% of incisional 36 and >75% of inguinal 23,27,40 hernia repairs. Apart from hernioplasty, mesh implants are also utilized in about 33% of ~300k pelvic organ prolapse repair surgeries and over 80% of ~260k stress urinary incontinence repair surgeries each year in the US 41–43.
As a focus of this review, meshes can potentiate and exacerbate post-surgical peritoneal adhesions (i.e. mesh-related adhesions) (Section 1.5.4). Meshes are especially likely to trigger adhesion formation and related complications with intraperitoneal placement 44, which becomes inescapable in laparoscopic surgery (i.e. minimally invasive abdominal surgery) for hernioplasty 45. The material properties (Section 1.5.4.1) and
30 structural characteristics (Section 1.5.4.2) of a particular mesh product can both have a considerable impact on the extent of adhesion formation.
Currently, over 150 unique mesh products are commercially available for hernia repair, each possessing distinct designs and features 46,47. Polypropylene (PP) is the most commonly implanted material in hernioplasty 46–50. PP meshes are widely selected because the material incorporates (i.e. integrates with surrounding tissue structures via ingrowth) quickly, does not degrade in vivo, and is sufficiently strong, easy to work with, and inexpensive 51,52. Mesh incorporation with the abdominal wall is desirable to hold the mesh in place long-term. This incorporation keeps the defect sealed while circumventing mesh migration, an event in which the mesh becomes detached from its original fixation site and lodged elsewhere in the body. Unfortunately, the same characteristics that promote incorporation of PP mesh with the abdominal wall also favor adhesion formation 51,52.
Due to the increasingly recognized importance of post-surgical peritoneal adhesions, it is generally accepted that the ideal hernia mesh should deter adhesion formation on the visceral (i.e. internal organ) side, whilst promoting tissue ingrowth on the parietal (i.e. external abdominal wall) side 49,52,53. These conflicting requirements call for sophisticated meshes (Section 1.6.2). However, in contrast to the vast body of literature on adhesions overall, the relationship between surgical meshes and adhesion formation/prevention is understudied (Figure 1-4). Further confusing matters, the mesh device regulatory landscape accelerates new product introduction without requirement for clinical data (Section 1.2.3). While this aids innovation, it also leads to numerous options for clinicians and hospital product selection committees to choose between, yet insufficient evidence regarding product adhesion prophylaxis to fully justify any solution.
31
Figure 1-4. Total number of articles published during 5y periods between the start of 1950 (which roughly coincides with the advent of synthetic polymer meshes for hernia repair 28,30,54) and the end of 2019, having a title pertaining to peritoneal or post-surgical adhesions, either alone or in combination with topic matter (includes title, abstract, and keywords) related to surgical meshes. Data was collected using Web of Science 20, searching across all databases. Search criteria were as follows. Title: ["*abdominal adhesio*" or "intestinal adhesio*" or "*operative adhesio*" or "*peritoneal adhesio*" or "*surgical adhesio*" or "visceral adhesio*"], alone or in combination with Topic: ["hernia mesh*" or "mesh biomaterial*" or "mesh prosthes*" or "polymeric mesh*" or "surgical mesh*"]. Note that mesh hernioplasty was not nearly as popular until after pioneering work by Lichtenstein in the late 1980’s 37.
1.2.3. Regulatory Environment for Mesh Adhesion Prophylaxis
The considerable need for adhesion prevention in the presence of meshes has stimulated rapid commercial innovation. The FDA 510(k) mechanism facilitates swift introduction of new commercial mesh products. All surgical meshes on the US market are considered Class II medical devices (moderate risk to patient health) and were cleared through the 510(k) pathway. This requires that new devices merely demonstrate
“substantial equivalence” to predicate devices (previously cleared for the same intended use), but rarely calls for independent clinical trials 55.
32 Through the 510(k), the most certain route to clearance for mesh product manufacturers entails marketing for the general indication 56 “to reinforce soft tissue where weakness exists.” Specific indications, especially adhesion prevention - for which a mesh would instead be regulated as a Class III (high risk to patient health) adhesion barrier device
- necessitate clinical trials to prove safety and efficacy 56. Furthermore, studying mesh- related adhesions in humans presents inherent difficulties. Namely, current “gold standard” evaluation methods tend to be invasive (e.g. diagnostic laparoscopy or second-look surgery) and subjective or at best semi-quantitative (e.g. scoring). Methods that are less invasive or more quantitative do not exist partly because it is challenging to directly visualize most meshes 57 and/or adhesions 58 with common imaging modalities. For these reasons, all mesh products - including those with anti-adhesion features (herein referred to as “adhesion-resistant” meshes) - are marketed for the general indication, while no product to date has specific indication for adhesion prevention 56.
In light of this regulatory environment, it becomes difficult to stay abreast of the rapid commercial developments in mesh adhesion prophylaxis. And yet published clinical data on efficacy are missing (for the majority of products), or lagging far behind these innovations.
1.2.4. Purpose
Given that we do not fully understand adhesions, let alone the role of surgical mesh devices in their formation or prevention, the overarching purpose of this review is to convey new insights concerning the detailed interactions of mesh biomaterials with adhesions, their pathogenesis, and their prophylaxis. This information will ultimately help
33 guide and motivate the intelligent design of future mesh products for improved performance and adhesion prevention.
Several efforts are made here to best achieve this objective. A comprehensive overview is presented to provide an up-to-date understanding of peritoneal post-surgical adhesions, and their general significance (Section 1.3), current treatment methods (which are suboptimal) (Section 1.4), formation (Section 1.5), and prevention (Section 1.6).
Within this broader scope and context, the relationships between peritoneal post-surgical adhesions and implanted surgical mesh biomaterials are examined, and factors are identified which influence adhesion formation (Section 1.5.4) and prevention (Section
1.6.2) specifically in the presence of a mesh device. Attention is drawn to this underexplored but important research area. Finally, emerging opportunities for progress within the mesh-related adhesions field are considered (Section 1.7).
1.3. Clinical Significance of Peritoneal Adhesions
1.3.1. Complications
Peritoneal adhesion formation is associated with numerous complications and sequelae that adversely impact patient well-being. These include abdominal pain and discomfort, bowel or ureteral obstruction, infertility, increased complexity and morbidity/mortality risk during reoperative abdominopelvic surgeries, and reduced safety/efficacy of peritoneal dialysis or chemotherapy.
Traction, compression, and restriction of normal free movement of viscera by peritoneal adhesions may produce intermittent or chronic abdominopelvic pain. In particular, one study demonstrated that traction applied to filmy adhesions between the
34 peritoneum and a relatively movable structure, such as an ovary, was very likely to produce pain 59. Additionally, nerve fibers have also been found in human peritoneal adhesions 1,60, though it remains unclear whether these can be associated with pain, as no difference in the presence of such nerve fibers was shown among patients with and without pelvic pain
61.
Constriction and distortion of organ structures by peritoneal adhesions can partially or completely impair transit of food, fluids, or gametes. Adhesions represent the single most common cause (being responsible for 37-74%) of small bowel obstructions 17,62–64.
These adhesive small bowel obstructions may even occur over a decade after the adhesion- preceding surgery 18. While less commonly reported, peritoneal adhesions have also led to ureteral obstruction 65. Additionally, infertility commonly follows adhesions of the ovaries or fallopian tubes, as adhesion-related constrictions interfere with the proper transport of ova 66. In one study of women that were infertile, 37% had adhesions, and of these, adhesions were found to be the sole infertility factor in 15% 67. Lysis of adhesions has been shown to improve pregnancy rates among such patients 68.
Peritoneal adhesions lead to an increased riskiness and complexity of subsequent abdominopelvic surgical procedures, whether performed for purposes of revision, treatment of unrelated conditions, adhesiolysis (i.e. surgical division of adhesions)
(Section 1.4.1), or a combination thereof. These effects of adhesions pose danger to patients. For example, adhesions increase the rate of hemorrhagic complications in women during Caesarean delivery 69,70. For surgeons, the presence of adhesions tends to obstruct visibility and distort anatomy. This increases the duration of surgery, as well as the likelihood of iatrogenic (i.e. accidental) injury to structures such as the intestines 71–73,
35 bladder 74–76, liver 73, and blood vessels 70,73,77,78. Lysis of adhesions in reoperative surgeries is associated with a mean increase in procedure time of 15-20 minutes and a 6-20% rate of inadvertent enterotomy (i.e. bowel perforation) 17,71,73,79,80. Such bowel injuries increase the risk of surgical site infection and sepsis 73.
When adhesions are present (e.g. in reoperative procedures), laparoscopic surgery, though typically less invasive, may become more dangerous than laparotomy (i.e. open abdominal surgery) 81. With laparoscopy, adhesions cause iatrogenic injuries more often
82, for example during placement of certain instruments (especially trocars or Veress needles) 72,83, insufflation 84, or division of adhesions 85. Such injuries that occur during laparoscopy typically call for conversion to an open procedure 85,86, and worst of all, might even go unrecognized due to limited field of vision 72. Such undetected injuries in laparoscopy can be disastrous. For instance, unrecognized bowel perforation is associated with a 20-50% mortality rate, which is considerably steeper than that for patients sustaining immediately recognized enterotomy 87.
Apart from increased complexity and riskiness of surgeries, peritoneal adhesions can also be a relative contraindication to peritoneal dialysis or chemotherapy. For the former, adhesions may lead to an inability to safely perform catheterization and increase the likelihood of catheter malfunction 88, while decreasing peritoneal capacity and clearance efficiency 89,90. Due to these adhesion-related risks in patients with a history of abdominal surgery, laparoscopy becomes necessary to establish peritoneal access and/or perform adhesiolysis 91–93. Similarly, for the latter, adhesions have been suggested to impede access of drugs to tumor cells 94,95 or may promote catheter malfunction 96,97.
36 With regards to meshes, adhesions may indirectly necessitate removal or delayed implantation of devices. Specifically, inadvertent enterotomy may occur as a result of adhesions, and presents a major risk for infection. If contamination occurs in the presence of surgical mesh, it usually requires challenging removal of previously implanted devices
98,99 and delays introduction of new prostheses for several months 85, given the likelihood for microbial colonization. Both scenarios are devastating for patients, as each heightens the risk for hernia recurrence and the need for further surgery 100.
For further information, adverse consequences of peritoneal adhesions have been a subject of several prior reviews 87,101.
1.3.2. Socioeconomic Burden
Apart from their adverse effects on patient health, adhesions place a tremendous toll on healthcare system resources, increasing clinical workload. During any given year in the US between 1988-2006, procedures to treat peritoneal adhesions necessitated an estimated >280k hospitalizations, >800k days of inpatient care, and >$1.1B in hospital and surgeon expenditures 102–105. These numbers have trended upwards over time. During the same period, the average length of hospital stay following primary adhesiolysis was >7 days 104. Large population studies in Scottish medical records have revealed that within
10y following an open abdominal, pelvic, or gynecologic surgery, 4.5-7.3% of all hospital readmissions were directly related to adhesions, and >32% of patients were readmitted for a condition directly or possibly attributed to adhesions 106–108.
Adhesions and subsequent complications are not restricted to adults. Studies have suggested that patients under the age of sixteen have a 1.1-6.2% rate of hospital readmission specifically due to adhesive small bowel obstruction within 5y following an
37 open abdominal surgery 109,110. Additionally, 8.3% of these young patients are readmitted within 4y for any condition directly or possibly attributed to adhesions 111. Given that pediatric patients generally have many years of physical growth and life ahead of them, this may elevate their risk of suffering adhesion-related complications in the long term.
Concerning mesh-related adhesions, 6.2% and 13.6% of Medicare beneficiary patients are re-hospitalized for adhesion-related complications within 2y and 5y, respectively, following ventral hernia repair 112. Unfortunately, it is not currently possible to generalize adhesion characteristics of surgical meshes in humans given a paucity of available data (Section 1.2.2), and wide variability in material and structural designs for different mesh devices (Section 1.5.4).
1.3.3. Insufficient Awareness
Despite evidence to date that adhesions plague patients and physicians, several studies have suggested that in general, there may be inadequate recognition and estimation of the risks and consequences of peritoneal adhesions (Section 1.3.1, Section 1.3.2) among both of these groups 16,113,114. Additionally, many surgeons may have uncertainties about various measures used to reduce adhesion formation 113,115, especially given a wide variety of prophylaxis strategies (Section 1.6) and lack of definitive information to guide ultimate selection among these. Altogether, however, insufficient awareness of the real burden of adhesions and uncertainty regarding prophylaxis strategies may hinder progress in preventing adhesions and their many potentially serious complications.
38 1.4. Clinical Treatment of Peritoneal Adhesions
1.4.1. Surgical Adhesiolysis
The predominant treatment method for peritoneal adhesions once they have formed is adhesiolysis (also known as adhesiotomy). This entails surgical visualization and dissection to physically disrupt the adhesions. However, adhesiolysis is far from ideal. The procedure is not only tedious and risky, but new adhesions may form or lysed ones may return 116–120. These are referred to as Type 1 (de novo) or Type 2 (reformed) adhesions, respectively 121. In humans, incidence rates of 12-58% and 97% by patient have been reported for Type 1 and Type 2 adhesions, respectively, following laparoscopic adhesiolysis 116,119,120. The important paradox of adhesiolysis is that it relies on surgically- induced damage, the typical cause of most adhesions, to treat them. Given its minimally invasive nature, laparoscopic adhesiolysis reduces the rate of adhesion recurrence compared to open adhesiolysis 117,118,122. However, the former may come with a higher danger of iatrogenic injury (Section 1.3.1). Furthermore, a considerable body of evidence indicates that adhesiolysis lacks efficacy in addressing abdominal pain attributed to adhesions 123–127.
1.4.2. Physiotherapy
Although adhesiolysis remains the current standard in practice and most frequently described treatment for peritoneal adhesions once formed, a growing body of evidence suggests that physical therapy may be able to attenuate the severity of complications related to such post-surgical adhesions 128. A systematic review on physiotherapy-based treatment of peritoneal adhesions concluded that existing evidence, though limited by few detailed
39 reports, small sample sizes, and lack of rigorous or consistent methodology, suggests that mobilization targeting post-surgical adhesions improves outcomes related to pain and dysfunction, with virtually no risk of treatment-related adverse effects 129. Publications from one research group in particular report that soft tissue manipulations may successfully ameliorate peritoneal adhesion-related complications, such as dyspareunia (i.e. painful intercourse) and dysmenorrhea (i.e. painful menstruation) 130–132, infertility 133–135, and bowel obstruction 136–139, thus providing a lower-risk alternative to subsequent surgery 140.
Case studies 128,141 and small trials 142 from other groups have similarly suggested relief from peritoneal adhesion-related pain and dysfunction following physical therapy. Finally, studies in rats demonstrated that scaled-down physiotherapy mobilizations in the early period after a surgical insult were able to attenuate 143 or even physically lyse 144 adhesions.
While encouraging, the published evidence on physiotherapy-based abdominal adhesion treatment presently remains sparse, and human trials with large patient sample sizes are needed 129.
1.5. Pathogenesis of Peritoneal Adhesions and Relation to Mesh Biomaterials
1.5.1. The Healthy Peritoneum
A thorough comprehension of the normal peritoneum physiology is necessary to understand the pathomechanisms behind formation of peritoneal adhesions, and to inspire strategies for their prevention. The peritoneum is the most extensive serous membrane in the body 145. It covers the surfaces of the abdominal cavity (i.e. parietal peritoneum) and the organs within it (i.e. visceral peritoneum). The peritoneum is composed of a specialized epithelial cell monolayer ~2.5-3 µm thick that is called the mesothelium, and an underlying
40 connective tissue layer 146. The two layers are separated at their interface by a basement membrane 146, an acellular matrix. The basement membrane is ~40 nm thick 147 and consists of laminin, collagen IV, fibronectin, and other matrix proteins 148–150. Adjacent mesothelial cells are linked together via desmosomes, and tight junctions at their apical surface, forming a continuous barrier 146. The subserosal connective tissue layer contains extracellular matrix, blood/lymph vessels, and cells of mesenchymal origin (e.g. fibroblasts) 145,151,152.
As the outermost layer, peritoneal mesothelial cells play important roles in maintaining peritoneal homeostasis. Two main functions traditionally ascribed to the mesothelium include physically hindering entry of pathogens and minimizing friction between apposing tissues and/or organs to promote their free movement 152,153. In addition to secreting free surfactant molecules (e.g. phospholipids) as a lubricating fluid 154–156, peritoneal mesothelial cells synthesize a matrix of carbohydrate-rich biomolecules such as glycolipids, glycosaminoglycans (e.g. hyaluronan), and proteoglycans, known as the glycocalyx, which is anchored to their apical surface 157. The polysaccharides in the glycocalyx entrap water molecules to create a slippery, non-adhesive surface that coats the mesothelium. This protects the underlying cells against abrasion. The glycocalyx also functions to limit non-specific protein adsorption 158,159. Together, the mesothelium and glycocalyx provide a natural defense against adhesions. Further, the glycocalyx could represent the ideal adhesion barrier - one that is native to the body and continuously renewed by mesothelial cells.
More recently, a large body of evidence has surmounted that peritoneal mesothelial cells also play active roles in fluid and solute transport, and regulation of inflammation and
41 wound healing (e.g. through synthesis of cytokines, growth factors, and matrix proteins)
153,157. Importantly, mesothelial cells can exert both procoagulant and fibrinolytic effects within the peritoneal cavity, and the complex and dynamic balance between these two is key in the formation or prevention of peritoneal adhesions. After a disturbance to peritoneal tissue (e.g. surgery), fibrin may accumulate on tissue surfaces. Fibrin deposits that accumulate and bridge between tissue surfaces (i.e. fibrinous adhesions), serve as the basis for permanent adhesions (i.e. fibrous adhesions) if allowed to persist and mature 160.
Fibrous adhesions are remarkably durable, both biologically and mechanically, and are generally only disrupted using adhesiolysis (Section 1.4.1).
Thus, inadequate fibrinolytic activity (i.e. hypofibrinolysis) in the early period following such a disturbance is largely responsible for adhesion formation 161. The fibrinolytic activity of mesothelial cells derives primarily from their secretion of tissue plasminogen activator (tPA), a protease that converts inactive plasminogen into active plasmin, the latter of which then enzymatically degrades fibrin. Mesothelial cells are the predominant source of tPA in the peritoneum. Adding another layer of complexity, upregulation of inflammatory factors tends to shift the balance in favor of coagulation.
Inflammatory factors decrease mesothelial cell production of tPA, while increasing synthesis of plasminogen activator inhibitor (PAI)-1, which directly inhibits tPA, thereby causing a marked reduction in fibrinolysis.
1.5.2. Adhesion-Free Peritoneal Healing
Under normal conditions, maintenance of peritoneal homeostasis is straightforward as disruption is atypical. However, this balance is fragile and can be very easily thrown into disorder through traumatic events, such as surgery, thus potentiating adhesion
42 formation. However, under ideal circumstances, adhesion-free restoration of peritoneal homeostasis is possible.
Peritoneal healing without adhesions requires timely restoration of damaged mesothelial surfaces. In theory, this must ideally occur either before: 1) Fibrin-containing fluids can become deposited and firmly anchored to a site of injury. 2) Another tissue surface can come into contact with non-polymerized fibrin at the injury site. 3) An apposed tissue surface can sustain increased friction from the injured tissue, thus generating a secondary site of injury at which the healing process starts anew. Or 4) bridging fibrinous adhesions can be infiltrated by fibroblasts and remodeled into fibrous adhesion tissue.
Fortunately, unlike in healing of skin, the injured peritoneal surface becomes re- epithelialized simultaneously/diffusely rather than gradually from the borders; as a result, large wounds are found to heal about as rapidly as smaller ones 162,163. Evidence suggests that this may be due in part to viable free-floating mesothelial cells in the peritoneal fluid that engraft as islands onto the injured surface 151. Together with mesothelial cells at the wound periphery that break cell-cell contacts and migrate onto the wound surface, these cells proliferate and reestablish a continuous mesothelial layer 151. Studies generally agree that complete remesothelialization of the surface occurs in approximately 5-8 days 163–166.
This process of remesothelialization has some important implications for meshes implanted into the peritoneum. Namely, in the absence of adhesion formation (similar to an injured tissue surface), the visceral aspect of a mesh can be quickly covered with a new mesothelial layer (i.e. neoperitoneum) 165,166 that should serve to deter future adhesions.
43 1.5.3. Abnormal Peritoneal Healing and Adhesion Formation
Unfortunately, peritoneal healing does not always take place before adhesions can form. Propensity to develop peritoneal adhesions can differ greatly between individuals 167.
The root of this variability is that adhesion formation is an enormously complex process.
It is regulated simultaneously by intricate biological pathways of inflammation, coagulation, and tissue repair, as well as the crosstalk between them. This immense complexity makes it inherently difficult to control for all variables in studies on peritoneal adhesions.
Regardless of complexity, detailed appreciation of adhesion formation is critical to any intervention/prevention strategy. Current understanding posits that adhesion formation is dependent on (or influenced by) several key factors, some of which may be needed in combination. These factors are described in the following subsections. For further information on the biology of peritoneal repair and adhesion formation, readers are directed to additional reviews 168.
1.5.3.1. Tissue Disturbance
Peritoneal adhesion formation generally requires some type of inducing disturbance to a peritoneal tissue surface. One such mode of disturbance is surgical handling/trauma, which may include mechanical or thermal injuries that abrade, denude, or destroy the slick mesothelial surfaces. Another example is tissue desiccation 169, which may occur during surgery and disrupt the ability of the glycocalyx to provide lubrication, protection, and resistance to protein adsorption to the underlying mesothelium. Such perturbations may enhance accumulation of intraperitoneal fibrin (Section 1.5.3.2), increase adherence of fibrin deposits to tissue surfaces, reduce the ability of apposed surfaces to slide across one
44 another without sticking (Section 1.5.3.3), prolong the contact between apposed surfaces
(Section 1.5.3.4), facilitate fibroblast access to fibrin deposits, and incite inflammation
(Section 1.5.3.5). Additionally, different tissues may exhibit disparate propensities for adhesion formation following a disturbance; some evidence suggests a higher susceptibility of visceral peritoneum than parietal peritoneum 170,171.
1.5.3.2. Fibrin Deposition
Deposition of fibrin near injured peritoneal surfaces seems to be essential in initiation of adhesion formation 169. Fibrin deposits may arise either from bleeding after vessel damage as a result of surgical trauma, or from inflammatory signals produced by the disturbed tissue that trigger vasodilation and leakage of exudate (i.e. protein- and cell- containing fluid discharged from the damaged tissue due to increased vascular permeability) in the hours following surgery 172. One such mechanism of inflammation- induced vasodilation involves promotion of mast cell degranulation and histamine release
173–175. In addition to fibrin, platelets in blood or exudate may adhere and become activated upon contact with exposed collagen (and/or von Willebrand Factor) at injured tissue surfaces, thus accelerating fibrin anchorage and polymerization.
1.5.3.3. Proximity of Surfaces
Another element that appears critical for adhesion formation is the close proximity of at least two apposed surfaces, upon which initial coagulation products (and ultimately fibrous adhesions) can firmly anchor and bridge across. These can be adjacent tissues, organs, or implanted biomaterials. Evidence suggests that just one injured surface is sufficient to produce adhesions, possibly by increasing friction on the apposed tissue
45 surface 176. However, the presence of two injured surfaces tends to be more likely to promote adhesion formation 177. Adhesions form more reliably when the apposed surfaces remain in close approximation with one another following the inducing disturbance for a prolonged period of time (Section 1.5.3.4). Conversely, adequate and sustained physical separation of involved tissue surfaces during the early post-operative healing period using adhesion barrier materials (Section 1.6.1.3), or even instilled gas 178, may effectively mitigate peritoneal adhesion formation.
1.5.3.4. Time for Adhesion Initiation and Maturation
Time is necessary to permit adhesion formation, which comprises two phases.
Initiation is the first phase, in which bridging fibrinous adhesions are established between surfaces. The subsequent phase is maturation, in which fibrinous adhesions are remodeled to become permanent fibrous adhesions. Initiation is the less time-intensive of the two processes. Maturation involves fibroblast migration into the deposited fibrinous scaffold, then reorganization of this provisional matrix through resorption of fibrin and concomitant deposition of extracellular matrix components, particularly collagen. At the same time, angiogenesis occurs and the maturing adhesion becomes vascularized. Overall, this remodeling process increases the mechanical and biological stability of the adhesion tissue.
As evidence of this, increased relative motion between apposed surfaces in the first few days following an inducing disturbance decreases adhesion formation 143,144,179–181, and conversely, decreased relative motion increases adhesion formation 182. Such movement likely weakens or disrupts adhesions while still in the fibrinous state.
The timings of the first and second phases of adhesion formation are ill-defined, as together they represent a continuous process, with no distinct transition between them. As
46 such, the kinetics of either phase are not precisely known. However, prior animal studies have proposed that initiation may occur within the first ~16h, after which point maturation proceeds 183. The transition of an adhesion from a fibrinous to a fibrous state is considered to occur around ~3-4 days after the inducing event 183.
As another important window, the “critical period,” during which adhesion barrier interventions are suggested to be most crucial for adhesion prevention, ranges from 36h 184 to 3-5 days 185 after the initial disturbance. This implies that to maximize effectiveness, barrier materials should persist at their site of application for longer than this critical period, ideally ~1 week at minimum. This duration is in rough agreement with the reported time frames for remesothelialization of peritoneal surfaces (Section 1.5.2), as restoration of mesothelium is a key event in deterring abnormal healing and adhesion formation.
1.5.3.5. Inflammation
Inflammation intensifies adhesion formation by both enhancing release of fibrinous exudate (Section 1.5.3.2), and reducing/delaying peritoneal fibrinolytic activity by locally altering levels of proteins such as tPA and PAI-1 (Section 1.5.1). This shift in favor of fibrin deposition is especially important in the first few hours/days of the triggering event
(Section 1.5.3.4). Inflammation and subsequent adhesion formation can be induced by the initial injury or disturbance, infection 186, endotoxins 187, and/or foreign bodies. Pro- inflammatory foreign bodies can include implanted devices (Section 1.5.4), fixation elements such as sutures 188,189 or tacks 190–193, starch 194–199 or talc 2,200–202 from surgical gloves, and fiber fragments or lint from gauze or sponges 197,203,204. Inflammation is also regulated at the cellular level, for example by macrophages, polarization of which (e.g. so- called M1 vs M2 phenotypes) has been suggested to impact adhesion formation 205–207.
47 1.5.3.6. Ischemia and Hypoxia
Ischemia and subsequent tissue hypoxia aggravate adhesion formation by promoting fibrin deposition, fibrosis, and inflammation, and causing dysregulation of fibroblast proliferation and apoptosis. Examples of events which have been proposed to trigger ischemia/hypoxia and subsequent adhesions include tissue strangulation following
186,208 suture-based defect closure , and insufflation with CO2 or helium gas during pneumoperitoneum in laparoscopy 209–211. Hypoxia promotes fibrin deposition by downregulating fibroblast expression of tPA and increasing PAI-1 expression 212. At the same time, it stimulates fibrosis by increasing fibroblast expression of transforming growth factor (TGF)-β1 and subsequent collagen synthesis 213–215. Hypoxic conditions incite inflammation (Section 1.5.3.5) through peritoneal fibroblast upregulation of interleukin-6 and tumor necrosis factor ɑ 216, which are pro-inflammatory cytokines, and of cyclooxygenase-2 217, expression of which triggers elevated production of pro- inflammatory prostaglandins. Finally, hypoxia preferentially increases proliferation and suppresses apoptosis of fibroblasts derived from adhesion tissue relative to normal peritoneal fibroblasts 218,219.
1.5.4. The Role of Surgical Meshes in Peritoneal Adhesion Formation
While surgical meshes and adhesions have frequently been described together, details regarding processes by which these prostheses influence adhesion pathogenesis seem to be unclear, elusive, or incomplete, and not all plausible mechanisms have been described or considered in literature. Consequently, many reports may be missing key aspects when interpreting adhesion formation in relation to mesh devices. Further, some studies that observe mesh-stimulated adhesiogenesis may offer little explanation apart from
48 attributing this finding to meshes acting as foreign bodies. This generalization is unsatisfying as it (falsely) implies that material selection and product design have no place in adhesion mitigation. Fortunately, broader research on biomaterials has established principles that may supply new insights concerning detailed mechanisms of mesh adhesion stimulation.
Meshes may impact adhesion formation through both their material properties
(including surface and bulk properties), and their structural characteristics. Material properties can modulate inflammation and nucleation of fibrin deposits, while structural characteristics can scale the inflammatory response, promote stability of adhesion anchorage, and potentiate injury to apposed tissue surfaces.
1.5.4.1. Effects of Material Properties
Material properties describe qualities of a substance that do not depend on its amount or spatial configuration. Examples include chemical composition, elastic modulus, and density. Material properties can be further subdivided into those unique to the surface versus the bulk. Mesh material properties can exacerbate adhesion formation through enhancement of the inflammatory response (Section 1.5.3.5), and through nucleation of fibrin deposits. In particular, the surface properties of a mesh material may influence adhesion formation through both of these mechanisms, while bulk properties only play a role in fibrin nucleation.
Let us consider PP, the most common surgical mesh implant material 46–50. Though
PP meshes possess advantages of rapid incorporation, biodegradation resistance, suitable strength, ease of use, and affordability, they are also notoriously adhesiogenic. Many studies have demonstrated the propensity of PP meshes to incite strong and persistent
49 inflammatory responses 220,221. A probable factor in this pro-inflammatory nature is that when surfaces of hydrophobic materials, such as PP, come into contact with proteinaceous aqueous solutions (e.g. body fluids), they have a strong tendency to adsorb proteins and denature (i.e. conformationally alter) them 222–226.
Within milliseconds after contacting a body fluid, the surface of an implanted material becomes covered in proteins, which adsorb (i.e. bind/adhere as molecules onto the surface). The biological response to such a biomaterial is then defined by the surrounding corona of adsorbed proteins, in terms of their identities, surface concentrations, and conformations. This adsorption process is inherently complex. The makeup of the protein conditioning film at any given moment represents a dynamic equilibrium influenced by the solution concentration, diffusivity, and affinity. Over short time frames, small and abundant proteins predominate on the surface, and gradually, higher-affinity proteins replace them. This process is known as the Vroman effect.
Ordinarily, protein solutes in physiologic fluid exist with polar and charged amino acids mostly localized to the exterior. Nonpolar segments are predominantly buried in the core, being shielded from interactions with the polar solvent 227,228. At nonpolar surfaces, water molecules are easily displaced by proteins. The proteins are then entropically driven to unfold, with core nonpolar domains becoming exposed to the surface where they become tightly associated through hydrophobic interactions 224. Vroman is credited with first recognizing this process 228.
Such denatured proteins at hydrophobic mesh implant surfaces may be misrecognized by the host immune system and elicit inflammatory responses. In particular, fibrinogen, found in blood and fibrinous exudate, is highly pro-inflammatory once
50 adsorbed and denatured 229–233. The increased inflammation pursuant to protein/fibrinogen denaturation would contribute to reduced fibrinolytic activity within the peritoneal cavity, thus exacerbating postoperative adhesion formation.
Furthermore, adsorbed and denatured fibrinogen on hydrophobic mesh implant surfaces may also directly participate in formation of fibrin scaffolds which ultimately mature into adhesions. Studies have demonstrated that fibrinogen denatured on hydrophobic surfaces may nucleate polymerization into long strands of fibrin 234, which may trigger adherence, activation, and aggregation of platelets 235 that further accelerate coagulation. Additionally, adsorption and denaturation of other clotting factors on the mesh surface may further promote coagulation.
Conversely, more hydrophilic surfaces, such as those of mesh hydrogel-based adhesion barriers (e.g. on Sepramesh), are much less likely to induce protein denaturation, thus mitigating pronounced inflammatory responses. Mesh hydrogel-based adhesion barriers can thus attenuate the inflammatory response by directly concealing a large fraction of the hydrophobic mesh surface, limiting its extent of protein interactions.
Additionally, proteins adsorb less tenaciously to hydrophilic surfaces. Finally, hydration of such hydrogel-based adhesion barriers makes them much softer (i.e. lower elastic modulus - a bulk property) than the rigid polymers which comprise hydrophobic meshes.
This may mechanically weaken the anchorage of fibrin deposits and subsequent adhesions.
Note that hydrophobicity cannot by itself explain adhesion formation to meshes.
For example, meshes composed of polytetrafluoroethylene (PTFE), including expanded
PTFE (ePTFE - a microporous form of PTFE), are hydrophobic like PP, yet are reportedly much less adhesiogenic 236,237. For one, the protein conditioning layer on ePTFE surfaces
51 may be vastly different than that on PP (e.g. less pro-inflammatory or pro-coagulatory). In support of this, lower-grade inflammatory responses have been seen with ePTFE than with
PP meshes 238, and PTFE surfaces have been shown to denature and inactivate thrombin
239, which otherwise promotes fibrin polymerization. Secondly, other differences between ePTFE and PP devices, such as their structural characteristics (Section 1.5.4.2), may be even more important factors in adhesion formation.
For further information on host-mesh surface interactions, readers are directed to the following review 240.
1.5.4.2. Effects of Structural Characteristics
Structural characteristics describe qualities of a material that are dependent on its amount or spatial configuration. Examples include surface area, texture, and stiffness.
Mesh structural characteristics may impact adhesion formation through scaling of the inflammatory response, enhancing adhesion anchorage, and/or potentiating tissue injury.
Terms commonly used to describe the structure of mesh components include reticular vs laminar, and macroporous vs microporous. A reticular mesh structure is made up of filaments/fibers that are knitted or woven together, producing a net-like architecture with an irregular/rough texture at tissue-level dimensions, whereas a laminar structure is constructed from a single sheet of material, resulting in a smoother and more continuous topography at the same length scale 237,241–244. In terms of mesh pore structure, macroporous and microporous are typically accepted to refer to pore diameters >75 μm (all dimensions) and <10 μm (at least one dimension), respectively 245–247. Note that a reticular mesh must be macroporous (knitting and weaving processes typically do not produce pores < 10 μm), that a single structure can be simultaneously macroporous and microporous at different
52 length scales (e.g. reticular ePTFE) 242, that a laminar material can be macroporous (e.g. perforations in the structure, or the parietal aspect of ePTFE Dualmesh) 237, and that it is also possible for certain mesh components to be essentially nonporous (e.g. hydrogel barriers). Among single-material prostheses, the visceral side of most current PP mesh devices is reticular and macroporous, while it tends to be laminar and microporous for ePTFE devices 237.
The structure of a mesh prosthesis, in particular its surface area, may dictate the extent of inflammation following its implantation, thus impacting adhesion formation.
Considering the hydrophobic PP mesh example from before (Section 1.5.4.1), the surface area of exposed mesh is proportional to the total amount of denatured protein on the mesh surface. These denatured proteins may commensurately impact the subsequent degree of inflammation in the peritoneal cavity. The surface area of a prosthesis can be influenced by the diameter and number per unit area of filaments (in the case of a reticular structure), and the total size of the implant required to close the defect. In many cases, despite an abundance of large pores, a reticular structure will possess a higher overall surface area than a laminar one.
Next, the structure of a mesh may influence the anchorage stability of adhesions. A reticular architecture facilitates strong anchorage of adhesion tissue. The connective tissue matrix not only can attach to the mesh visceral surface, but it can also wrap around or completely envelop the mesh filaments, and connect to the underlying tissue exposed by the open pores. Furthermore, this type of structure is permeable to exudate that may traverse directly from the parietal side to the mesh visceral surface, thus nucleating adhesions between the mesh and viscera. On the other hand, a laminar topography provides
53 only a relatively flat surface, making firm adhesion anchorage less likely. Because this type of structure is also impermeable to fluids, exudate trapped between the parietal peritoneum and the mesh can only move off to the side of the mesh rather than directly onto the mesh visceral surface, leading to a greater proportion of adhesions seen at the mesh border 248.
Finally, the structure of a mesh might even impact the propensity of the prosthesis to inflict tissue injury through abrasion of the apposed tissue surface. The coarse surface of a reticular mesh may be more likely to inflict tissue damage to the apposed surface than the smooth surface of a laminar structure. In support of this view, fistulization, erosion, and extrusion have been described far more commonly for PP meshes 249–256 than for ePTFE meshes 257,258. Each of these complications involves a mesh wearing completely through a layer of soft tissue (usually into the lumen of a hollow organ) over time, likely as a result of sustained or repeated pressure/abrasion. A causal relationship between mesh texture and such complications has never been proven - though it has been suggested 259.
However, an argument could be made that if a mesh could inflict these complications through abrasion of tissue surfaces, it should be similarly capable of inducing tissue injuries that trigger adhesions.
The increased propensity of reticular structures to provide a firm anchorage site and inflict damage to an apposed tissue surface may partly explain the higher incidence of tissue adhesion generally reported in literature for PP mesh relative to laminar ePTFE meshes 237. In support of this, reticular ePTFE mesh has demonstrated a greater number of adhesions, and adhesions that required more traction force to separate, than laminar ePTFE mesh 242. Likewise, PP mesh possessing a more laminar structure has demonstrated less severe and more easily separable adhesions than reticular PP mesh 260.
54 Looking beyond adhesions, there are important drawbacks of microporous mesh structures. For example, the architecture of ePTFE makes this material highly susceptible to permanent microbial colonization and biofilm formation, reducing the chance for successful eradication of device infection, and increasing the need for mesh explantation if infection occurs 261,262. Bacteria can occupy the microscopic pores in ePTFE where they are protected indefinitely, as the material never degrades in vivo and host immune cells are too large to pursue them. Additionally, tissue integration with the abdominal wall tends to be weaker for microporous ePTFE than for macroporous PP 260. Finally, low water permeability may make meshes with a non-macroporous ePTFE layer relatively more susceptible to formation of a seroma (i.e. local accumulation of serous liquid), as fluids are not readily cleared through the microscopic pores of the material 263,264.
55 1.6. Prophylaxis of Peritoneal Adhesions
1.6.1. General Preventive Strategies
Prevention of adhesions is always a desirable outcome. Given the many processes that contribute to their formation, a multitude of strategies have been applied to combat adhesions. General preventive strategies for peritoneal adhesions can be categorized as control of surgery-related factors, and use of adjuncts such as drugs or barrier materials, each explored in sections below.
1.6.1.1. Surgery-Related Factors - Techniques and Instrumentation
Skillful control of surgery-related factors, which include techniques and instrumentation, can limit adhesions through minimization of several of the elements that promote adhesion formation (Section 1.5.3). Surgical approaches reported to limit adhesion formation tend to fall in general agreement with Halstedian principles (named after the surgeon William Halsted). However, even optimal practices exercised by the most experienced surgeons are unlikely to prevent adhesions entirely on their own 265,266. At the same time, the effectiveness for a number of these strategies is controversial. Surgery- related techniques (Table 1-1) and instrumentation (Table 1-2) that have been reported to impact adhesion formation are summarized below, along with their supposed mechanisms, literature evidence to support or discredit their use, and other primary or secondary sources that explore their relation to adhesions.
56 Table 1-1. Effects of Surgery-Related Techniques on Peritoneal Adhesions Evidence: Evidence: Surgical Practice for Supposed Explanation / Theory for Adhesion See Adhesion Lack of Mitigating Adhesions Prophylaxis Also Reduction Reduction
Laparoscopy / Laparoscopy may reduce trauma, bleeding, inflammation, Minimally Invasive and inadvertent introduction of bacteria or foreign 267,268 269 270 Surgery material, as compared to laparotomy.
Microsurgery may reduce trauma, bleeding, and 272, Microsurgery 271 inflammation as compared to macrosurgery. 273
Minimizing peritoneal introduction of bacteria, feces, and 186,187, Maintaining Asepsis 275 endotoxins may mitigate infection and inflammation. 274
Reducing or altering bacterial flora in the bowel prior to Pre-Surgical Bowel surgeries, especially those involving possible leakage of 276–278 Preparation / Cleansing bowel contents (e.g. anastomosis), may mitigate infection and inflammation.
Handling of tissue in a gentle way and minimization of Gentle Tissue Handling unnecessary tissue handling (as comes with surgical 279–283 experience) may reduce trauma and inflammation.
Minimizing Duration of Decreasing duration of insufflation or surgery may 209,211, 285 Insufflation or Surgery minimize desiccation and ischemia. 284
Minimizing Pressure of Decreasing insufflation pressure may minimize 211,285 286 Insufflation desiccation and ischemia.
Decreasing Suture Minimizing tension on tissue may reduce ischemia. 287 Tightness
Leaving behind long suture ends on knots may increase Trimming of Suture abrasion of apposed tissue, inflammation, and adhesion 288 Ends from Knots anchorage to knots.
Fewer (e.g. via use of continuous vs interrupted suture) or Suture/Knot 171,290, smaller knots may facilitate sliding of apposed tissue 287,289 Configuration 291 surfaces and minimize adhesion anchorage to knots.
Pre- or Intraoperative Hydration may minimize desiccation, and irrigation may Hydration of Tissues flush away bacteria or foreign debris. However, 292 293 169 with Crystalloid crystalloid solutions could dilute/disrupt the glycocalyx. Solutions
Preoperative Coating of Coating tissues with hydrophilic polymers (e.g. Tissue Surfaces with hyaluronic acid) before manipulation may minimize 294–296 Hydrophilic Polymers trauma and desiccation by augmenting the glycocalyx.
Postoperative Peritoneal Instilled volumes may reverse desiccation and physically 269,298, Instillation of separate tissue surfaces. However, crystalloid solutions 292,297 300 299 Crystalloid Solutions are quickly absorbed in <24h after instilling.
Maintaining Meticulous Reducing the amount of blood in the peritoneal cavity 2,169,301,
Hemostasis may minimize fibrin deposition. 302
Closure of peritoneal defects may promote ischemia and 186,291, Omittance of Peritoneal 171,186, 308– inflammation (e.g. due to suture material) as compared to 303,305– Closure 303,304 310 non-closure. 307
57 Table 1-2. Impact of Common Surgical Instrumentation on Peritoneal Adhesions Instrumentation Evidence: Evidence: Supposed Explanation / Theory for See Strategy for Mitigating Adhesion Lack of Adhesion Prophylaxis Also Adhesions Reduction Reduction
Hemostatic agents stem bleeding in Utilization of Hemostatic 320– the peritoneum and may minimize 311–316 317–319 Agents 323 fibrin deposition.
Fiber fragments left behind from Avoidance of Lint 197,203, gauze or surgical sponges may Material 204 promote inflammatory responses.
Avoidance of Starch- Starch powder from gloves may be left 194–198
Powdered Gloves behind and promote inflammation.
Avoidance of Talc- Talcum powder from gloves may be 2,201,202 Powdered Gloves left behind and promote inflammation.
Altering Temperature / Warming and humidifying insufflation Humidity of Insufflation gas to physiologic conditions may 324,325 326 Gas mitigate tissue desiccation.
Barbed sutures may promote adhesion Avoidance of Sutures with formation compared to non-barbed 327 328 Barbed Structures sutures.
186,274, Appropriate Selection of Certain suture materials may stimulate 291,303, 289,290, 189 Suture Material inflammation more than others. 330 329
Reducing suture diameter may reduce Use of Smaller Diameter 189,288, inflammation and ability of adhesions 271 Sutures 290 to firmly anchor to the sutures.
Adding a small amount of O or N O 2 2 209,211, Altering Insufflation Gas to CO insufflation gas may reduce 2 284,301, 286 Composition tissue hypoxia or inflammation, 331 respectively.
Use of Less Abrasive Decreased abrasiveness may reduce 332 Gauze trauma.
Minimizing the number of trocars may Reduction in Number of reduce trauma, bleeding, and 333 Trocars in Laparoscopy inflammation.
Application of Laser, Advanced dissection / coagulation Harmonic, Waterjet, or devices may reduce trauma and 335– 271,334 Electrosurgical Dissection bleeding as compared to conventional 338 / Coagulation Devices scalpels / hemostats.
58 1.6.1.2. Pharmacologic Strategies
Numerous drugs have been experimentally investigated toward application for prevention of post-surgical adhesions. Given that most (if not all) drugs lack specific indication for post-surgical adhesion prevention, these pharmacologic strategies tend to be more frequently explored in pre-clinical rather than clinical scenarios. In Table 1-3, drug compound classes, their putative mechanisms, and specific drugs within those classes which have been investigated for adhesion prophylaxis are summarized. This table is not fully comprehensive of all investigated drug classes or compounds; only pure drug compounds for which there is currently a simple, clear scientific premise for adhesion prophylaxis were included. Most of the represented drug classes limit adhesions through interference with fibrin deposition (e.g. anti-histamines, anti-coagulants, fibrinolytics), adhesion maturation (e.g. prokinetics, anti-proliferatives, anti-fibrotics), inflammation (e.g. anti-inflammatories, anti-microbials), hypoxia (e.g. anti-oxidants), or a combination thereof. Other pure drug compounds investigated for adhesion prophylaxis, currently for which there is either a relatively convoluted or potentially unclear theoretical basis, are listed in Table 1-4.
59 Table 1-3. Effects of Pharmacologic Agents on Peritoneal Adhesion Formation Drug Classes for Supposed Explanation / Theory for Mitigating Literature Examples Adhesion Prophylaxis Adhesions
Anti-histamines, Anti-histamines: Dimetindene maleate339, Limit downstream effects or initial Mast Cell Diphenhydramine340,341, Promethazine175,342 release of histamine, thus reducing Stabilizers, and Mast Cell Stabilizers: Cromoglycate173, vascular permeability and exudation of Histamine Receptor Tranilast343,344 fibrin. Antagonists H1 Receptor Antagonists: Ketanserin345
Inhibit polymerization and deposition Anti-coagulants Citrate346,347, Dicumarol348, Heparin349–353 of fibrin (e.g. on surfaces).
Ancrod354, Mesna355, Neurokinin-1 receptor Fibrinolytics or Promote degradation of polymeric antagonist356,357, Reteplase358,359, Fibrinogenolytics fibrin and/or of monomeric fibrinogen. Streptokinase360,361, tPA360,362,363, Urokinase plasminogen activator360
Promote intestinal motility to increase relative motion between peritoneal Prokinetics Cisapride180, Escin179 surfaces in the early postoperative period.
Inhibit proliferation of cells (e.g. Anti-proliferatives Bevacizumab364, Mitomycin C365,366, Paclitaxel367, fibroblasts) or growth of blood vessels or Anti-angiogenics Sunitinib368 involved in adhesion maturation.
Reduce synthesis/deposition of Antibodies to TGFꞵ369,370, Halofuginone371, Anti-fibrotics collagen, limiting adhesion maturation. Collagenase372
Steroids: Budesonide373,374, Dexamethasone172,342,375–377, Hydrocortisone378,379, Methylprednisolone340 Non-selective NSAIDs: Aspirin380, Ibuprofen381– Limit inflammation by targeting 383, Indomethacin384, Oxyphenbutazone385,386, Anti-inflammatories various inflammatory pathway Piroxicam387, Tolmetin381,388,389, Tenoxicam390 components. Selective NSAIDs: Celecoxib384,391, Meclofenemate392, Nimesulide384,393, Parecoxib394, Rofecoxib395 Cytokines: Interleukin-10396,397 Macrophage Polarizers: Pioglitazone205, YC-1206
Limit infection and subsequent Cefepime398,399, Metronidazole398,399, Anti-microbials inflammation. Sulfanilamide400, Taurolidine401,402
Reduce tissue hypoxia by supporting Anti-hypoxants oxygen delivery or carbon dioxide Protoporphyrin403,404 removal for cells and tissues.
Allopurinol405, Berberine406, Catalase407, Directly scavenge or indirectly reduce Ligustrazine408, Melatonin409,410, Methylene formation of reactive oxygen species Anti-oxidants blue411,412, Resveratrol413,414, Superoxide that promote oxidative stress and dismutase415, Tirilazad416, Trimetazidine417, subsequent inflammation. Vitamin C418, Vitamin E419–422
Multiple Combination of other mechanisms in Activated Protein C423, Aescinate424, Iloprost425, Mechanisms parallel. Sirolimus426,427
60 Table 1-4. Other Pharmacologic Agents Studied for Peritoneal Adhesion Prevention Drug Class Literature Examples
Bupivacaine 428, Lidocaine 428,429, Prilocaine 428,429, Anesthetics / Analgesics Propofol 430
Anti-fibrinolytics Aprotinin 376,431
Calcium Channel Blockers Bepridil 432, Verapamil 433,434
Cytokines Interferon-Gamma 435,436
Growth Factors Epidermal Growth Factor 437
Histone Deacetylase Valproate 438,439 Inhibitors
Hormones Estrogen 440, Ghrelin 441, Progesterone 442,443
Immuno-stimulants Plerixafor 444
Immuno-suppressants Cyclosporine 445, Pimecrolimus 446, Tacrolimus 444
Proteases Papain 447
Cilostazol 448, Pentoxifylline 448, Rolipram 449, Phosphodiesterase Inhibitors Sildenafil 450
Serotonin Synthesis P-Chlorophenylalanine 451 Inhibitors
Statins Atorvastatin 452, Lovastatin 453, Simvastatin 452,454,455
Toxins Botox 456
61 Note that these lists do not consider gene 457,458 or gene-silencing 459 therapies, cell therapies 460–462, or experimental agents which may impact peritoneal adhesions through barrier (rather than pharmacologic) functions, such as phospholipids 302,463–466, honey 467, trehalose 468,469, fucoidan 470,471, garlic oil 472, canola oil 473, soybean oil 474,475, olive oil 420, or aloe vera gel 476. Also not considered are mixtures that don’t represent a single pure compound, such as cell-conditioned media 477,478, platelet rich plasma 353,479, Kombucha
480, Changtong oral liquid 481, or the prokinetic ricinus oil 482.
Additionally, it is important to recognize that delivery route (e.g. local vs systemic), rate, and dosage are also key for determining adhesion outcomes and side effects with pharmacologic strategies. Localized delivery is preferable to systemic delivery, as it reduces the dose required for therapeutic benefit, while limiting adverse effects related to excessive dosage, and off-target effects. As an example highlighted by others previously
434, similar post-surgical adhesion prevention was possible with a 600-fold reduction in the minimum effective dose of ibuprofen when using local intraperitoneal delivery as opposed to systemic administration 383. Control of drug delivery rate also helps to limit deleterious effects associated with “burst release,” a phenomenon in which there is an excessively high drug concentration for a very brief window (promoting adverse effects), followed by sub- therapeutic levels. Drug delivery vehicles are one strategy to improve control over drug localization and release kinetics. As an ideal solution that may and effectively safely maximize adhesion prevention, adhesion barriers could represent such vehicles for controlled local delivery of pharmacologic agents 360,373,374.
For more detailed information on pharmacologic agents for adhesion prevention, readers are directed to additional reviews 483.
62 1.6.1.3. Barrier Materials - Membranes and Gels
Barrier materials limit adhesion formation primarily by providing physical separation between surfaces of internal structures during the critical period of time following a tissue disturbance. Adhesion barriers can be classified as inert membranes, inert gels, and tissue-adherent membranes/gels. Membranes have characteristics of a solid or semisolid material prior to delivery, but may gradually convert to a gel or liquid state afterward. Gels are viscous solutions, typically consisting of hydrophilic polymers suspended in water, that are able to flow prior to delivery. These may thicken immediately after delivery in the event that crosslinking occurs in situ, as is the case for some of the tissue-adherent gels described below. Inert membranes and gels are not intended to react with application sites following delivery, but merely provide passive separation between surfaces. In contrast, tissue-adherent membranes and gels undergo reaction with tissue surfaces in situ to remain in place at the application site following delivery, thereby providing stable interposition between surfaces.
Considering inert membranes versus gels, the former have advantages in that a membrane is not as readily displaced from between adjacent surfaces and tends to have a longer residence time to allow for healing to take place. However, membranes are not as easily delivered, require precise placement at sites where adhesions are anticipated to develop in order to provide benefit, and may need suturing, which can lead to adhesions at suture sites484. On the other hand, gels can be effortlessly instilled in large volumes to provide broad coverage and a large degree of separation through what is frequently described as the “hydroflotation” effect 485. Furthermore, given their ability to flow, gels are not limited to the site of application but can even provide protection in spaces that might
63 otherwise be difficult to access. However, this can also be a drawback, as gels may be expelled from between surfaces that need protection, and excess may flow into tight spaces between adjacent tissues where coalescence is actually desirable (e.g. at an anastomosis or incision). Tissue-adherent barrier materials possess a unique advantage in that they can circumvent the problematic undesired displacement seen in inert membranes and gels. The main downside of a tissue-adherent barrier is that once inadvertently applied at an undesired site, it may be practically impossible to remove.
Many different materials, for each type of adhesion barrier, have been widely investigated for adhesion prophylaxis in both pre-clinical and clinical scenarios. The literature on adhesion barrier materials is vast, the majority concentrated in animal studies.
Many ideas have shown pre-clinical promise but have not achieved translation to human application. Given the importance of translational impact, the focus in this section will be restricted to clinically-explored concepts and products, specifically those first tested in humans after 1976. This excludes Preamendment Devices that may have been introduced under looser regulations 486. Also not considered are decellularized tissue matrices such as amniotic membrane 487 or Cargile membrane 488. Finally, given the large variety of materials/products and potential unique indications for them, no attempts are made here to comment on their effectiveness, or advocate any material/product over another.
Past and current clinically-explored inert barrier membranes, inert barrier gels, and tissue-adherent barrier membranes/gels for peritoneal adhesion prophylaxis are summarized in Table 1-5, Table 1-6, and Table 1-7, respectively.
64 Table 1-5. Clinically-Explored Inert Barrier Membranes for Peritoneal Adhesion Prophylaxis Earliest Published Device Clinical Trial Device Description (Manufacturer) Studying Adhesion Prophylaxis
Interceed Knitted fabric of oxidized regenerated 1989 490 (Ethicon) cellulose that is absorbed within 2 weeks 489.
Thin sheet of ePTFE that is nonabsorbable; Preclude (Gore) must be sutured in place to prevent device 1992 491 migration 489.
Seprafilm Film composed of sodium hyaluronate and 1996 492 (Genzyme) CMC that is absorbed within 7 days 489.
Film composed of amorphous 70:30 poly(L- lactide-co-D,L-lactide) that is fully absorbed SurgiWrap within 6 months 489. Is chemically identical (MAST 2008 493 to the products CardioWrap and OrthoWrap, Biosurgery) respectively indicated for cardiovascular and orthopaedic applications.
65 Table 1-6. Clinically-Explored Inert Barrier Gels for Peritoneal Adhesion Prophylaxis Earliest Published Device Clinical Trial Device Description (Manufacturer) Studying Adhesion Prophylaxis
Viscous solution of 32% Dextran-70 in 10% Hyskon (Medisan dextrose, which is absorbed within 5-7 1983 495 Pharmaceuticals) days489,494.
Viscous solution of 0.5% hyaluronic acid ionically crosslinked with ferric ions; product Intergel (Lifecore was discontinued in 2003489. Has an 1998 496 Biomedical) intraperitoneal half-life of ~51 h496. This formulation also went by the trade name Lubricoat497.
Hyaluronic acid and CMC in a viscous gel Sepracoat form that is absorbed within 7 days; product 1998 498 (Genzyme) was discontinued in 1997489.
Adept (Baxter Viscous solution of 4% Icodextrin that is 2002 500 Healthcare) absorbed within 4 days489,499.
Viscous gel produced through condensation of Hyalobarrier hyaluronic acid via an auto-crosslinking (Anika mechanism501. This formulation has also gone 2003 503 Therapeutics) by the trade name Hyaloglide for indications in tendon or peripheral nerve reconstruction502.
Viscoelastic gel of poly(ethylene glycol) (PEG) and CMC, stabilized with calcium489, that is Oxiplex/AP absorbed within 30 days504. This formulation 2005 505,508 (FzioMed) has gone by other trade names for various indications: Oxiplex/SP505, Intercoat506, Dynavisc504, and Medishield507.
HyaCorp Endo Gel composed of crosslinked sodium Gel (Bioscience 2013 509 hyaluronate of non-animal origin509. GmbH)
HyaRegen NCH Gel of crosslinked hyaluronan that is absorbed (BioRegen within 1-2 weeks510. Crosslinking occurs via a 2015 510 Biomedical) thiolated chemistry511.
A-part gel (B Gel composed of PVA and CMC that is 2015 512,513 Braun) absorbed within ~6 weeks512.
66 Table 1-7. Clinically-Explored Tissue-Adherent Barriers for Peritoneal Adhesion Prophylaxis *Clinical trial described but results not published. Earliest Published Device Clinical Trial Device Description (Manufacturer) Studying Adhesion Prophylaxis
Two-component solution consisting of PEG succinimidyl SprayGel succinate (SS) and PEG-amine that forms a tissue-adherent gel (Confluent upon mixing/spraying, then is hydrolyzed and absorbed within 2003 514,515 Surgical) 20 days; SprayGel was discontinued and replaced by the next generation, reformulated product SprayShield in 2008 489.
Bi-layer membrane that converts into a hydrogel once hydrated and is absorbed within 2-3 weeks; one side is nonporous and Prevadh consists of porcine collagen, PEG, and glycerol, while the other 2008 518 (Covidien) side is porous and consists of lyophilized porcine collagen fleece designed to be placed in contact with a bleeding surface where it may promote hemostasis and adhere/integrate 489,516,517.
Tissue-adherent hydrogel formed by mixing two synthetic PEGs, a dilute hydrogen chloride solution, and a sodium phosphate/sodium carbonate solution 519. These components are CoSeal (Baxter) 2008 520 co-extruded from a syringe. This formulation is absorbed within 30 days and had previously gone by the trade name Adhibit (Angiotech Pharmaceuticals) 489,520,521.
Two-component system that consists of 4-arm 10-kD PEG-SS and trilysine amine in one component solution, and a borate SprayShield buffer in the other component solution 522. The two components 2010 523 (Covidien) gel quickly upon mixing/spraying, and are hydrolyzed and absorbed within 7 days 489. SprayShield is chemically identical to the product DuraSeal 489.
Two-component spray that consists of PEG-SS in one solution, Progel AB and human serum albumin in another, which combine to form a 2010 526 (Neomend) tissue-adherent hydrogel that is absorbed within 2 weeks 524,525.
Bi-layer membrane that consists of a non-adherent omega-3 C-Qur Film fatty acid (O3FA) layer on one side and a sodium CMC tissue- (Atrium Medical 2014* 527 adherent coating on the other side; it is absorbed within ~60 Corporation) days 527.
Two-component spray that consists of N-hydroxysuccinimide- modified carboxymethyl dextrin and trehalose stabilizer in one AdSpray (Terumo solution, and sodium carbonate and sodium bicarbonate in 2016 529 Corporation) another 528. The two solutions combine to form a tissue-adherent hydrogel, absorbed within 3 days 529. This formulation previously went by the trade name AdBlock 529.
Actamax Adhesion Barrier Dextran aldehyde and two different types of PEG-amine that are (Actamax 2017 530 sprayed together to form a tissue-adherent hydrogel 530. Surgical Materials)
67 For further information on barrier materials for adhesion prevention, readers are directed to additional reviews 483,489,531,532.
1.6.2. Mesh-Specific Preventive Strategies
Mesh-specific adhesion prophylaxis most commonly entails application of mesh devices specially designed such that the side facing the viscera mitigates tissue attachments while the parietal side promotes tissue ingrowth (i.e. adhesion-resistant meshes).
Alternative approaches that have been investigated typically entail combining mesh implantation with aforementioned general preventive strategies.
Precise and accurate terminology is important in describing nuanced designs of different meshes 46, including those intended for adhesion resistance, in order to avoid confusion. While they seem similar, mesh coatings and mesh barriers are structurally and conceptually different. Mesh coatings cover individual mesh fibers but do not span across mesh pores, while mesh barriers do span across pores and create a continuous and distinct layer parallel to the mesh 46. Meshes with barriers (i.e. barrier meshes) can be further classified as composite or noncomposite, with the former implying that the barrier layer is made of a material distinct from that of the mesh layer, and the latter implying that both layers consist of the same material but different structural features 46.
Typical component materials used in commercial hernia mesh products may be categorized based on their primary intended function: structural, barrier/coating, plasticizing, or bonding. However, note that in different products the same material may serve separate purposes, and in certain cases, one material in a given product may serve more than one function. Structural materials impart most of the mechanical properties to a mesh and can include non-resorbable polymers such as PP, poly(ethylene terephthalate)
68 (PETE, a polyester), ePTFE, and poly(vinylidene fluoride) (PVDF), resorbable polymers such as poly(glycolic acid) (PGA), and poly(lactic acid-co-glycolic acid) (PLGA), and decellularized tissues from human or animal sources. Barrier/coating materials modulate biocompatibility of and/or tissue attachments (i.e. adhesions) to other mesh components and include ePTFE, collagen, carboxymethylcellulose (CMC), hyaluronic acid (HA), PEG, oxidized regenerated cellulose (ORC), O3FA, titanium, poliglecaprone, and polyurethane.
Plasticizing materials improve the mechanical flexibility of one of the mesh components
(e.g. the barrier layer) and include PEG and glycerol. Bonding materials improve the adherence between separate layers in the mesh (e.g. between the structural layer and the barrier layer) and include PGA and polydioxanone.
With regards to commercial mesh devices that have been explored for mesh- specific adhesion resistance, products studied with relation to adhesions in publicly available literature, along with published human/animal studies reporting adhesion-related outcomes, are summarized in Table 1-8 (composite meshes) and Table 1-9 (noncomposite and other meshes). Simple meshes of PP or polyester, which are often studied as positive controls for adhesion formation, are not specifically included, but most mesh products in these tables demonstrated fewer adhesions than bare PP mesh. Given the wide variety of products to select from, the lack of specific indication for adhesion prophylaxis for any product, and the shortage of efficacy data of most products to definitively guide clinical decision-making (Section 1.2.3), no attempts are made here to comment on product effectiveness, or advocate any product over another. The difficulty in recommending one adhesion-resistant mesh product over another on the basis of adhesion formation has also been recognized elsewhere 248,263,533.
69 Table 1-8. Commercial Composite Meshes Explored for Adhesion Prophylaxis *Studies referring to meshes as “Composix” without context were assumed to pertain to Composix E/X. **Studies that described mesh composition but did not mention a specific product name are excluded. Device Human Device Description** Animal Studies See Also (Manufacturer) Studies
Composite of PP mesh and a hydrogel barrier (HA/CMC/PEG) that resorbs within 30 days. Sepramesh (Bard / 533,550, The two layers are joined with the use of 534 49,53,241,248,535–549 BD) 551 resorbable PGA filaments co-knitted with the PP.
Parietene Composite Composite of PP mesh and resorbable barrier 49,248,548,552–554 (Sofradim) layer of collagen, PEG, and glycerol.
Parietex Composite Composite of PETE mesh with resorbable 534, 52,53,248,537,538,544, 533,550, (Covidien / collagen barrier (type I atelocollagen, PEG, and 555–562 545,548,549,563–574 551 Medtronic) glycerol).
C-Qur Mesh (Atrium Composite of PP mesh with resorbable O3FA 534 248,538,567,575 533,551 Medical) barrier.
Intramesh T1 Composite of monofilament PP mesh and 248,576 (Cousin Biotech) ePTFE barrier.
Composix E/X* Composite of PP mesh and microporous ePTFE 49,236,535–538, 534,577 533,551 (Bard/BD) barrier. 548,569,578–582
Composite of lightweight monofilament PP Composix L/P mesh and microporous ePTFE barrier. L/P 533,551 (Bard/BD) stands for low profile or large pore.
Physiomesh Composite of PP mesh sandwiched between 166,583,584 533,551 (Ethicon) two layers of poliglecaprone-25.
Composite of PP mesh sandwiched between two layers of absorbable polydioxanone film, 53,236,538,539, Proceed (Ethicon) with an absorbable barrier layer of ORC Fabric 534 549,553,566–568, 533,551 on the visceral side. The polydioxanone 580,585,586 provides a bond to the ORC layer.
Hi-Tex Endo-IP Composite of multifilament PETE mesh with a 583 (THT Bioscience) smooth barrier layer of polyurethane.
Combi Mesh Composite of monofilament PP mesh with a 578 (Angiologica) barrier layer of polyurethane.
Composite of monofilament PP mesh and a Ventralight ST hydrogel barrier (HA/CMC/PEG) that resorbs 584 (Bard/BD) within 30 days. The two layers are joined with the use of resorbable PGA filaments.
Composite of monofilament PP mesh and Ventrio (Bard/BD) 551 submicronic ePTFE barrier layer.
Composite of monofilament PP mesh and a Ventrio ST hydrogel barrier (HA/CMC/PEG) that resorbs 539 551 (Bard/BD) within 30 days. The two layers are joined with the use of resorbable PGA filaments.
70 Table 1-9. Commercial Noncomposite and Other Meshes Explored for Adhesion Prophylaxis **Studies that described mesh composition but did not mention a specific product name are excluded. Device Human See Device Description** Animal Studies (Manufacturer) Studies Also
Noncomposite mesh of 52,53,192,236,237, ePTFE that is microporous 534,587, 535,537,538,540, 533, Dualmesh (Gore) and smooth on one side, 588 548,554,563,580, 551 and textured on the other. 585,589–592
Dualmesh impregnated with a coating of silver carbonate and chlorhexidine diacetate to Dualmesh Plus inhibit bacterial 594–596 (Gore) colonization of the device for up to 14 days after implantation. Product has been discontinued 593.
A single-layer mesh of condensed PTFE having Omyra Mesh (B star-shaped macropores. 583,586,597,598 Braun) Also goes by the trade name MotifMesh.
Mesh consisting of Ultrapro interwoven monofilaments 53,567,574 (Ethicon) of PP and absorbable glycolide/ε-caprolactone.
Timesh (Biomet PP mesh with a covalently 53,244,567,590, Biologics/GfE bonded coating of 551 599,600 Med. GmbH) titanium.
Mesh consisting of interwoven non- Dynamesh absorbable filaments of IPOM (FEG PVDF (88%) and PP 244,553,554,564, 558,559 551 Textiltechnik (12%), with the PP 601 mbH) filaments selectively localized on the parietal- facing side.
71 Several interesting common themes are observed across studies involving adhesion-resistant meshes. First, proper orientation of barrier meshes is critically important for both adhesion prevention and abdominal wall integration. A barrier mesh placed upside down cannot be as effective in adhesion prophylaxis as it would be in its intended orientation; interestingly however, an upside down barrier mesh might still prevent more adhesions than a bare PP mesh 248. Second, preferential sites of adhesion for barrier meshes seem to be at the mesh edges (whether cut or not) or at sites of fixation elements
248,537,546,547,567,583,595,602. Third, there is little comparability between mesh adhesion studies, especially in terms of methodology for evaluating outcomes. This prohibits correlation of adhesion metrics in animals or humans to clinically relevant outcomes, and emphasizes the need for a consistent scoring system, such as the recently proposed consensus standard
Mesh Tissue Adhesion (META) score 603. Lack of inter-study comparability has likewise been deemed an obstacle in hernia research in general 604. Finally, it has been suggested that complete prevention of mesh adhesions is impossible regardless of device features 603.
Thus, adhesion-resistant meshes might ultimately need to be implemented in combination with other general approaches to best realize complete adhesion prevention 543.
Alternative mesh-related adhesion prevention strategies, which have been attempted with varying degrees of success, most commonly include application of various materials as interpositional barriers on the visceral side of bare meshes. These materials have included CoSeal 605, Hyalobarrier 49, fibrin glue or sealant 49,594, 4DryField PH 566,606,
Prevadh 607, Surgiwrap 599,607, Seprafilm 51,581,607–613, PLGA Vicryl mesh 241,614–617,
Preclude 618,619, Interceed 619,620, Sepracoat 545, icodextrin 545, amniotic membrane 608, and omentum 616,621. Additionally, application of adhesion barrier materials may not only be
72 beneficial for adhesion prophylaxis in the presence of bare mesh, but also even in the presence of an already adhesion-resistant prosthesis, as was exemplified by combinatorial application of Seprafilm on top of composite Sepramesh 543. Other alternative strategies for preventing mesh adhesions involves combining pharmacologic agent administration with mesh implantation 402, and evaluating different materials or approaches for mesh fixation 192,573,576,586,591,597,622–624.
For further information on mesh-related adhesion prophylaxis, readers are directed to the following reviews 533,550,625–627.
1.7. Emerging Opportunities
Tremendous efforts have clearly been devoted by numerous research teams and companies to the prophylaxis of peritoneal adhesions, and considerable progress has been made. More work remains to be done, however, especially with regards to prevention of mesh adhesions, for which there are various obstacles to progress that must be overcome, and unmet clinical needs to address. Underexplored opportunities that could advance the field of mesh adhesion prevention include the advancement of non-invasive approaches to evaluate mesh adhesions, the development of multi-functional materials as adhesion- resistant barriers or coatings for meshes, and the use of currently uncommon processes to directly modify mesh surface properties and subsequent biological response.
First, there is immense room for improvement in non-invasive monitoring of mesh adhesions in both pre-clinical and clinical settings. As mentioned previously (Section
1.2.3), the gold standard for evaluating adhesions in human patients is invasive and subjective, and neither adhesions nor common mesh devices are readily seen using
73 conventional imaging approaches. This greatly hinders the study (and even diagnosis) of mesh adhesions in animals and humans.
One novel strategy for improving pre-clinical detection of mesh adhesions involves maintaining insufflation of the abdominal cavity while utilizing an imaging modality such as magnetic resonance imaging (MRI) (i.e. pneumoperitoneal MRI) 628 or computed tomography (CT) (i.e. pneumoperitoneal CT) (Figure 1-5). For mesh materials not readily visible with either modality, such as PP 57, high-radiocontrast markers, such as FibermarX
Radiopaque Tissue Marker (Viscus Biologics LLC, Cleveland, OH), may be utilized to delineate the mesh borders (Figure 1-5A). Such approaches may allow for improved quantification and determination of various adhesion characteristics. For instance, the surface area of mesh affected by adhesions could be quantified through standardized segmentation procedures. Additionally, adhesion tissue composition (e.g. vascularity, adiposity) and organ involvement could be readily evaluated. Because pneumoperitoneal
MRI and CT are non-invasive and non-destructive, these techniques would open the door to longitudinal monitoring of adhesions. At the same time, these techniques would not interfere with more common terminal adhesion outcome measures, such as scoring or mechanical testing, allowing them to be performed in parallel. Adoption and improvement of non-invasive imaging methods (such as pneumoperitoneal MRI or CT) and materials that facilitate these techniques (such as high-radiocontrast markers or meshes) will permit new insights in the study of mesh adhesions regarding their formation, prevention, treatment, and association with adverse clinical outcomes.
74
Figure 1-5. Adhesions to PP mesh are clearly visible with pneumoperitoneal CT in a rat (same animal from Figure 1-2). A) Axial section. An intestinal loop (red arrow) is visibly adhered to the mesh (yellow dashed line). FibermarX Radiopaque Tissue Marker (blue arrow) can be seen and aids with delineating mesh edges. Inset: View of implanted PP mesh, with tissue markers (blue filaments) fixed along mesh borders, prior to implantation (scale bar = 1 cm). B) Sagittal section from the same scan.
Second, apart from adhesion formation, another (particularly dreaded) complication affecting implanted hernia meshes is prosthetic infection. Reported rates of mesh infection can be as high as 10% of hernioplasty cases 629–631. As mentioned previously, such contamination often necessitates device removal or delayed implantation
(Section 1.3.1), and worsens adhesion formation (Section 1.5.3.5). One attractive solution to address infectious complications of implanted meshes involves development of unique materials as mesh coating or barrier components that can simultaneously deter adhesion and resolve infection. This might be achieved, for example, through use of the mesh coating/barrier material as a reservoir for controlled delivery of anti-microbial agents. This idea has been posed previously 632, but to date there has been slow and limited development in this area, with a select few publications 339,426. Only one commercial device, Dualmesh
Plus, has ever been explored in humans for simultaneous resistance to adhesion and
75 infection. However, the product has been discontinued 593, pointing to an unmet clinical need for meshes that can concurrently combat post-surgical adhesions and infections.
One candidate class of materials that could be highly suitable for simultaneous adhesion and infection prophylaxis could be materials based on cyclodextrins (CDs). CDs are cyclic oligosaccharides (of which there are several possible sizes), synthesized through enzymatic digestion of starch. They possess a truncated cone structure with a lipophilic core (or “pocket”) and hydrophilic exterior. CDs are generally recognized as safe (GRAS) by the FDA, and are widely used in many foods and pharmaceutical formulations. The structure of CDs imparts a unique ability to complex reversibly through thermodynamically-driven non-covalent (i.e. “affinity”) interactions with small molecules
(e.g. drugs), or their domains, which “fit” stably in the pocket 633. Insoluble polymers can be constructed from subunits of CD and used as vehicles for local and controlled drug delivery 634. Controlled local delivery is important for any pharmacologic agent (Section
1.6.1.2), but for anti-microbial drugs in particular, systemic delivery or burst release may promote hazardous bacterial resistance, as well as off-target or adverse effects to human cells and tissues. Because drug release from CD-based polymers is mediated by both affinity and diffusion, release profiles for these materials can be extended far beyond those for ordinary polymers, which release through diffusional mechanisms alone 634–636.
Additionally, by virtue of their reversible, thermodynamically-driven affinity properties,
CD polymer materials can be efficiently filled or refilled with drugs (e.g. via bolus injection of drug next to the vehicle) to allow for multiple windows of release 637. CD polymers are thus uniquely suited for controlled local release of pharmacologic agents. Furthermore, given hydrophilic surface properties conferred by the inherently polar exterior of CD
76 subunits, certain formulations of CD polymers applied as coatings have demonstrated passive resistance to non-specific protein adsorption, mammalian cell attachment, and bacterial adherence 638. CD polymers thus represent remarkable multi-functional materials that possess unparalleled advantages for simultaneous prophylaxis/eradication of mesh infection 639–641 and prevention of adhesions. Finally, drugs released from CD polymers are not limited to antibiotics, but could also include analgesic, fibrinolytic, anti-inflammatory, anti-oxidant, anti-fibrotic, and other agents that may provide additional benefits (e.g. pain reduction or further adhesion mitigation) following hernioplasty.
Lastly, opportunities to advance the field of mesh adhesion prophylaxis may also lie in the application of currently underexplored processes to directly modify surface properties of prosthetic meshes, and to investigate subsequent effects on mesh adhesions.
For example, plasma is an ionized gas that can directly alter surface properties of polymeric materials. With appropriate selection of parameters, mesh polymer surfaces can be modified using plasma to decrease nonspecific adsorption/denaturation of proteins (e.g. fibrinogen), reduce adherence of bacteria, and increase fibroblast attachment, all while maintaining prosthesis mechanical properties 642. These plasma-induced changes could respectively produce desirable effects on mesh biological response (e.g. inflammation and subsequent adhesion formation), infection, and incorporation. Alternatively, plasma can be used to directly enhance the bonding quality between a hydrophilic material, such as a hydrogel coating or adhesion barrier, and a hydrophobic polymer substrate, such as a PP mesh 643. Plasma thus represents an unexplored strategy to improve the adherence between different material components in composite mesh devices. As an added advantage, plasma may even be suitable as a terminal sterilization method for surgical meshes, as it has been
77 used to sterilize other polymeric devices such as ultra-high molecular weight polyethylene components in joint replacements 644.
1.8. Conclusions and Future Outlook
Peritoneal adhesions pose a substantial, widespread, and persistent burden in modern medicine that plagues clinicians and patients. While these internal scar-like tissues have been the subject of an expansive and enduring area of research, they are not yet fully understood. Furthermore, even less is known, appreciated, or reported about the interactions between surgical mesh biomaterials and adhesions. Certain meshes have been shown to exacerbate adhesions, and the role of meshes in adhesion formation and prevention adds an additional layer of complexity to both of these already intricate processes. Overall, the field of mesh adhesions represents an important but underexplored research area.
In efforts to address these gaps, this review has sought to synthesize and reveal novel insights regarding detailed mesh-adhesion interactions and their effects on adhesion pathogenesis and prophylaxis. In doing so, this review has broadly summarized the current knowledge of peritoneal adhesions, exhaustively examined the relationships between peritoneal adhesions and surgical mesh biomaterials, and considered emerging opportunities for progress within the field of mesh-related adhesions.
While post-surgical adhesions will always be troublesome, great advances in general and mesh adhesion prophylaxis have already been made. Further developments are still needed, however, especially in mesh adhesion prevention. Currently, there are several obstacles to future progress in this area. These include: 1) A wide variety of cleared mesh devices and little available clinical data regarding their association with adhesions (Section
78 1.2.3). 2) Lack of non-invasive and/or objective methodologies for studying mesh adhesions (Section 1.2.3). 3) General unawareness or uncertainty regarding adhesion risks, consequences, or prevention strategies (Section 1.3.3). And 4) lack of comparability between mesh adhesion studies and uncertain clinical relevance of adhesion metrics
(Section 1.6.2). On the other hand, there are several underexplored opportunities for potential advancement of the field of mesh adhesions (Section 1.7). Future studies should also seek to further elucidate mechanisms by which mesh devices impact adhesion formation and prevention, and clinical trials on efficacy of adhesion prophylaxis for mesh devices are ultimately needed.
Adhesion formation is certainly not the only concern with regards to use of surgical meshes, and other factors apart from adhesions (e.g. case-specific context, device handling, risks for recurrence or infection, etc.) must guide clinical decision-making. Additionally, not all adhesions are necessarily problematic. Despite this, peritoneal adhesions remain a genuine threat to human health and quality of life, thus their prevention is always desirable.
Thorough knowledge and awareness of adhesions, and of the impacts of surgical mesh biomaterials on their formation and prevention, is essential for biomaterials researchers, device manufacturers, physicians, and patients, to collectively achieve improved clinical outcomes.
79 1.9. Acknowledgments
The authors acknowledge support through National Institutes of Health: NIH R01
GM121477 (HvR), and NIH Ruth L. Kirschstein NRSA T32 AR007505 Training Program in Musculoskeletal Research (GDL). Additional support was provided by the Center for
Stem Cell and Regenerative Medicine Undergraduate Student Summer Program
(ENGAGE) at Case Western Reserve University (EJL). The authors also thank Erika
Cyphert, Alan Dogan, Ashley Djuhadi, and Katherine Chapin for valuable writing feedback, Aldo Fafaj for performing animal surgeries, and Kathleen Derwin for donation of the Viscus Biologics FibermarX Radiopaque Tissue Marker product.
80 2. CHAPTER 2: NONTHERMAL PLASMA TREATMENT IMPROVES
UNIFORMITY AND ADHERENCE OF CYCLODEXTRIN-BASED
COATINGS ON HYDROPHOBIC POLYMER SUBSTRATES
This chapter was adapted from a published article 643, all content being reused with permission (see APPENDIX).
Authors: Greg D. Learn, Emerson J. Lai, Horst A. von Recum
2.1. Abstract
Low surface energy substrates, which include many plastics and polymers, present challenges toward achieving uniform, adherent coatings, thus limiting intended coating function. These inert materials are common in various applications due to favorable bulk, despite suboptimal surface, properties. The ability to functionally coat low surface energy substrates holds broad value for uses across medicine and industry. Cyclodextrin-based materials represent an emerging, widely useful class of coatings, which have previously been explored for numerous purposes involving sustained release, enhanced sorption, and reversible reuse thereof. In this study, substrate exposure to nonthermal plasma was explored as a novel means to improve uniformity and adherence of cyclodextrin-based polyurethane coatings upon unreceptive polypropylene substrates. Plasma effects on substrates were investigated using contact angle goniometry and x-ray photoelectron spectroscopy (XPS). Plasma impact on coating uniformity was assessed through visualization directly and microscopically. Plasma effects on coating adhesion and bonding were studied with mechanical lap-shear testing and XPS, respectively. Substrate surface wettability and oxygen content increased with plasma exposure, and these modifications
81 were associated with improved coating uniformity, adhesion, and interfacial covalent bonding. Findings demonstrate utility of, and elucidate mechanisms behind, plasma-based surface activation for improving coating uniformity, adherence, and performance on inert polymeric substrates.
2.2. Introduction
The surface plays a critical role in the performance and success of many applied solid materials. For example, on surgically implanted devices, events such as protein adsorption, cell or bacterial attachment, biofilm formation, blood coagulation, tissue adhesion, foreign body response, and corrosion can all transpire at the host-material interface, ultimately determining the fate and function of the prosthesis, as well as the clinical outcome for the patient. Likewise, for non-medical commercial products, the surface of an item can impact its appearance, operation, and even durability, all of which influence consumer demand and profitability. Large bodies of research have been directed toward modifying material surfaces to achieve desirable function without compromising bulk properties. Within this vein, one of the most common surface modification approaches entails the application of coatings, which may serve many purposes.
Many biomedical and non-medical commercial products are comprised of polymers that possess low surface energy. Such polymeric materials include polypropylene (PP), polyethylene (PE), polydimethylsiloxane (PDMS), and polytetrafluoroethylene (PTFE).
As a focus in this work, PP is a polyolefin, being one of the most abundantly used plastics today 645. PP has many uses that range from textiles 646 (e.g. surgical meshes, sutures, clothing, and upholstery fabric), to filters/membranes 647 (e.g. in water or ventilation
82 systems, and in surgical respirators or masks), to packaging 648 (e.g. containers for food, cosmetics, or chemicals), and beyond.
The low surface energy of such materials can be disadvantageous for several reasons. First, it decreases the receptiveness of these polymers toward coatings 649, which could otherwise allow for improved product function, appeal, longevity, safety, and in the case of implantable devices, host response. In general, the surface energy of a substrate should exceed that of a coating to achieve reasonable spreading and adhesion. If not, applied coatings suffer from lack of uniformity, adherence, and durability, limiting the intended function of the coating over the product lifetime. The second disadvantage, particularly important for implants and water treatment filters/membranes, is that the nonpolar nature of low surface energy substrates promotes non-specific adsorption and denaturation of proteins on the bare surface. It is thermodynamically favorable for proteins to displace water at a hydrophobic interface and to change conformation (core nonpolar amino acids associate with the material) 222. On implant substrates, denatured proteins in this protein conditioning layer may assume altered functions or be misidentified by the host immune system, thus triggering adverse thrombotic, inflammatory, and foreign body responses 229,231,234. Furthermore, this protein conditioning layer may serve as a site for undesired attachment of biofoulants such as bacteria. On water treatment filters/membranes, adherent bacteria and their biofilms can reduce flux and contaminate treated water 650,651.
An attractive solution to promote uniformity and adherence of coatings on such challenging materials involves treatment of the substrate with nonthermal plasma 652,653.
Plasma is a state of matter composed of a dynamic mixture of negative and positive ions,
83 uncharged molecules/atoms, radicals, photons, and free electrons. It is created when sufficient energy, either in the form of heat (thermal plasma) or electromagnetic fields
(nonthermal plasma) is applied to a gas. Thermal plasmas exist only at very high temperatures (>4000K) that would destroy most polymeric materials 654,655, thus only nonthermal plasmas are used for surface modification as considered in this paper.
Interaction of nonthermal plasma with a polymer surface results in cleaning or etching, bond scission or rearrangement, and the introduction of new functional groups as determined by the carrier gas composition. These plasma-induced changes can promote substrate receptiveness to coatings 656, and also directly alter biological responses at the substrate surface 642. Furthermore, plasma treatments circumvent the need for hazardous chemicals (e.g. solvents) as adhesion promoters, are amenable for substrates with complex geometries, and can achieve desirable surface changes while minimizing impact on the substrate material bulk properties 642.
Our group and others have previously investigated polysaccharide-based materials, constructed from subunits of cyclodextrin (CD), as coatings that can serve a wide variety of unique purposes. CDs are cyclic oligosaccharides, produced through bacterial enzymatic digestion of starch, having a truncated cone geometry with a hydrophilic exterior and a lipophilic core or “pocket”. They have remarkable abilities to reversibly complex, through non-covalent interactions, with small nonpolar molecules or domains that “fit” in the pocket. Applications of these CD-based coatings include sustained delivery of antibiotics
635,637,640,657, drugs 658–660, pesticides 661, or fragrances 662, and efficient uptake and retention of pollutants 663,664 or dyes 665–667. Additionally, due to the reversible nature of their molecular interactions, these coatings may be regenerated and reused multiple times.
84 Furthermore, owing to hydrophilic properties imparted by the intrinsically polar exterior of CD subunits, certain formulations of polymerized CD (pCD) applied as coatings have recently shown potential for mitigating events of biofouling, through passive resistance to non-specific protein adsorption, mammalian cell adhesion, and bacterial attachment 638.
Within our group, pCD coatings have previously been applied to polyester surgical fabrics and metallic orthopedic screws 635,639–641. However, given that many low surface energy polymers are used as commercial materials in medicine and industry, in addition to the perpetual need for functional and stable coatings, it would be advantageous to explore the application of pCD coatings upon such difficult substrates in efforts to maximize coating uniformity and adherence.
The objective of this work is therefore to investigate effects of nonthermal plasma activation of PP substrates on the quality of pCD coatings. PP is chosen as a model substrate material given its inherently low surface energy, and its extensive use in biomedical and non-medical commercial products. A few notable examples include: (i) PP surgical textile implants, for which pCD coatings could help mitigate inflammatory, infective, and/or adhesive complications, whether through delivery-based or passive means; (ii) PP water treatment filters/membranes, for which pCD coatings could simultaneously resist biofouling and repeatedly scavenge pollutants; and (iii) PP packaging, for which pCD coatings could benefit the quality and shelf life of food through delivery of preservatives and/or extraction of undesirable spoilage byproducts.
An overview of this work is presented in Figure 2-1. The hypothesis of this study is that nonthermal plasma treatment enhances the uniformity and adherence of pCD coatings on PP substrates. To test this hypothesis, the time-dependent effects of nonthermal
85 plasma exposure on PP substrate surface characteristics were first evaluated using contact angle goniometry and X-ray photoelectron spectroscopy (XPS). Next, the effects on pCD coating uniformity and adherence were investigated. Uniformity was qualitatively assessed through both direct visualization and use of scanning electron microscopy (SEM), and semi-quantitatively investigated through spreading experiments under direct visualization.
Adherence was evaluated using lap-shear testing to evaluate mechanical adhesion in accordance with ASTM standard, and XPS to survey for chemical evidence of interfacial covalent bonds.
Figure 2-1. Chapter 2 study overview. Effects of nonthermal plasma on PP substrates were investigated in terms of wettability and surface chemistry, and effects on pCD coatings were explored in terms of uniformity and adherence.
2.3. Materials and Methods
2.3.1. Materials
Soluble, lightly epichlorohydrin-crosslinked β-CD polymer precursor (bCD) was purchased from CycloLab R&D (#CY-2009, batch CYL-4160, MW ~116 kDa; Budapest,
86 Hungary). Hexamethylene diisocyanate (HDI) crosslinker (#52649) and 2-
(trifluoromethyl)phenyl isocyanate (2-TPI) (#159379) were purchased from Sigma
Aldrich. N,N-dimethylformamide (DMF) solvent (#D119-4) was purchased from Fisher
Scientific. PP 24-well plates (#1185U58) and lids (#1185U62) for macroscopic coating observation were purchased from Thomas Scientific. PP sheet stock (#8742K133) for XPS and contact angle goniometry, and bar stock (#8782K11) for lap-shear testing were purchased from McMaster-Carr. PP 4-0 Prolene blue suture (#8592G) was purchased from eSutures.
2.3.2. Plasma Cleaning and Activation of PP Substrate Surfaces
In order to systematically examine effects of plasma treatment on pCD coatings and PP substrates, PP materials were placed in a 4” (10 cm) diameter x 8” (20 cm) length cylindrical quartz reaction chamber of a Branson/IPC Model #1005-248 Gas Plasma
Cleaner and treated with low-pressure nonthermal plasma (500 mTorr, 50 W, 13.56 MHz) using an inlet gas mixture of argon bubbled through water (Ar/H2O). Given equipment scheduling constraints, effects of plasma treatment on PP substrate wettability and surface chemistry were investigated using treatments of durations from 1-20 min performed within
6 h or 12 h of contact angle measurement or XPS analysis, respectively. The impact of plasma treatment on pCD coating uniformity and adherence was next assessed using a substrate treatment duration of 10 min (unless otherwise specified) within 1 h of pCD coating application. Non-treated (0 min) PP samples without any known prior exposure to plasma or ultraviolet light were included as controls in all experiments.
87 2.3.3. Effects of Plasma Treatment on PP Substrates
2.3.3.1. Wettability - Contact Angle Goniometry
For polar pCD coatings to spread uniformly on nonpolar PP substrates, the surface must be rendered more wettable, therefore the effect of plasma treatment on wettability of
PP was examined using contact angle goniometry. PP sheet stock was cut to dimensions of
~1.5 x 1.5 in2 (38 x 38 mm2), gently sanded to expose fresh surface using a graded series of SiC sandpaper (1200, 2500, and 5000 grit), and rinsed thoroughly with a stream of deionized water prior to performing plasma treatments (0, 1, 2.5, 5, 10, and 20 min). After sanding and/or plasma treatments, care was taken to ensure that faces to be analyzed were not inadvertently exposed directly to any liquid or solid materials before analysis. PP surfaces were then evaluated for wettability by static contact angle measurement using a
KSV Instruments CAM 200 Optical Contact Angle Meter. Deionized water droplets (n =
9-14 unique droplets per sample) of 8 µL volume were dispensed onto each PP surface and allowed to equilibrate for 30 s prior to photographing and measurement. The measurement for each droplet reflects the average of the angles on the left and right sides. Measurements were performed using KSV CAM 2008 software. Results shown represent findings from one experiment, with the same trends having also been observed in 2 similar independent experiments.
2.3.3.2. Surface Chemistry - XPS
Effects of plasma treatment on PP surface chemistry were studied using XPS to better understand the mechanistic basis by which plasma impacts spreading and adhesion of pCD coatings onto PP substrates. PP sheet stock was cut to dimensions of ~7.5 x 7.5
88 mm2, gently sanded to expose fresh surface using a graded series of SiC sandpaper, and rinsed thoroughly with a stream of deionized water prior to performing plasma treatments
(0, 1, 2.5, 5, 10, and 20 min). After sanding and/or plasma treatments, care was taken to ensure that faces to be analyzed were not inadvertently exposed directly to any liquid or solid materials before analysis. PP surfaces were then analyzed for elemental content using a PHI Versaprobe 5000 Scanning X-Ray Photoelectron Spectrometer equipped with Al Kα source (hν = 1486.6 eV). Scans were acquired on a total of n = 3-4 samples per plasma treatment duration across 2 pooled experiments, with 2 unique scan locations averaged per sample. Survey scans were collected using a 200 µm spot size, 45 W power, 15 kV acceleration voltage, 117.40 eV pass energy, 0.40 eV step size, 25 ms/step, 8 cycles, 44.7° take-off angle, and 0-1100 eV range. The C1s peak was auto-shifted to 284.8 eV, and the ratios of the elements carbon, nitrogen, and oxygen were analyzed. The areas of peaks were taken with background set using a Shirley function from 280-292 eV for C1s, 396-404 eV for N1s, and 526-538 eV for O1s. Auger peaks were not used for analysis. After survey scans, high-resolution scans were collected using a 100 µm spot size, 25.2 W power, 15 kV acceleration voltage, 23.50 eV pass energy, 0.20 eV step size, 50 ms/step, 16 cycles,
44.7° take-off angle, and 278-298 eV range for C1s, or 523-543 eV range for O1s. The vertical sampling depth ζ (from which 95% of signal originates) for take-off angle θ = 44.7° and reported inelastic mean free path of λ = 3.5 nm at a photoelectron kinetic energy of 1 keV for PP surfaces 668, is determined to be ~7.4 nm based on the relation 669 ζ = 3λcos(θ).
Analysis was performed using MultiPak software version 9.8.0.19 (Physical Electronics,
Inc.). For interpretation of high-resolution scans, C-C, C-O, C=O, and O-C=O peak positions 670–674 were constrained at 284.8±0.1, 286.3±0.1, 288±0.2, and 289±0.2 eV,
89 respectively, and full widths at half maximum (FWHM) were constrained to be within 10% that of the C-C peak FWHM value.
2.3.4. pCD Synthesis and Coating onto Surfaces
pCD coatings were synthesized in three steps using HDI as a crosslinker for bCD.
First, bCD was weighed and placed in a PP tube, and DMF was added to dissolve it, maintaining a ratio of 3 mL DMF per gram bCD. Second, HDI was added so as to initiate crosslinking, and this pre-polymer mixture was thoroughly vortexed. Third, pre-polymer mixtures were cast as coatings either: (i) into wells of PP multiwell plates for production of coated well surfaces for qualitative and semi-quantitative direct visualization of coatings, (ii) onto PP sutures for qualitative microscopic visualization of coating uniformity under SEM, or (iii) onto flat PP bar stock pieces (newly abraded using SiC sandpaper to expose fresh surface) for preparing coated specimens for lap-shear testing.
Unless otherwise specified, the amount of HDI used in experiments was 320 µL per gram bCD to achieve an intermediate crosslink ratio of 0.32. For experiments where polymer curing was necessary, cast pre-polymer mixtures were kept covered with Parafilm and allowed to cure for between 2-6 days, as specified for each experiment, under static conditions at ambient temperature and pressure. Cured pCD coatings were used promptly and directly in subsequent experiments before any drying could take place.
2.3.5. Effects of Plasma Treatment on pCD Coatings
2.3.5.1. Qualitative Uniformity - Direct Visualization
For qualitative visual assessment of pCD coating uniformity on PP substrates following plasma treatment, pCD coatings were applied to 24-well plates with or without 90 a prior 10 min exposure to plasma. In this experiment, to observe potential effects of pCD formulation on polymer spreading, the amount of HDI crosslinker was varied between extremes of 80 µL and 640 µL per gram bCD, to achieve approximate crosslink molar ratios (HDI per glucose residue) from 0.08 to 0.64, respectively, chosen to span a range from the minimum limit for gelation up to brittle materials. The volume of pre-polymer mixture added was 140 µL/well, with each unique crosslink ratio in a different column (i.e.
4 wells/formulation), then all plates were gently but thoroughly agitated to help promote complete coverage prior to covering them with Parafilm and allowing the polymers 2 days to cure at ambient temperature and pressure. Plates were then photographed and representative rows were selected to demonstrate the completeness of spreading for each pCD formulation.
2.3.5.2. Qualitative Uniformity - SEM
SEM was performed to qualitatively study pCD coating uniformity on PP substrates with or without prior plasma surface activation. Prolene sutures were removed from sterile packaging, cut into 1” long pieces, and either treated with nonthermal plasma for 10 min or left untreated (0 min). Sutures were then coated with pCD by dipping them in freshly prepared pre-polymer mixture, then were placed in a Parafilm-covered Teflon dish and allowed to cure for 2 days at ambient temperature and pressure. pCD-coated sutures were then gently adhered to a stub using carbon tape, and sputter-coated with 5nm of palladium under vacuum. Sutures were characterized using a JSM-6510 series JEOL scanning electron microscope. Images were taken at 50x magnification and an excitation voltage of
25kV.
91 2.3.5.3. Semi-Quantitative Uniformity - Direct Visualization
For semi-quantitative direct visual assessment of pCD coating uniformity on PP substrates following plasma treatment, pCD pre-polymer solutions were applied to PP 24- well plates treated with plasma for different lengths of time (0, 1, 2.5, 5, 10, and 20 min).
PP 24-well plates were held still on a stable fixture on an isolated lab bench, and solutions were slowly and gently added to each well via pipette by the same operator. The volume of pre-polymer mixture added to each well was varied for each plate to evaluate the completeness of spreading across a narrow range of volumes. All wells were used in each plate, for a total of 24 volumes tested per plasma treatment duration, with each unique volume being tested in at least triplicate. Pre-polymer solutions were allowed to spread on well surfaces for 1 min under static conditions (no agitation), then visually determined to be either completely spread across the well surface (passing result), or not (failing result).
Volumes were initially selected based on results of a preliminary study (not included), and refined slightly if needed as the experiment progressed, such that roughly half of the wells would demonstrate complete coverage, while the other half would not. Binary logistic regression was then performed on the results using a logit link function in Minitab 19 software to estimate best-fit equations, which were used to determine volumes at which
50% of wells would be expected to have a passing result after 1 min of spreading (V50).
2.3.5.4. Adherence - Lap-Shear Testing
To examine the impact of plasma surface activation on the adhesion of PP substrates to pCD coatings, lap-shear testing was performed. Lap-shear testing and specimen preparation were performed in accordance with the ASTM D3163-01(2014) standard 675. pCD pre-polymer mixtures (150 µL) were cast between two rectangular strips
92 of PP sheet substrate with overall dimensions of 4” x 1” x 1/8” (101.6 mm x 25.4 mm x
3.2 mm), on a 1” x 1” (25.4 mm x 25.4 mm) overlap region, and allowed to cure for 6 days at ambient temperature and pressure. Paired PP strips were either left untreated, or treated with nonthermal plasma for 10 min, prior to pCD application. Overlap regions were securely held together over the curing period with the use of paired medium-size binder clips and excess pre-polymer that spilled out of the joint was carefully wiped off with a
KimWipe. Lap joints (n = 5/group) were tested in tension until failure at a rate of 0.05”/min
(1.3 mm/min) and a sampling rate of 72 Hz on an Instru-Met renewed load frame (#1130) operated with Testworks 4 software and hand-tightened vice grips capable of being horizontally offset so as to minimize peel effects. Load was monitored using an Instron
100 lbf (445 N) tension load cell (cat #2512–103). The maximum load attained during the test was divided by the overlap area and recorded as the ultimate lap-shear strength
(reported in units of kPa). The work to failure was divided by the overlap area and recorded as the lap-shear toughness (reported in units of J/m2). Grips were maintained with 2.5”
(63.5 mm) between the grip edge and the bonded region, providing a 0.5” (12.7 mm) length of specimen/grip overlap. Results shown represent findings from one experiment, with the same trend having also been observed in one similar independent experiment.
2.3.5.5. Interfacial Covalent Bonding - XPS
In order to better understand the mechanisms behind plasma treatment effects on
PP substrate adhesion to pCD coatings, we sought to determine whether covalent bonding takes place between pCD coatings and the plasma-treated PP surface. However, to do this an isocyanate compound was needed that could be uniquely identified at the activated PP surface. This was done using XPS to detect fluorine as a unique marker of the isocyanate
93 compound 2-TPI after exposure of this compound to untreated and plasma-treated PP surfaces. PP sheet stock was first cut to dimensions of ~7.5 x 7.5 mm2, gently sanded to expose fresh surface using a graded series of SiC sandpaper, and rinsed thoroughly with deionized water prior to performing plasma treatments (0 versus 10 min). Sample surfaces were then directly exposed to 2-TPI overnight, followed by two sequential rinses with copious amounts of toluene, acetone, isopropanol, then deionized water. Finally, a stream of dry nitrogen gas was used to remove excess water from sample surfaces. Untreated and plasma treated PP surfaces that were not exposed to isocyanate were prepared as controls as well. After sanding, plasma, and isocyanate exposure, care was taken to ensure that faces to be analyzed were not inadvertently touched by any solid surface before or during analysis. PP surfaces were then analyzed for elemental content using XPS as above. XPS was performed within 48h of plasma treatment. Survey scans were acquired on 2 samples per plasma treatment duration, with 2 unique scan locations per sample. The ratios of carbon, nitrogen, oxygen, and fluorine were analyzed using Multipak software. The areas of peaks were taken with background set using a Shirley function as before for C1s, N1s, and O1s, and from 675-695 eV for F1s. Auger peaks were not used for analysis. Increased fluorine content was attributed to covalent urethane bond formation between hydroxyl groups on the PP surface and the isocyanate group on 2-TPI.
2.3.6. Statistical Analysis
All data is presented as mean ± standard deviation. Unless otherwise specified, statistical analysis tests were carried out in Microsoft Excel 2016, with two-sample two- tailed Student’s t-tests with unequal variance used for all comparisons. Statistical
94 significance was set at p < α = 0.05. Plots depicting representative data were constructed using the sample closest to the group mean value(s).
2.4. Results
2.4.1. Effects of Plasma on PP Substrate Wettability and Surface Chemistry
First, we sought to evaluate the effects of plasma on the substrate surface, particularly in terms of wettability and chemistry. Wettability was studied through contact angle goniometry using the sessile drop method. Plasma treatment for any length of time was found to increase the wettability of PP substrates (p < 0.001) (Table 2-1), with longer treatments correlating with decreased static water contact angles. Without plasma treatment, water contact angles on PP averaged >120°, but after 20 min of plasma exposure, this value decreased to <60°.
PP substrate surface chemistry was studied using XPS. Quantitative analysis of survey scan spectra (Figure 2-2) indicated that plasma exposure for any duration enhanced the amount of oxygen on PP surfaces (p < 0.001), and decreased the proportion of carbon
(p < 0.001), with no significant impact on surface nitrogen content (p > 0.1) until 10 or 20 min (p < 0.014) of treatment (Table 2-1). Longer treatments were associated with increased oxygen, slightly increased nitrogen, and decreased carbon concentrations. Engraftment of oxygen-containing groups likely explains the enhanced surface wettability of PP following plasma treatment.
High-resolution C1s spectra were nearly symmetric for the untreated PP surface, but demonstrated an increasing skew to the left as plasma exposure time increased (Figure
2-3), suggesting increasing abundance of oxygen-containing functionalities including C-
95 O, C=O, and O-C=O. Deconvolutions of high-resolution C1s scans suggested that most engrafted oxygen was incorporated as hydroxyl groups, with lesser amounts as ketones/aldehydes and carboxyls/esters, for all tested plasma durations. This finding is in agreement with previous reports 642,670. However, the abundance of hydroxyl groups leveled off after 10 min of treatment, as existing hydroxyls were oxidized further to species such as ketones, carboxylic acids, or possibly even carbonates 671,672,674. These latter two oxygen-rich functional groups can only be created through chain scission at the substrate surface, and such degraded polymer chains may embrittle the PP material 642. Furthermore, loss of hydroxyl groups at the PP surface might be predicted to decrease the extent of interfacial covalent bonding with the polyurethane coating, as hydroxyls undergo expected reaction with isocyanates 676,677. This would tend to decrease coating adherence. For these reasons, a treatment duration of 10 min was selected for most subsequent experiments.
Table 2-1. Effect of plasma treatment duration on PP wettability, and surface chemistry. *Significant difference to 0 min. ⁑Significant difference to all subsequent time points. Plasma Contact Angle Atomic % Atomic % Atomic % Duration (°) Carbon Nitrogen Oxygen (Minutes) 0 126.8 ± 4.8⁑ 98.27 ± 0.35⁑ 0.29 ± 0.19 1.45 ± 0.24⁑
1 95.0 ± 7.4 88.47 ± 0.70 0.34 ± 0.19 11.19 ± 0.51
2.5 87.1 ± 10.4 88.18 ± 0.75 0.45 ± 0.40 11.37 ± 0.63
5 70.1 ± 9.5 85.79 ± 0.29 0.53 ± 0.17 13.68 ± 0.28
10 65.8 ± 9.9 82.18 ± 0.89 0.71 ± 0.07* 17.11 ± 0.85
20 57.9 ± 9.7 79.93 ± 1.28 0.72 ± 0.15* 19.36 ± 1.20
96
Figure 2-2. Representative XPS survey scans (rescaled to normalize areas under curves) of the 0 and 20 min time points demonstrate increasing height of the O1s peak with plasma treatment. KLL peaks indicate atomic relaxation via the Auger effect, and were not used for analysis. Spectra are intentionally offset from each other along the vertical axis.
97
Figure 2-3. Representative XPS high-resolution scans demonstrate increasing abundance of oxygen-containing functional groups with increasing plasma treatment time. Spectra are intentionally offset from each other along the vertical axis.
98 2.4.2. Effects of Plasma on pCD Coating Uniformity, Adherence, and Interfacial
Covalent Bonding
Having characterized the effects of plasma treatments on PP substrates, we next aimed to evaluate the effects of the plasma on pCD coating uniformity and adherence.
First, coating uniformity was qualitatively assessed based on direct and SEM visualizations. pCD coatings uniformly and stably covered the surfaces of PP substrates after plasma treatment, but not otherwise. This was observed both directly on initially- agitated PP 24-well plate surfaces (Figure 2-4), and microscopically on PP suture surfaces
(Figure 2-5).
For PP 24-well plates, coatings tended to bead up at well edges without prior substrate plasma treatment, resulting in inadequate surface coverage at well centers. These well coatings were also observed to delaminate more easily from well surfaces, especially at higher crosslink ratios when pCD tended to contract more upon curing. Following plasma treatment, coatings applied in the same manner covered the entire surface of each well, regardless of the HDI crosslink ratio.
For PP sutures, SEM images demonstrated that pCD coatings beaded up and covered very little of the untreated PP suture surface (Figure 2-5A). At the same time, pCD coatings applied in the same manner were able to spread and coat a much larger fraction of the surface on plasma treated sutures (Figure 2-5B).
99
Figure 2-4. Plasma treatment of PP substrates for 10 min improves macroscopic pCD coating uniformity on wells of a PP 24-well plate. This remained true regardless of the amount of HDI crosslinker added.
Figure 2-5. Scanning electron micrographs demonstrate that plasma treatment of PP suture improves microscopic pCD coating uniformity. A) An untreated and pCD-coated PP suture. The pCD coating beads up instead of spreading on untreated PP suture; B) A 10 min plasma-treated and pCD-coated PP suture. The pCD coating is thin and well-spread on the plasma-treated suture, as made apparent based on the presence of cracks caused by sample drying upon sputtering and imaging.
Next, coating uniformity was semi-quantitatively assessed by evaluating, across a range of volumes, whether or not pCD pre-polymer solutions were able to completely spread within 1 min under static conditions on PP 24-well plate surfaces that had been treated with plasma for various durations of time. Binary logistic regression was then performed on the results so as to determine the volume for each plasma exposure duration at which 50% of wells would show complete coverage, V50. In agreement with the
100 wettability study results, V50 was lower for plasma-treated wells than for untreated wells
(Table 2-2), implying that plasma treatment decreased the volume of pCD pre-polymer solution required to achieve complete coverage of PP wells. Additionally, V50 tended to decrease with increasing plasma duration, with the minor exception of the 20 min group
(though this small deviation might reflect low sensitivity of the test method to detect such slight differences).
Table 2-2. Effect of plasma duration on calculated pCD pre-polymer solution volume needed for complete coverage of half of tested PP well surfaces (V50). Plasma Duration V (µL) (Minutes) 50 0 246.67 1 163.75 2.5 163.75 5 162.34 10 159.65 20 166.81
Taken together, results of the qualitative and semi-quantitative coating uniformity studies indicate that plasma treatments of PP substrates improve pCD coating uniformity.
In light of the PP wettability findings, this improved uniformity likely reflects decreased substrate hydrophobicity, which promotes spreading of the polar pre-polymer mixtures on the surface.
Coating adherence was next investigated using lap-shear testing, an approach used for measuring the bond shear strength of coatings and adhesives, which here involved pulling apart two rectangular PP adherends parallel to a bond line of pCD applied over a defined area. Lap-shear testing revealed that plasma treatment of PP substrates for 10 min increased average ultimate lap-shear strength by 43% (p = 0.003) and doubled average lap-
101 shear toughness (p = 0.026) of pCD coatings (Table 2-3). The ultimate lap-shear strength and toughness are related to the maximum load (peak) and work (integral) for the load- displacement curve (Figure 2-6), respectively. For general reference, the ultimate lap-shear strength of untreated PP adherends bonded with cyanoacrylate has been reported as 0.22
MPa 678, approximately the strength observed here for untreated PP bonded with 0.32 pCD, though cyanoacrylate most likely possesses higher cohesive strength than the tested pCD.
The enhanced coating adherence following plasma treatment may result from the evident engraftment of hydroxyl groups onto the PP surface, which might react to form covalent urethane linkages between the coating and substrate upon diisocyanate crosslinking 676,677.
Table 2-3. Effect of plasma treatment on pCD coating adherence to PP substrates. *Significant difference to 0 min control. Plasma Duration Ultimate Lap-Shear Lap-Shear (Minutes) Strength (kPa) Toughness (J/m2) 0 222.3 ± 31.7 32.9 ± 12.3
10 318.7 ± 40.0* 66.4 ± 22.5*
Figure 2-6. Load-displacement curves from lap-shear testing representative for untreated and plasma-treated substrates.
102 To validate this explanation that interfacial covalent bonding was partly responsible for enhanced coating adherence following plasma treatment, we next aimed to evaluate the formation of covalent urethane bonds between functional groups at the coating-substrate interface. This was done using XPS to detect fluorine as a unique marker of an isocyanate compound, 2-TPI (Figure 2-7), after exposure of this compound to untreated and plasma- treated PP surfaces followed by copious rinsing. In comparison to non-2-TPI-exposed controls, incubation of untreated PP samples with 2-TPI resulted in no significant change in surface fluorine content (p = 0.5), while incubation of plasma-treated PP samples with
2-TPI led to a significant increase in surface fluorine content (p = 0.018) (Figure 2-8,
Table 2-4). This supports that plasma treatment of PP substrates enables covalent linkages between the PP surface (most likely via hydroxyl groups) and compounds containing isocyanate groups (such as pCD coatings during HDI crosslinking). This result corroborates the finding that plasma treatment improved the adherence of pCD coatings, and suggests that this effect can be attributed, at least in part, to covalent bond formation at the coating-substrate interface. Note that the 10 min plasma treated PP that was not exposed to 2-TPI in this experiment showed lower surface oxygen content than was seen before (Table 2-1). This could reflect an extended period between plasma treatment and
XPS (given overnight 2-TPI incubation), as well as the employed rinsing protocol.
Figure 2-7. Chemical structure of 2-TPI. Figure created using ChemDraw Professional software (PerkinElmer Informatics, Inc). 103
Figure 2-8. Representative XPS survey scans (rescaled to normalize areas under curves) of PP after 0 or 10 min plasma, plus equivalent exposure to 2-TPI. Increased isocyanate binding is indicated by the appearance of an F1s peak (arrow) only for the 10 min plasma sample. KLL peaks were not used for analysis. Spectra are intentionally offset from each other along the vertical axis.
Table 2-4. Effect of plasma treatment on isocyanate-mediated fluorination of PP surfaces. *Significant difference to non-2-TPI-exposed control. Plasma 2-TPI Atomic % Atomic % Atomic % Atomic % Duration Exposure Carbon Nitrogen Oxygen Fluorine (Minutes) 0 - 99.34 ± 0.37 0.42 ± 0.02 0.25 ± 0.35 0.00 ± 0.00
0 + 99.55 ± 0.64 0.00 ± 0.00 0.15 ± 0.21 0.31 ± 0.43
10 - 87.09 ± 0.11 0.89 ± 0.06 11.97 ± 0.26 0.07 ± 0.09
10 + 90.41 ± 0.49 0.72 ± 0.32 7.48 ± 0.53 1.39 ± 0.08*
104 2.5. Discussion
Uniformity and adherence to the substrate are crucial for the performance and function of any coating. For example, delamination of a drug-eluting polymer coating on an intravascular implant could pose risk for lethal thrombosis or embolism 679. Similarly, improper adherence of paint to a metal component (e.g. for an automobile) could accelerate corrosion and potential structural failure of the part 680,681. Likewise, a lack of adequate substrate coverage of an anti-biofouling coating could enable organisms to exploit coating surface defects and attach to the underlying substrate 682–684, creating problems such as infection on a medical device or increased drag on a marine vessel.
CD-based coatings have numerous potential applications relevant for inert polymeric substrates like PP. For example, on PP surgical textile implants, they could deliver antibiotics/drugs 640,660, or provide intrinsic resistance to protein/cell/bacterial attachment 638 to help mitigate inflammatory, infective, and/or adhesive complications. On
PP non-medical textiles, they could be used to improve fabric dyeability 666, for sustained and controlled release of insect repellent compounds 685–687 and fragrances 688–690, and for incorporating protectants against ultraviolet radiation 691,692 to which PP has known degradation susceptibility 693. In PP water treatment filters/membranes, they could be useful as reusable adsorbents for scavenging of pollutants from water 663,694,695, and for limiting attachment and growth of biofouling microbial biofilms 638. In PP packaging materials, they could be used to improve the quality and shelf life of food, through release of anti-microbial or preservative compounds 696,697, or through scavenging of undesirable components 698,699, either directly from the food or from the package headspace.
105 This study was designed to test the hypothesis that nonthermal plasma treatment enhances the uniformity and adherence of pCD coatings on PP substrates. The observations from contact angle goniometry, XPS, direct and SEM visualization, and lap-shear testing support this hypothesis. Plasma activation was found here to be a suitable strategy for improving both uniformity and adherence of polymer coatings such as pCD, on inert and otherwise incompatible polymer substrates like PP.
This conclusion is in agreement with those from prior studies that have investigated effects of plasma on uniformity and adherence of other polyurethane coating materials to low surface energy polymer substrates 677,700. Martinez et al. determined that atmospheric plasma treatment of PDMS and acrylonitrile-butadiene-styrene substrates could enhance the adhesion of polyurethane-based paints in scratch, cross-cut, and pull-off tests 700. Bao et al. found that atmospheric plasma treatment engrafted hydroxyl and amine groups onto styrene-butadiene-styrene rubber surfaces, and that these specific changes in surface chemistry, more so than incorporation of halogen functionalities, enabled increases in T- peel strength of polyurethane coatings after addition of an isocyanate-terminated hardener
677. Sanbhal et al. showed that oxygen cold plasma treatment of PP mesh substrates enhanced the uniformity of pCD coatings 676. Extending the knowledge gained from these previous investigations, through detailed surface analysis in combination with mechanical lap-shear adhesion testing, this study revealed direct evidence for formation of covalent bonds at the interface between a plasma-treated polyolefin substrate and a polyurethane coating, clarified the basis for these interfacial connections, and identified their contribution to improved strength of adhesion.
106 Substrate plasma activation is advantageous for improving uniformity and adherence of coatings upon inert polymer substrates when compared to conventional treatments, which utilize harsh chemicals as adhesion promoters. The former strategy represents a greener chemistry process, and eliminates the risk for residual hazardous compounds, which is especially important for any materials that may ultimately come into contact (directly or indirectly) with human cells or tissues (e.g. medical devices, water treatment filters/membranes). Furthermore, appropriate selection of plasma treatment parameters can ensure minimal impact on bulk mechanical properties of load-bearing substrates 642. Apart from CD-based coatings, findings herein support the use of substrate plasma activation for improving uniformity and adherence of adhesives or paints
(especially polyurethane-based formulations) upon low surface energy substrates such as polyolefins.
Future studies could implement these findings to achieve enhanced uniformity and adherence of coatings for improved functionality and durability in important applications, such as those mentioned previously, involving low surface energy substrates. Additionally, effects of plasma treatment on pCD coating uniformity and adherence could be investigated in detail for other polymer substrate materials, different plasma processes, and other CD- based polymer chemistries/formulations. Also, changes in coating uniformity or adherence over long periods after application could be assessed, as the effects of plasma on surface chemistry of bare substrates are known to diminish (i.e. age) over time. Prior studies suggest that hydrophobic recovery of bare plasma-treated PP normally occurs over the course of several weeks, the rate being dependent on environmental factors and polymer crystallinity 701–704. However, the application of a covalently anchored coating, such as
107 pCD in this study, may limit conformational changes and reorientation of the activated substrate surface, in theory preserving the coating uniformity and adherence long term.
2.6. Acknowledgments
The authors gratefully acknowledge support from National Institutes of Health:
NIH R01 GM121477 (HvR), and NIH Ruth L. Kirschstein NRSA T32 AR007505 Training
Program in Musculoskeletal Research (GDL). Additional support was provided by the
Center for Stem Cell and Regenerative Medicine Student Summer Program (ENGAGE) at
Case Western Reserve University (EJL). Core facility services provided by the Swagelok
Center for Surface Analysis of Materials and the Advanced Manufacturing and Mechanical
Reliability Center at CWRU are also appreciated. The authors also thank the Advincula group for use of the contact angle goniometer, the Eppell group and Eloise Miller for use of and guidance with operating the plasma cleaner, Kevin Abbassi for expertise and assistance with XPS, Kathleen Young for expertise and assistance with SEM, Katherine
Yan for assistance with sample preparation, and Erika Cyphert and Nathan Rohner for valuable writing feedback.
108 3. CHAPTER 3: NONTHERMAL PLASMA TREATMENT OF POLYMERS
MODULATES BIOLOGICAL FOULING BUT CAN CAUSE MATERIAL
EMBRITTLEMENT
This chapter was adapted from a published article 642, all content being reused with permission (see APPENDIX).
Authors: Greg D. Learn, Emerson J. Lai, Emily J. Wilson, Horst A. von Recum
3.1. Abstract
Plasma-based treatment is a prevalent strategy to alter biological response and enhance biomaterial coating quality at the surfaces of biomedical devices and implants, especially polymeric materials. Plasma, an ionized gas, is often thought to have negligible effects on the bulk properties of prosthetic substrates given that it alters the surface chemistry on only the outermost few nanometers of material. However, no studies to date have systematically explored the effects of plasma exposure on both the surface and bulk properties of a biomaterial. This work examines the time-dependent effects of a nonthermal plasma on the surface and bulk (i.e. mechanical) properties of polymeric implants, specifically polypropylene surgical meshes and sutures. Findings suggest that plasma exposure improved resistance to fibrinogen adsorption and Escherichia coli attachment, and promoted mammalian fibroblast attachment, although increased duration of exposure resulted in a state of diminishing returns. At the same time, it was observed that plasma exposure can be detrimental to the material properties of individual filaments (i.e. sutures), as well as the structural characteristics of knitted meshes, with longer exposures resulting in further embrittlement and larger changes in anisotropic behavior. Though there are few
109 guidelines regarding appropriate mechanical properties of surgical textiles, the results from this investigation imply that there are ultimate exposure limits for plasma-based treatments of polymeric implant materials when structural properties must be preserved, and that the effects of a plasma on a given biomaterial should be examined carefully before translation to a clinical scenario.
3.2. Introduction
Upon implantation, biomedical materials are rapidly covered by proteins, followed by host cells, or bacteria in the case of infection. The nature of these biological attachments often dictate the success of a surgically implanted device. For example, in hernia repair surgery, polymer fabrics called “meshes” are typically used for abdominal wall defect closure. These meshes are commonly associated with complications, particularly post- surgical adhesions and prosthetic infection. Specifically, out of over 350,000 hernia repair procedures performed each year in the US 21, 13.6% of patients are rehospitalized for adhesion-related complications within 5 years of surgery 112, and between 1-10% of implanted meshes become infected 630,631. Tissue adhesions to mesh are initiated by deposition and maturation of a fibrin matrix on the mesh surface as well as an adjacent tissue/organ 705, whereas prosthetic infection results from bacterial attachment to the protein-conditioned surface and subsequent biofilm formation. To mitigate such adverse events, many research efforts have focused on altering surface properties of synthetic meshes, either directly or through application of anti-adhesion 248,533 or antibiotic coatings
706,707. However, many synthetic mesh polymers, especially polypropylene (PP), which is the most common mesh material 46–50, are inert, possess low surface energy, and bond poorly (weakly and/or with nonuniform coverage) with coatings.
110 One strategy that has been explored for direct surface modification, as well as enhancing the uniformity, durability, and adherence of coatings, on difficult substrates like
PP is surface treatment with a man-made plasma. Plasma, the fourth state of matter, is composed of a mixture of positive and negative ions, neutral atoms/molecules, radicals, free electrons, and photons (particularly visible and ultraviolet light). It is created when a gas becomes ionized through application of sufficient energy, either through heat (thermal plasma) or electromagnetic fields (nonthermal plasma). Thermal plasmas exist only at high temperatures that would destroy most substrate materials 654, thus only nonthermal plasmas are used for surface modification as considered in this paper.
Interaction of plasma with a polymer surface results in surface cleaning/etching, scission or rearrangement of bonds, and the introduction of new functional groups (as determined by the composition of the carrier gas). In the absence of coatings, these plasma- induced changes can be used to tune biological response to the material. For example, surface engraftment of oxygen can enhance wettability, in turn reducing protein adsorption and conformational denaturation that might otherwise trigger detrimental inflammatory responses on a hydrophobic implant surface 229,708, and modulating subsequent cell/bacterial attachment. Alternatively, these plasma-induced changes can contribute to improved receptiveness/adherence to coatings 656, circumventing the need for hazardous chemicals as coating adhesion promoters, on inert polymers such as PP. Additionally, plasma can easily treat objects with complex geometries and high surface areas. Numerous studies have therefore explored plasma surface treatments for surgical meshes 676,709–728.
Another frequently-argued advantage of plasma-based surface treatment for biomedical polymers such as mesh substrates is that plasma exposure is thought to have
111 negligible effects on bulk mechanical properties, given that the treatment modifies only the outermost few nanometers of the material surface 676,709,710,716,722–725,729–731. To our knowledge, however, no prior studies have systematically verified this lack of effect on bulk properties or examined what limits there are to the established beneficial effects of plasma exposure. This paucity of data is concerning for a few reasons: 1) most surgical meshes have relatively high surface-to-volume ratios, 2) plasmas emit photons in the ultraviolet range 674,732,733, 3) ionizing radiation is known to embrittle PP 734–737 and other polymeric materials, and 4) degraded polymer chains at the material surface may resemble nano-scale cracks that could propagate into the bulk of the material upon application of mechanical stress. Considering this, the objective of this work is to evaluate the time- dependent effects of typical nonthermal plasma on both the surface and bulk mechanical properties of surgical mesh devices. We hypothesize that plasma treatment alters biofouling by proteins, cells, and bacteria, but extending exposure accelerates the mechanical failure of PP biomedical textiles.
An overview of this work is depicted in Figure 3-1. PP substrates, primarily surgical textiles and microplates, were treated with nonthermal plasma for selected durations from 0 to 20 minutes. Effects on PP surface composition, protein adsorption, bacterial attachment, and mammalian cell attachment were investigated with X-ray photoelectron spectroscopy (XPS), fluorescence spectroscopy, bioluminescence spectroscopy, and hemacytometry, respectively. Next, effects on mesh material and structural properties were evaluated. Uniaxial tension tests were performed on individual
PP filaments and ASTM standard dogbone-shaped mesh specimens cut along two
112 orthogonal axes. Finally, suture retention, tear-resistance, and ball burst tests were performed according to standardized methods 264,738 using rectangular mesh pieces.
Figure 3-1. Chapter 3 study overview. Effects of nonthermal plasma on properties of PP substrates was explored at both surface and bulk levels. PP surfaces were investigated in terms of surface chemistry (mesh substrates), and resistance to protein, bacterial, and fibroblast attachment (microplate substrates). PP bulk was investigated in terms of mechanical properties of knitted meshes overall and at the filament level.
3.3. Materials and Methods
3.3.1. Materials
Prolene PP mesh, Prolene PP Soft mesh, and 4-0 Prolene PP monofilaments were manufactured by Ethicon, Inc. The extruded PP in each of these products is identical in composition according to manufacturer instructions for use. Prolene Soft mesh was used for XPS analysis due to the presence of bundled filaments with wider overall diameters
(facilitating higher photoelectron counts), while Prolene mesh and monofilament sutures were used for mechanical tests. The monofilament size of 4-0 was selected to closely approximate the ~130 µm diameter 738 of filaments in Prolene mesh. For measurement of protein adsorption and bacterial attachment, PP 96-well plates (#M9685) were purchased
113 from Sigma Aldrich, fluorescein isothiocyanate (FITC)-labeled fibrinogen (#FIB-FITC) was purchased from Molecular Innovations, Luria Bertani (LB) broth (#BP1426-500) and tissue culture polystyrene 96-well plates (#07-200-90) were purchased from Fisher
Scientific, and bioluminescent ilux pGEX(-) Escherichia coli 739 was purchased from
Addgene (plasmid #107879). For measurement of fibroblast attachment, PP 24-well plates
(#1185U58) and lids (#1185U62) were purchased from Thomas Scientific, tissue culture polystyrene 24-well plates (#08-772-1H), DMEM with low glucose (#10-567-014), fetal bovine serum (FBS) (#16-141-079), penicillin-streptomycin (#15-140-122) were purchased from Fisher Scientific, and NIH/3T3 fibroblasts (#AKR-214) were purchased from Cell BioLabs, Inc.
3.3.2. Plasma Treatment
PP materials were exposed to plasma at varying durations to examine the effects of treatment time on both surface and bulk properties. For studies on surface chemistry and mechanical properties, meshes and monofilaments were removed from sterile packaging, cut to size, and pre-treated for 1, 2.5, 5, 10, or 20 minutes with a low-pressure 500 mTorr
(67 Pa) Ar/H2O nonthermal plasma (50 W, 13.56 MHz) using a Branson/IPC Model #1005-
248 Gas Plasma Cleaner. Plasma treatment was performed within 18 h of XPS analysis and
6 h of mechanical testing. For protein, bacterial, and mammalian cell attachment experiments, PP 96-well and PP 24-well plates were plasma-treated in similar fashion within 2 h of biofoulant exposure. Non-treated (0 min) PP samples with no known prior exposure to ultraviolet light were included as controls in all experiments, and tissue culture polystyrene (a common substrate for cell culture and attachment in vitro) microplate surfaces were included in protein, bacterial, and mammalian cell attachment experiments.
114 3.3.3. Surface Characterization
Effects of plasma exposure on PP surfaces were evaluated in terms of surface composition, protein adsorption, bacterial attachment, and mammalian cell attachment.
These were investigated using XPS, fluorescence spectroscopy, bioluminescence spectroscopy, and hemacytometry, respectively.
3.3.3.1. X-ray Photoelectron Spectroscopy
Effects of plasma exposure on PP implant surface chemistry were determined using
XPS. Prolene Soft mesh surfaces were analyzed for elemental content using a PHI
Versaprobe 5000 Scanning X-Ray Photoelectron Spectrometer equipped with Al Kα source (hν = 1486.6 eV). Scans were acquired on 1 scan location per sample per plasma treatment duration. Survey scans were collected using a 200 µm spot size, 45 W power, 15 kV acceleration voltage, 117.40 eV pass energy, 0.40 eV step size, 25 ms/step, 8 cycles,
44.7° take-off angle, and 0-1100 eV range. The C1s peak was auto-shifted to 284.8 eV, and the ratios of the elements carbon, nitrogen, oxygen, and fluorine were analyzed. The areas of peaks were taken with background set using a Shirley function from 280-292 eV for C1s, 396-404 eV for N1s, 526-538 eV for O1s, and 675-695 eV for F1s. After survey scans, high-resolution scans were collected using a 100 µm spot size, 25.2 W power, 15 kV acceleration voltage, 23.50 eV pass energy, 0.20 eV step size, 50 ms/step, 16 cycles,
44.7° take-off angle, and 278-298 eV range for C1s. The vertical sampling depth ζ for take- off angle θ = 44.7° and reported inelastic mean free path of λ = 3.5nm at a photoelectron kinetic energy of 1 keV for PP surfaces 668, is estimated to be ~7.4 nm based on the relation
669 ζ = 3λcos(θ). Analysis was performed using MultiPak software version 9.8.0.19
(Physical Electronics, Inc.).
115 3.3.3.2. Fibrinogen Adsorption
Protein adsorption is one of the first stages in biological fouling occurring after implantation. While many proteins adsorb and compete for an implant surface, fibrinogen is one of the most studied, and was chosen as a model protein here given its common involvement in inflammatory responses to biomaterials 229,231 and post-surgical adhesion formation 705. FITC-labeled fibrinogen was thawed and diluted to a working concentration of 50 µg/mL in sterile PBS containing 2 mM sodium azide. This solution was pipetted onto
PP 96-well plates at 100 μL/well before covering plates with Parafilm and foil, and incubating them overnight at 37°C. Following incubation, all wells were rinsed with sterile
PBS such that aspiration was performed 5 times, standard curves with known FITC- fibrinogen concentrations were pipetted in empty rows for each microplate, and a Biotek
Synergy H1 plate reader was used to read the fluorescence in each well using excitation and emission wavelengths of 490 nm and 525 nm, respectively. Data presented reflects 12 wells/condition from one experiment, and the same trends were confirmed across one additional independent experiment (data not shown).
3.3.3.3. Bacterial Attachment
Attachment of bacteria to a biomaterial surface is a key step that precedes any prosthetic infection, and E. coli was chosen as a model microorganism for bacterial attachment to PP given its common involvement in mesh infection following repair of incarcerated abdominal hernias 740, being derived from the gut. Bioluminescent E. coli were thawed from frozen stock, inoculated into a 14 mL round-bottom tube of sterile LB broth with vented lid, and expanded in suspension for 18 h in a dedicated 37°C incubator with the tube lid vented. The tube was then removed from the incubator and stored at 4°C to
116 maintain bacteria in the stationary phase. Prior to seeding onto experimental surfaces, the bacterial suspension was diluted to an optical density at 600 nm (OD600) of 0.50 relative to sterile LB broth. The bacterial suspension was then directly seeded onto PP 96-well plates at 100 µL/well, cultured under static conditions for 24 h at 37°C, and following incubation, wells were rinsed 4 times with 200 µL/well sterile PBS, then emptied and filled a final time with 100 µL/well sterile LB broth. Several known dilutions of bacteria were included for creation of standard curves in duplicate for each type of surface.
Bioluminescence measurement was then performed using a Biotek Synergy H1 plate reader
(luminescence endpoint scan, 5 s integration, 1 mm read height, full light emission, 135 gain, top optics, 100 ms delay, extended dynamic range). Data presented reflects 14-16 wells/condition from one experiment, and the same trends were observed across one additional independent experiment using PP mesh substrates (data not shown).
3.3.3.4. Fibroblast Attachment
Attachment of host cells to a biomaterial surface is an important event that can lead to tissue ingrowth around the implant as adherent cells deposit and remodel matrix proteins.
Fibroblasts were chosen as a model mammalian cell type given their role in incorporation of mesh implants with host abdominal wall tissue, a step that is necessary to anchor the mesh in place and prevent its migration. NIH/3T3 fibroblasts were cultivated on 100 mm tissue culture polystyrene dishes in culture medium consisting of 89% DMEM high glucose, 10% FBS, and 1% penicillin-streptomycin, and cells were subcultured before reaching confluence. Fibroblasts were trypsinized, resuspended, counted, and seeded at
30,000 cells/cm2 onto PP 24-well plates at 500 μL/well, then allowed to attach for 18h in a 37°C incubator under 5% CO2 and 95% humidity. To wash away non-adherent cells, each
117 well was gently rinsed thrice with 750 μL sterile PBS. Next, 200 μL/well trypsin, followed by two rinses with 350 μL/well culture medium, were used to detach cells and collected for cell quantification using a hemacytometer. Data presented reflects 6 wells/condition from one experiment, with one hemacytometry count performed for each well.
3.3.4. Bulk Characterization
A variety of mechanical tests were performed to determine the extent of bulk property changes following plasma treatments. These tests included uniaxial tension for individual PP filaments and dogbone-shaped mesh specimens, and suture retention, tear- resistance, and ball burst tests of rectangular mesh specimens. Unless otherwise specified, tests were performed using an Instru-Met renewed load frame (#1130) operated using
Testworks 4 software, an Instron 100 lbf (445 N) tension load cell (cat #2512-103), and a sampling rate of 100 Hz. Load to failure was defined as the maximum load (N) sustained by the specimen during the test, which for some tests may not have coincided with displacement to failure. Displacement to failure was defined as the magnitude of crosshead travel (mm) at which the specimen completely ruptured and could no longer bear load.
Work to failure was defined as the energy (J) absorbed by the specimen from the beginning of the test (0 mm) up until the displacement to failure.
3.3.4.1. Uniaxial Tension Testing of PP Monofilaments
PP monofilament sutures were tested in uniaxial tension to examine effects of plasma exposure on mechanical properties of mesh implants at the level of their most basic component – the individual filaments. A total of three Prolene 4-0 monofilaments were removed from sterile packaging and cut in half, with each half being assigned to a 0 min
118 or 20 min plasma treatment (n = 3 samples per plasma treatment duration). Each monofilament half was then cut into 7 pieces (technical replicates) of 2.5” (63.5 mm) length, such that a total of 21 pieces were tested per treatment group. Pieces were gripped on each end (wrapped in tape to reduce risk of grip failure) to a depth of 0.75” (19 mm) to leave a gauge length of 1” (25.4 mm), and tested in uniaxial tension at 10 mm/min until failure. All samples were observed to rupture between the grips. An Instron 10 lbf (44.5 N) tension load cell (cat #2512-111) and a sampling rate of 100 Hz were used. Initial stiffness values were determined by performing linear regression on the elastic region of load- displacement curves from 3 N to 5 N load, and slope of the best-fit line was taken to be stiffness (all R^2 values exceeded 0.999). Data presented reflects 3 samples per plasma treatment duration.
3.3.4.2. Uniaxial Tension Testing of PP Meshes
Unlike individual filaments, meshes tend to exhibit a complex geometric structure, so were tested in uniaxial tension to examine effects of plasma exposure on mechanical properties along two orthogonal axes. Methods for uniaxial tension testing were adapted from Deeken et al. 264,738, with slight modification. Meshes were die-cut into ASTM D412
Type A dogbone shaped samples along both the longitudinal and transverse directions.
Samples were gripped on each end (wrapped in tape to reduce risk of grip failure) to a depth of 1” (25.4 mm) to leave a gauge length of 3.5” (88.9 mm). Uniaxial tension testing was performed at 50 mm/min until mesh failure. All samples were observed to rupture within the gauge region. Initial stiffness values were determined by performing linear regression on the elastic region of load-displacement curves from 10 mm to 15 mm displacement, and slope of the best-fit line was taken to be stiffness (all R^2 values
119 exceeded 0.92). Data presented reflects 4-7 longitudinal mesh samples per plasma treatment duration, and 4-5 transverse samples per plasma treatment duration. A typical mesh uniaxial tension test (at 10X speed) can be viewed in the supplementary material of the published article 642.
3.3.4.3. Suture Retention Testing of PP Meshes
Suture retention testing was performed to simulate a loading scenario which might be anticipated in clinical use, specifically force applied to a mesh through a fixation point or “suture” that passes through it, which in this case is represented by a stainless steel wire to ensure that failure is localized to the mesh and not the fixation element. Methods for suture retention testing were adapted from Deeken et al. 264,738. A custom test fixture was machined that consisted of a 0.014” (0.36 mm) diameter stainless steel wire that could be clamped in place on an aluminum frame, after passing the wire through a mesh 1 cm from its bottom edge. Mesh specimens were cut to 1” x 2” (25.4 mm x 50.8 mm). Meshes were installed with a gauge length of 1” (25.4 mm) and clamped at the top end (wrapped in tape to reduce risk of grip failure) using pneumatic grips. Each mesh specimen was tested in tension at a rate of 300 mm/min until the suture pulled through the mesh. The “pull-out force” was defined as the failure load (N). A sampling rate of 200 Hz was used. Meshes were oriented such that the direction of pull was parallel to the longitudinal axis. Data presented reflects 4-10 mesh samples per plasma treatment duration. A typical mesh suture retention test can be viewed in the supplementary material of the published article 642.
120 3.3.4.4. Tear Resistance Testing of PP Meshes
Tear resistance testing was performed to simulate the ability of a mesh to resist further tearing once a tear has already begun. Methods for tear resistance testing were adapted from Deeken et al. 264,738. Specimens were cut to 1” x 2” (25.4 mm x 50.8 mm). A
1” (25.4 mm) slit was cut parallel to the longitudinal axis from the middle of the short edge toward the center of the mesh to form 2 tabs or “pant legs.” The left and right tabs (wrapped in tape to reduce risk of grip failure) were clamped to a ½” (12.7 mm) depth in the upper and lower grips, respectively. This arrangement yielded a 1” (25.4 mm) gauge length. The test was conducted at a rate of 300 mm/min until the specimen tore in half. The “tear force” was defined as the failure load (N). A sampling rate of 200 Hz was used. Meshes were oriented such that the direction of pull was parallel to the mesh longitudinal axis. Data presented reflects 6-8 mesh samples per plasma treatment duration. A typical mesh tear resistance test can be viewed in the supplementary material of the published article 642.
3.3.4.5. Ball Burst Testing of PP Meshes
Ball burst testing was performed to simulate a biaxial loading scenario which meshes might be anticipated to experience in clinical use (e.g. during coughing or sneezing). Methods for ball burst testing were adapted from Deeken et al. 264,738 and ASTM
D3787-16 741. Fixtures were machined that consisted of a polished stainless steel ball having a 1” (25.4 mm) diameter for the top fixture, and an aluminum ring clamp having a
1.75” (44.5 mm) internal diameter with neoprene-lined faces to prevent mesh slippage for the bottom fixture. Specimens were cut to 2.75” x 2.75” (70 mm x 70 mm). The steel ball was pressed transversely through the mesh at a rate of 300 mm/min until the mesh burst.
The “burst force” was defined as the failure load (N). Circumference at failure refers to the
121 circumference (cm) of the contact zone that the mesh made with the steel ball at the instant of failure. The circumference at failure for each specimen was determined as follows: 1) the ball burst test geometry was first modeled over a comprehensive range of experimentally-relevant crosshead displacements using Solidworks 2017 Student Edition
(Dassault Systèmes SolidWorks Corporation), 2) a polynomial regression equation was derived from Solidworks model results to relate displacement to failure to circumference at failure, and 3) circumference at failure was calculated for each specimen using the regression equation based on the specimen’s known displacement to failure. “Burst strength” refers to the maximum spherical wall tension sustained by the mesh and was calculated by dividing the burst force by the circumference at failure. A Material Testing
System MTS 810 frame (MTS Systems Corporation) was used along with a 2000 lbf (8900
N) load cell. Data presented reflects 5-6 mesh samples per plasma treatment duration. A typical mesh ball burst test can be viewed in the supplementary material of the published article 642.
3.3.5. Statistical Analysis
All data is presented as mean ± standard deviation. Statistical analysis tests were carried out in Microsoft Excel 2016 using two-sample two-tailed Student’s t-tests with unequal variance, with statistical significance being set at p < α = 0.05. A Bonferroni correction was applied in fibrinogen adsorption and bacterial attachment experiments to set statistical significance at p-values less than α divided by 15, the number of Student’s t- test comparisons performed among the 6 conditions (i.e. unique plasma treatment durations). Data points in any individual group data set that fell more than 2 standard
122 deviations away from the mean of that data set were considered outliers and excluded from analyses.
3.4. Results
3.4.1. Surface Characterization
Plasma-induced effects on PP material surfaces were evaluated in terms of surface chemistry, and attachment of proteins, bacteria, and mammalian cells.
3.4.1.1. X-ray Photoelectron Spectroscopy of PP Meshes
XPS was used to follow the extent of changes in PP surface chemistry in response to plasma exposure time. Analysis of XPS survey scan spectra (Figure 3-2) indicated that plasma enhanced the atomic percentage of oxygen on PP substrate surfaces in a duration- dependent manner (Figure 3-3) with little apparent impact on surface nitrogen content. For the untreated substrate, trace amounts of fluorine were detected on the mesh surface, possibly derived from device packaging or the textile production process. The high- resolution C1s spectrum was nearly symmetric for the untreated mesh surface, while spectra were skewed left following plasma treatment for any duration (Figure 3-4). High- resolution spectra also showed an increasing signal at ~289 eV with longer plasma exposures, attributed to carboxylic acids and/or carbonates on the surface, species that can only be produced through PP chain scission.
123
Figure 3-2. XPS survey scans, normalized by area under the curve, for 0 min and 20 min plasma treatment durations on Prolene Soft mesh. KLL peaks indicate atomic relaxation via the Auger effect, and were not used for analysis.
Figure 3-3. Quantification of XPS survey scans to determine the atomic composition of mesh surfaces (n = 1 mesh/group). Data is fit with an exponential growth curve.
124
Figure 3-4. XPS high-res C1s scans, normalized by area under the curve, demonstrated increasing signal at ~289 eV with increasing treatment duration.
125 3.4.1.2. Fibrinogen Adsorption
Fibrinogen was used as a model protein to evaluate changes in protein adsorption to PP surfaces following plasma exposure. Plasma treatment for any length of time reduced fibrinogen adsorption by roughly half (p < 0.001) relative to untreated PP (Figure 3-5).
Increasing the length of plasma treatment time beyond 1 min resulted in diminishing resistance to fibrinogen adsorption (p < 0.002). For reference, the average rate of fibrinogen adsorption to tissue culture polystyrene in the same experiment was 10.2 ± 0.8%.
Figure 3-5. Effect of plasma treatment duration on FITC-fibrinogen protein adsorption. The initially seeded amount (i.e. 100% of seeded) represents 15,625 ng fibrinogen per square cm. ♢Significant difference to all other time points.
126 3.4.1.3. Bacterial Attachment
Materials were exposed to E. coli to evaluate the capacity for plasma treatment to reduce bacterial attachment to PP surfaces. Plasma treatment for any length of time reduced attachment of E. coli (p < 0.001) by a proportion of at least 75% relative to untreated PP
(Figure 3-6). Bacterial attachment appeared to level off after 1 min of plasma treatment (p
> 0.048), until attachment rates of roughly half that seen at 1 min (p < 0.001) were seen at both 10 min and 20 min of plasma treatment. For reference, the average rate of E. coli attachment to tissue culture polystyrene in the same experiment was 2.36 ± 0.59%.
Figure 3-6. Effect of plasma treatment duration on E. coli bacterial attachment. ♢Significant difference to all other time points.
127 3.4.1.4. Mammalian Fibroblast Attachment
Substrates were exposed to fibroblasts to evaluate the capacity for plasma treatment to enhance mammalian cell attachment to PP surfaces. Plasma treatment for any length of time increased attachment of NIH/3T3 fibroblasts (p < 0.018) relative to untreated PP
(Figure 3-7). Fibroblast attachment to untreated PP was not detectable. Increasing the length of plasma treatment time beyond 1 min did not result in any significant improvements in fibroblast attachment (p > 0.1). For reference, the average rate of
NIH/3T3 attachment to tissue culture polystyrene in the same experiment was 64.3 ± 7.4%.
Figure 3-7. Effect of plasma treatment duration on mammalian fibroblast attachment. The initially seeded amount (i.e. 100% of seeded) represents 30,000 cells per square cm. ♢Significant difference to all other time points.
128 3.4.2. Bulk Characterization
A variety of mechanical tests were performed to assess plasma-induced bulk property changes of PP biomedical textiles. The tests used for evaluating these effects on textile implant structural properties are depicted in Figure 3-8.
Figure 3-8. Tests performed to assess the effects of plasma on material and structural properties of PP surgical meshes: A) uniaxial tension tests on PP monofilaments, and B) uniaxial tension, C) suture retention, D) tear- resistance, and E) ball burst tests on PP meshes.
129 3.4.2.1. PP Monofilament Uniaxial Tension Testing
PP monofilament sutures were tested in tension to explore the effects of plasma exposure on bulk property changes of mesh implants at the level of their simplest component. Plasma treatment for 20 min caused an 11% decrease in load to failure (p =
0.046), a 43% decline in work to failure (p = 0.01), and a 32% reduction in displacement to failure (p = 0.013), but had no effect on initial stiffness (p > 0.5) for Prolene monofilaments (Table 3-1, Figure 3-9).
Table 3-1. Effects of plasma exposure time on PP monofilament tensile properties. *Significant difference to untreated.
Plasma Monofilament Uniaxial Tension Duration (min) Load to Work to Failure Displacement to Initial Stiffness Failure (N) (mJ) Failure (mm) (N/mm)
0 16.70 ± 0.22 234.79 ± 18.62 20.09 ± 1.41 1.62 ± 0.05
20 14.80 ± 0.80* 133.13 ± 28.42* 13.72 ± 1.96* 1.60 ± 0.01
Figure 3-9. Plasma treatment resulted in embrittlement of Prolene 4-0 monofilaments, as depicted by representative (closest to average values) load-displacement curves.
130 3.4.2.2. PP Mesh Uniaxial Tension Testing
Meshes were tested in uniaxial tension to evaluate effects of plasma exposure on textile implant mechanical properties along two orthogonal axes (Figure 3-10), referred to here and in prior literature 742,743 as longitudinal and transverse. Plasma treatment decreased failure load (p < 0.001), work to failure (p < 0.006), and displacement to failure (p < 0.001) of Prolene mesh in the longitudinal direction (Figure 3-11A, Table 3-2). Additionally, plasma treatment increased mesh longitudinal initial stiffness for all durations (p < 0.02) except 20 min (p = 0.07). Longer treatments were associated with earlier failure of meshes in the longitudinal direction. For example, mesh plasma treatment for 20 min caused a 41% decrease in failure load, a 27% decrease in failure displacement, and a 54% decrease in work to failure. While the declines in failure displacement are comparable, the decreases in failure load and work to failure of meshes are proportionally greater than those observed for monofilaments treated for the same length of time. Additionally, with the exception of a ~5% decrease in displacement to failure for all durations (p < 0.03) and a ~12% decrease in work to failure at 20 min (p = 0.048), there were otherwise no apparent effects of plasma on Prolene mesh failure load (p > 0.3), work to failure (p > 0.1), or initial stiffness (p > 0.1) in the transverse direction (Figure 3-11B, Table 3-3). In agreement with prior literature, untreated Prolene mesh demonstrated higher stiffness, higher load to failure, and lower displacement to failure in the longitudinal direction than in the transverse direction 743. The relatively higher initial stiffness of Prolene mesh in the longitudinal axis highlights the important effect filament alignment has on the mesh mechanical behavior early in the loading process. Altogether, these results suggest that plasma-induced effects on textile bulk mechanics are influenced by structure (e.g. the geometry and orientation of the knit)
131 in addition to material properties, with effects being more drastic for the predominant load- bearing axis.
Figure 3-10. Longitudinal and transverse orientations, defined as being parallel to the direction of crosshead travel for uniaxial tension tests on Prolene meshes. Scale bar = 1 mm.
A B
Figure 3-11. Representative load-displacement curves from uniaxial tension tests of Prolene mesh in the A) longitudinal direction, and B) transverse direction.
132 Table 3-2. Effects of plasma exposure time on mesh longitudinal mechanics. *Significant difference to untreated. ⁑Significant difference to all subsequent time points.
Plasma Longitudinal Mesh Uniaxial Tension Duration (min) Load to Failure Work to Displacement to Initial Stiffness (N) Failure (J) Failure (mm) (N/mm)
0 134.94 ± 4.75⁑ 2.69 ± 0.21⁑ 59.87 ± 2.23⁑ 0.47 ± 0.06
1 118.62 ± 5.75 2.29 ± 0.20 53.88 ± 2.60 0.57 ± 0.05*
2.5 108.18 ± 6.34 1.95 ± 0.18 50.84 ± 1.85 0.57 ± 0.04*
5 102.99 ± 3.87 1.80 ± 0.09 49.62 ± 1.06 0.57 ± 0.04*
10 94.25 ± 6.71 1.54 ± 0.17 47.52 ± 0.91 0.59 ± 0.06*
20 80.25 ± 4.46 1.24 ± 0.08 43.95 ± 0.74 0.52 ± 0.02
Table 3-3. Effects of plasma exposure time on mesh transverse mechanics. *Significant difference to untreated. ⁑Significant difference to all subsequent time points.
Plasma Transverse Mesh Uniaxial Tension Duration (min) Load to Work to Failure Displacement to Initial Stiffness Failure (N) (J) Failure (mm) (N/mm)
0 73.60 ± 4.11 2.18 ± 0.14 101.43 ± 1.54⁑ 0.06 ± 0.02
1 75.63 ± 4.14 2.13 ± 0.18 96.44 ± 2.76 0.06 ± 0.01
2.5 77.11 ± 4.94 2.22 ± 0.20 96.51 ± 2.09 0.08 ± 0.01
5 74.11 ± 3.44 2.02 ± 0.13 93.79 ± 2.48 0.07 ± 0.01
10 72.08 ± 6.04 1.93 ± 0.22 91.91 ± 3.65 0.08 ± 0.01
20 72.47 ± 3.55 1.92 ± 0.18* 93.89 ± 4.96 0.08 ± 0.02
133 3.4.2.3. Mesh Suture Retention Testing
Suture retention testing was conducted to examine the impact of plasma treatment on the ability of mesh samples to withstand loading through a single fixation point. Plasma treatment for 20 min was observed to result in a 17% reduction in pull-out force (p = 0.006) compared to untreated meshes (Figure 3-12). Representative load-displacement curves are shown in Figure 3-12b, where each peak represents the rupture of an individual (or group of) filament(s) as the wire breaks through. No other treatment duration showed a significant difference to untreated meshes in pull-out force (p > 0.08). Note that displacement and work to failure are not considered precise metrics in this test, given that the wire may have captured different numbers of filaments to break through on different samples, depending on how the mesh was cut and where in the knit pattern the wire was inserted. The longitudinal suture retention strength of 58.9±7.5 N measured in this study for untreated
Prolene is in agreement with the 61.2 N previously reported in literature 742.
A B
Figure 3-12. Mesh suture retention tests in the longitudinal direction showing effect of plasma duration on A) maximum pull-out force, and B) representative load-displacement curves. *Represents significant difference to untreated.
134 3.4.2.4. Mesh Tear-Resistance Testing
Tear-resistance testing was carried out to explore the effect of plasma treatment on the resistance of mesh samples to tear propagation once a tear has already been initiated.
Plasma treatment had no discernible impact (p > 0.2) on mesh tear force (Figure 3-13).
Representative load-displacement curves are shown in Figure 3-13B, where each peak represents breakage of an individual (or group of) filament(s). Results from tear-resistance testing demonstrated high variability in terms of tear force due to the lack of consistency in how the tear propagated through the mesh among different trials. For example, it was observed that the tear may propagate and exit through either the side (long edge) or center
(short edge) of the rectangular specimen. These tear propagation directions are equivalent to those seen in the longitudinal and transverse tensile tests, respectively, given that during tensile tests the tear propagates perpendicular to the direction of pull. Interestingly, a higher proportion of side failures were observed for plasma-treated meshes at all time points than untreated meshes (Table 3-4), and this is consistent with the finding that plasma treatments preferentially embrittled meshes tested in the longitudinal direction. At any rate, attempting to account for the different tear propagation patterns would reduce sample size by different proportions among groups, therefore any effects of plasma treatment on tear force may be obscured by this variability. The longitudinal tear force of 27.8±12.8 N measured in this study for untreated Prolene is in reasonable agreement with the 33.66 N previously reported in literature 742.
135 A B
Figure 3-13. Effects of plasma treatment duration on A) maximum load sustained in PP mesh tear-resistance tests. B) Representative load-displacement curves.
Table 3-4. Failure patterns for tear-resistance testing among samples after plasma treatments. Plasma Duration n Side Failures Center Failures (min)
0 7 1 6
1 7 5 2
2.5 7 4 3
5 6 4 2
10 6 4 2
20 8 5 3
136 3.4.2.5. Mesh Ball Burst Testing
Ball burst testing was conducted to examine the impact of plasma treatment on the ability of meshes to withstand biaxial loading that might be encountered during changes in intra-abdominal pressure in vivo. For ball burst tests, plasma treatments for 10 and 20 min caused significant declines in burst force (p < 0.006), displacement to failure (p < 0.02), and work to failure (p < 0.02) compared to untreated (Figure 3-14, Table 3-5). In particular, after 20 min of treatment, burst force diminished by 19%, work to failure decreased by 39%, and displacement to failure declined by 13%. Interestingly, this proportional decrease in work to failure is roughly equal to that seen for Prolene monofilaments. No significant differences to untreated meshes were observed for earlier time points (p > 0.09). The burst strength measured in this study for untreated Prolene is slightly less than the 156.6 N previously reported in literature 742. This discrepancy could perhaps be due to minor procedural differences in determination of displacement and subsequent calculation of circumference at burst.
Figure 3-14. Representative load-displacement curves for ball-burst tests. 137 Table 3-5. Effects of plasma treatment duration on ball burst properties. *Significant difference to untreated. Plasma Displacement Work to Circumference Burst Strength Duration Burst Force (N) to Failure Failure (J) at Burst (cm) (N/cm) (min) (mm)
0 847.03 ± 72.62 6.31 ± 1.30 23.06 ± 1.57 6.62 ± 0.21 127.71 ± 7.76
1 838.78 ± 100.38 6.05 ± 1.12 22.93 ± 1.16 6.61 ± 0.16 126.66 ± 12.47
2.5 796.21 ± 82.78 5.28 ± 1.04 21.89 ± 1.14 6.45 ± 0.19 123.21 ± 9.33
5 773.19 ± 86.70 5.04 ± 0.93 21.56 ± 1.12 6.42 ± 0.17 120.22 ± 10.24
10 718.89 ± 28.43* 4.44 ± 0.33* 20.80 ± 0.47* 6.30 ± 0.07* 114.03 ± 3.50*
20 689.21 ± 30.32* 3.86 ± 0.28* 20.05 ± 0.42* 6.18 ± 0.07* 111.40 ± 3.58*
3.5. Discussion
These experiments were designed to test the hypothesis that plasma treatment alters biological fouling by proteins, cells, and bacteria, but increasing the duration of exposure accelerates the mechanical failure of PP biomedical textiles. Findings regarding fibrinogen, fibroblast, and E. coli attachment, as well as findings from uniaxial tension tests of monofilaments (sutures), and longitudinal tension, suture retention, and ball burst tests of meshes support this hypothesis. Effects of plasma duration on mesh transverse tension and tear resistance properties were less severe, however both of these tests emphasized failure in the mesh axis that bears less load. We can state with certainty, however, that while there were beneficial effects on protein, cell, and bacterial attachment, the nonthermal plasma in this study had non-negligible effects on the bulk properties of biomedical textiles.
138 Additionally, the plasma treatment parameters used here are similar to those of plasmas applied towards surface modification of biomaterials currently used clinically, or being investigated for such use. Power outputs between 40 W and 200 W, treatment durations between 30 s and 10 min, and carrier gases such as Ar, N2, O2, and NH3 have been reported for plasma treatments of surgical meshes 676,710,713,714,716,719,723,724.
There are several possible explanations for the demonstrated embrittlement effect of plasma on PP. As demonstrated by XPS high-resolution C1s scans, longer plasma treatments appeared to result in an increase in the proportion of carboxylic acids and carbonates on the PP surface. Considering that the tertiary carbon atom in the PP repeat unit has the highest radical stability (hence is the most likely reaction site), these oxygen- rich functional groups can only be created through some degree of chain scission at the polymer surface. These degraded molecular chains at the surface may act like nano-scale cracks, increasing in size and number with longer plasma exposure. Additionally, while changes in the plasma-treated PP XPS spectra between 1 and 20 min may have appeared less drastic than those seen for certain mesh mechanical properties, the calculated sampling depth of XPS was only ~7 nm so plasma-induced crack growth and etching would likely not be reflected completely in the spectra. Though very small, such cracks could propagate into the bulk of the material upon application of sufficient mechanical stress. In support of this reasoning, prior studies have demonstrated that plasma treatments roughen the topography of polymer surfaces 676,730. Enhanced surface roughness would also explain the increase in Prolene mesh initial stiffness seen in tensile tests – surface irregularities could reduce the ability of individual filaments to slide smoothly across one another, resulting in a faster increase in force for a given displacement.
139 There have been several studies on hernia mesh mechanical properties 264,738,743–745.
In considering the mechanical properties of untreated Prolene mesh, results from this research were largely in agreement with those of prior studies, with few minor exceptions.
For example, prior reports have indicated a 20x higher uniaxial tensile strength in the transverse or “perpendicular” direction than in the longitudinal or “parallel” direction for untreated Prolene mesh 738,742. We believe this discrepancy may result from differences in the dimensions of dogbone-shaped test specimens (the gauge region of their samples appeared narrower, thus capturing very few mesh interstices across its width), and perhaps in the naming convention used to refer to the different mesh axes.
Though there are few definitive guidelines for the mechanical properties that a surgical mesh should possess, it is important to be aware of the detrimental effects that plasma treatments may have on hernia mesh bulk mechanical properties, and to try to minimize these through careful selection of plasma processing parameters. Based on our results, tension tests along the predominant load-bearing direction and ball burst tests would appear to be the most sensitive assays at detecting detrimental effects of plasma treatment on a given mesh fabric. The test metrics which are arguably most relevant to hernia repair would be suture retention and ball burst testing, as mesh devices in vivo are most likely to experience loading that is biaxial in nature, with stresses concentrated at the fixation elements. The guidelines that exist regarding mesh mechanical properties recommend a suture retention strength that exceeds 20 N and a burst pressure of at least 50
N/cm 742. Although plasma treatments for up to 20 min did not result in clinically meaningful embrittlement of Prolene mesh to the extent that it fell short of either of these specific failure values, a PP mesh having lower baseline mechanical properties (e.g.
140 Prolene Soft mesh) or higher surface-to-volume ratio could potentially have been weakened to such a degree.
A notable finding from this study was that plasma treatments of PP for all durations tested significantly reduced fibrinogen adsorption and E. coli attachment relative to untreated surfaces, and increasing exposure time had diminishing returns on this passive repellence. Likewise, exposure to plasma for any length of time significantly enhanced fibroblast attachment relative to untreated surfaces, while increasing plasma duration had no additive effect on this attachment. Hydrophobic materials such as untreated PP have a strong tendency to adsorb and unfold (i.e. denature) protein solutes. It is energetically favorable for proteins to displace water molecules repelled by a hydrophobic surface and then change conformation such that core nonpolar segments can associate tightly with the surface via nonpolar interactions. Proteins adsorbed on the material then serve as a conditioning layer requisite for further biofouling events, such as bacterial or cell attachment. Plasma treatment renders the surface more hydrophilic, thereby altering the concentrations and conformations of the adsorbed proteins, and modulating subsequent biofouling events. It has previously been shown that plasma treatment decreases protein adsorption of both albumin and fibrinogen to PP 656. But to our knowledge, no studies have assessed the effects of plasma exposure time on protein, bacterial, or cell resistance of PP.
Our findings regarding protein and bacterial resistance of plasma-treated PP suggest that plasma treatment may point to a strategy to reduce the likelihood or severity of adhesion formation and prosthetic infection on PP mesh surfaces. At the same time, our findings regarding fibroblast attachment on plasma-treated PP suggest that plasma treatment may promote mesh incorporation with the abdominal wall. Finally, our findings suggest that
141 long exposures (i.e. beyond 1 minute using the materials and system described here), with their associated decline in bulk mechanical properties, are not necessary to realize these beneficial effects.
This study sets the stage for several possible additional investigations. Future studies are needed to determine whether the effects of plasma seen in these studies on PP meshes extend to meshes that exhibit different geometries, or meshes made from other polymeric materials such as polyester or poly(tetrafluoroethylene). In particular, meshes that exhibit a higher surface-to-volume ratio, or meshes that are more lightweight and flexible at baseline should be considered, as plasma-induced embrittlement of these could pose more dangerous consequences. Prolene mesh is a relatively rigid, heavyweight mesh with high baseline mechanical properties 738, and it is considered to maintain its strength indefinitely in clinical use according to manufacturer product literature. It was chosen as a representative mesh for several reasons: 1) its homogenous composition of monofilament
PP, the most common mesh material 46–50, 2) its longstanding clinical use in abdominal surgery for over 4 decades, and 3) its constituent Prolene material is used in several Ethicon surgical products (including Prolene sutures, Prolene Soft mesh, and Proceed composite mesh) that span a wide variety of surgical purposes. Additional studies on plasma-treated meshes should also explore fatigue properties, which are more physiologically relevant than monotonic tests given the periodic nature of loading in vivo (e.g. breathing, coughing, jumping, etc.), to evaluate whether plasma treatments predispose meshes to failure after repetitive movements. Further investigations should also consider the effects of plasma- treated mesh storage time and environment on biofouling resistance and bulk properties; hydrophobic recovery or “aging” of plasma-treated PP surfaces normally occurs over
142 several weeks in the absence of a coating 701,703, but whether protein/cell/bacterial attachment or mechanical properties might also recover is uncertain and worth studying.
Finally, the effects of plasma treatment on tissue adhesion formation and prosthetic infection of surgical textiles could also be investigated in vivo.
3.6. Conclusions
These experiments sought to systematically quantify the positive and negative effects of nonthermal plasma on polymeric implants. While plasma treatment for any length of time increased surface oxygen content, and reduced fibrinogen adsorption and E. coli attachment on PP surfaces, increasing the length of exposure resulted in diminishing benefit with regards to surface chemistry, protein and bacterial resistance, and mammalian cell attachment. Simultaneously, plasma treatments were found to result in bulk embrittlement of PP biomedical textiles across several standardized testing modalities, with the duration of exposure being correlated to the decline in mechanical properties.
Taken together, the results from this study indicate that plasma-based treatments of PP surgical meshes should be optimized for both surface qualities and implant structural properties (if they are of chief importance), and the effects of a plasma process on a given biomedical material at both the surface and bulk levels should be carefully evaluated before translation to any clinical scenario.
143 3.7. Acknowledgments
The authors acknowledge support through National Institutes of Health: NIH R01
GM121477 (HvR), and NIH Ruth L. Kirschstein NRSA T32 AR007505 Training Program in Musculoskeletal Research (GDL). Additional support was provided by the Center for
Stem Cell and Regenerative Medicine Undergraduate Student Summer Program
(ENGAGE) at Case Western Reserve University (EJL). Valuable core facility services were provided by the Swagelok Center for Surface Analysis of Materials, the Advanced
Manufacturing and Mechanical Reliability Center, and Think[box] at CWRU. The authors also thank Kevin Abbassi for expertise with operation of XPS, Chris Tuma for expert assistance with operation of mechanical testing equipment, Katherine Yan for technical help, and Nathan Rohner and Alan Dogan for revision suggestions.
144 4. CHAPTER 4: CYCLODEXTRIN POLYMER COATINGS RESIST FOULING
BY PROTEINS, MAMMALIAN CELLS, AND BACTERIA
This chapter was adapted from a bioRxiv preprint 638. No permissions are needed.
Authors: Greg D. Learn, Emerson J. Lai, Horst A. von Recum
4.1. Abstract
Undesired attachment of proteins, cells/bacteria, and organisms on material surfaces is problematic in industrial and health care settings. In this study, polymer coatings are synthesized from subunits of cyclodextrin, an additive/excipient found in food/pharmaceutical formulations. These unique polymers, which have been applied mainly towards sustained drug delivery applications, are evaluated in this study for their ability to mitigate non-specific protein adsorption, mammalian cell (NIH/3T3) adhesion, and bacterial cell (Staphylococcus aureus, Escherichia coli) attachment. Effects of cyclodextrin polymer composition, particularly incorporation of nonpolar crosslinks, on material properties and passive anti-biofouling performance are investigated. Results suggest that lightly-crosslinked cyclodextrin polymers possess excellent passive resistance to protein, cell, and bacterial attachment, likely due to the hydrophilic and electrically neutral surface properties of these coatings. At the same time, anti-biofouling performance decreased with increasing crosslink ratios, possibly a reflection of decreased polymer mobility, increased rigidity, and increased hydrophobic character. Cyclodextrin-based materials may be broadly useful as coatings in industrial or medical applications where biofouling-resistant and/or drug-delivering surfaces are required.
145 4.2. Introduction
Biological fouling, or “biofouling,” is the undesired accumulation of biological contaminants (biofoulants) on a material surface, particularly at an aqueous liquid/solid interface. Biofouling poses major challenges in the health care 746, water treatment 651, and marine industries 747, among others 748,749. For example, medical implants are susceptible to uncontrolled surface accumulation of proteins, cells, and bacterial biofilms, resulting in serious complications such as foreign body response 750,751, thrombosis 752,753, and prosthetic infection 746,754,755. Water purification and desalination membranes are vulnerable to biofilm colonization that reduces flux and contaminates treated water
650,651,756. Buildup of aquatic life on ship hulls accelerates corrosion, reduces vessel maneuverability/speed, raises fuel consumption, promotes invasive species migration, and necessitates periodic dry-docking maintenance 747,757–759.
Biofouling is a cumulative process that often starts with non-specific protein adsorption 684,760,761. Immediately upon exposure to a bare surface, protein solutes adsorb in an equilibrium determined by concentration, diffusivity, and affinity 760. Small, abundant proteins predominate on the surface over short time frames, then are gradually replaced by higher-affinity proteins, a process known as the Vroman effect 762. With few exceptions, proteins adsorb onto surfaces in a near-monolayer arrangement 760. This “conditioning film” then mediates further biofouling events – typically cell adhesion, then matrix/biofilm formation as cells deposit new proteins, then macro-organism attachment (in the case of marine biofouling) – dependent on the types, concentrations, and conformations of the adsorbed molecules 758,763–766. For this reason, surfaces that prevent protein adsorption are theorized to block cell adhesion 222,758. Conversely, increased adsorption is typically
146 expected to enhance cell adhesion 760,767. However, many findings challenge this correlation 768,769, indicating the importance of testing materials in biofouling scenarios that recapitulate conditions of intended use.
Anti-biofouling (ABF) materials possess surfaces that resist accumulation of proteins, cells, and/or organisms 222,770, thus reducing consequences linked to excessive buildup of biological material on critical interfaces. ABF has typically been achieved by two major strategies: active ABF and passive ABF. These strategies are based on degrading adherent biofoulants or preventing their attachment, respectively 752,764.
Active ABF approaches use biocidal agents presented at, or released from, the material surface to injure or destroy any cells or organisms that stick 771. Typical agents presented at medical implant surfaces include silver compounds and antibiotic drugs.
Biocides released from marine surfaces include toxic metal compounds 757,759,772–774. Key limitations of active ABF strategies include: 1.) potential for off-target effects detrimental to surroundings, and 2.) transient effectiveness, given the finite biocide reservoir and the propensity for target micro-organisms to develop resistance. Additionally, depending on the biodiversity encountered, the biocidal agents chosen may not provide a sufficiently broad effect to resist biofouling by all relevant organisms.
Passive ABF approaches utilize anti-adhesive material surfaces to reduce the ability of biological contaminants to physically settle and adhere. Upon wetting, the foulability of a surface is influenced by numerous factors, including fluid movement near the interface, the area of surface exposed, the duration of biofoulant exposure, and the surface chemistry, charge, wettability, stiffness, and topography 756,757,765,775. Surfaces that resist biofouling typically do so by minimizing contact (e.g. by trapping an air layer at the surface to
147 minimize wetted area, or by attracting a water layer that sterically hinders protein adsorption) and attractive forces (e.g. electrostatic) between biofoulants and the surface.
The wettability of a surface plays a large role in the onset of protein adsorption.
Specifically, hydrophobic materials have a strong tendency to adsorb and unfold (i.e. denature) protein solutes 222–224,756,776–778. It is energetically favorable for proteins to displace polar water molecules at a hydrophobic surface, and further to change conformation such that core nonpolar amino acid segments are exposed and able to associate with the surface via nonpolar interactions. Even superhydrophobic substrates, which are hydrophobic materials that possess a nano-structured surface that traps air to minimize wetted area and biofoulant exposure, are prone to irreversible wetting and subsequent biofouling with extended submersion in protein solutions 682,759,779–781. Beyond serving as a conditioning layer for further biofouling, denatured proteins on a hydrophobic medical device may elicit thrombotic, fibrotic, or inflammatory responses that limit the device’s biocompatibility 220,229,231,240,778,782–785.
Conversely, hydrophilic material surfaces attract water molecules that sterically hinder protein adsorption while minimizing the chance of protein denaturation. Pioneering studies have indicated that electrically-neutral hydrophilic surfaces, especially those presenting abundant H-bond acceptors without H-bond donors, tend to be the most protein- resistant 786. Overall charge neutrality minimizes electrostatic interactions that could otherwise attract proteins (or cells) to a surface 787,788. The importance of lacking H-bond donors is less obvious 789 as glycocalyx-mimetic carbohydrate surfaces (which display many H-bond donors) also effectively resist non-specific protein adsorption 790–792. It is clear, however, that water held at the surface through H-bonding is critical for resistance
148 to protein adsorption. For this reason, neutrally-charged, H-bond-acceptor-rich, hydrophilic polymers, such as polyzwitterions 793, poly(ethylene glycol) (PEG) 794,795, poly(2-hydroxyethyl methacrylate) (pHEMA) 796, and many polysaccharides 790–792 present key advantages for passive ABF. The major limitation of passive ABF strategies is that their performance may be compromised by surface defects (e.g. cracking or delamination of passive ABF coating exposes underlying substrate to biofoulants).
Our group has previously studied polymers composed of cyclic oligosaccharides, in which cyclodextrin (CD) molecules are crosslinked together to form insoluble polymer networks 635. CDs are toroid-shaped molecules with a hydrophilic exterior and relatively hydrophobic core. They have long been used in the food and pharmaceutical industries, being Generally Recognized as Safe by the US FDA, to solubilize or prevent aggregation of nonpolar compounds in aqueous mixtures. CD polymers have principally been investigated for application as implantable drug delivery depots, based on the unique ability of CD subunits to reversibly bind and release drug compounds (e.g. antibiotics) through non-covalent (“affinity”) interactions 797. These affinity properties make CD polymers very well-suited for active ABF, because when compared to purely diffusion-based active ABF materials, CD polymers can release biocides over longer durations 635,636,640, and can be refilled more easily once the reservoir is depleted 637. Additionally, given their neutrally- charged, H-bond acceptor-rich composition of hydrated saccharide units, CD polymers share many similarities with glycocalyx-mimetic passive ABF materials. Furthermore, proteins such as albumin have demonstrated weak affinity for cyclodextrin subunits 798.
These observations imply a possibility that CD polymers might be useful for passive ABF in addition to active ABF, a unique property not seen in traditional ABF materials 799. This
149 potential has not been sufficiently explored and could be of value for many medical and industrial applications. Therefore, the objective of this work was to assess protein adsorption, and attachment of mammalian and bacterial cells on CD-based polymer surfaces.
An overview of this work is presented in Figure 4-1. In this study, polymerized CD
(pCD) was applied as a coating for polypropylene (PP), chosen as a model substrate material because of its susceptibility to uncontrolled biofouling in medical 709,800–803 and industrial 804–809 applications. Given the low surface energy of olefin polymers, PP substrates were treated with nonthermal plasma to enhance pCD coating uniformity and adherence 643. We hypothesized that pCD coatings would deter protein adsorption, cell adhesion, and bacterial attachment to PP substrates, in a manner dependent on crosslinking.
In this case, hexamethylene diisocyanate (HDI) was chosen as a crosslinker in order to allow investigation of neutrally-charged pCD materials having different overall hydrophobicity. Four different crosslinking formulations of pCD were first characterized in terms of material properties: swellability, rigidity, and wettability. Chemical composition of these formulations was also characterized using Fourier-transform infrared
(FTIR) spectroscopy, and the impact of antibiotic drug loading on pCD material properties was preliminarily examined. Next, the four pCD formulations were studied in terms of their passive ABF performance relative to bare PP and polystyrene (PS) control surfaces in terms of protein adsorption from bovine plasma, mammalian fibroblast adhesion and viability, and bacterial (S. aureus, E. coli) attachment.
150
Figure 4-1. Chapter 4 study overview. Effects of HDI crosslinking on pCD material properties (swellability, rigidity, and wettability) and resistance to biofouling (protein adsorption, fibroblast attachment, and bacterial adhesion) were explored.
4.3. Materials and Methods
4.3.1. Materials
Soluble, lightly epichlorohydrin-crosslinked β-CD polymer precursor (bCD) was purchased from CycloLab R&D (#CY-2009, batch CYL-4160, MW ~116 kDa; Budapest,
Hungary). HDI crosslinker (#52649), black PP 96-well plates (#M9685), pHEMA
(#P3932), and Pluronic F108 (#542342) were purchased from Sigma Aldrich. N,N- dimethylformamide (DMF) solvent (#D119-4), tissue culture PS (TCPS) 24-well plates
(#08-772-1H), non-TCPS 12-well plates (#08-772-50), Dulbecco’s Modified Eagle
Medium (DMEM) with high glucose (#11-995-040), fetal bovine serum (FBS) (#16-141-
079), penicillin-streptomycin (#15-140-122), Trypsin-EDTA (#25-200-056), LB broth
(#BP1426-500), and 1.5 mL tubes (#05-408-129) were purchased from Fisher Scientific.
BBL broth (#211768) and agar (#214010) was purchased from Becton Dickinson. PP 24- well plates (#1185U58) and lids (#1185U62) for fibroblast attachment were purchased from Thomas Scientific. Poly(tetrafluoroethylene) (PTFE) evaporating dishes with inner
151 diameters 63 mm (#355314-0025) and 30 mm (#355304-0025) were obtained from Lab
Depot. The antibiotic rifampicin (#R64000) was purchased from Research Products
International. Sterile bovine plasma with sodium heparin anti-coagulant (#IBV-N) was purchased from Innovative Research. PP sheet stock (#8742K133) for wettability and protein adsorption, and rod stock (#8658K51) for S. aureus attachment, were purchased from McMaster-Carr. Guava ViaCount reagent was purchased from Millipore Sigma
(#4000-0040). PP 4-0 Prolene blue suture (#8592G) was purchased from eSutures.
Bioluminescent ilux pGEX(-) E. coli was purchased from Addgene, a gift from Dr. Stefan
Hell (plasmid #107879) 739. Methicillin-resistant S. aureus strain Xen30 was purchased from Caliper Life Sciences. Tissue homogenizer (#TH-01) and blades (#30750H) were purchased from Omni International.
4.3.2. Plasma Cleaning and Activation of PP Substrate Surfaces
To maximize pCD coating adherence and stability, PP substrates were placed in a
4” diameter x 8” length quartz reaction chamber of a Branson/IPC Model #1005-248 Gas
Plasma Cleaner and treated with nonthermal plasma (500 mTorr, 50 W, 13.56 MHz) using
642,643 an inlet gas mixture of argon bubbled through water (Ar/H2O) . PP substrates were treated for a fixed 10 min duration within 1 h of pCD coating application or direct biofoulant exposure (for plasma-treated bare PP controls). Non-treated (0 min) PP samples without any known prior exposure to plasma or ultraviolet light were included as controls in all experiments.
152 4.3.3. pCD Synthesis and Coating onto Surfaces
In order to examine the impact of crosslinking on material properties and ABF performance, pCD was synthesized using HDI as a crosslinker for bCD at approximate crosslinker / glucose residue molar ratios of 0.08, 0.16, 0.32, and 0.64 (Table 4-1). bCD was weighed and placed in PP tubes, then DMF was added to dissolve it at 33% w/v. HDI was added to solutions to achieve the desired crosslink ratios, and pre-polymer mixtures were thoroughly vortexed, then cast either: (i) into clean PTFE dishes to produce free films of pCD for subsequent punching of disks for measurement of swelling ratio, elastic modulus, and contact angle, (ii) onto flat PP sheet/rod stock pieces (newly abraded using
1200, 2500, then 5000 grit SiC sandpaper to expose fresh surface) for preparing coated specimens for protein adsorption and S. aureus attachment, (iii) into wells of PP multiwell plates for production of coated well surfaces to be used in measurement of mammalian fibroblast adhesion/viability and in measurement of E. coli attachment. For coated well surfaces, coatings were applied to plates in a sterile biosafety cabinet, the volume of pre- polymer mixture added was 140 µL/well for 24-well plates, and 42 µL/well for 96-well plates, then plates were agitated to promote complete coverage, and surfaces were visually examined post-curing to exclude defective coatings prior to use. Cast pre-polymer mixtures were kept covered with Parafilm and typically allowed to cure for at least 4 days at ambient temperature and pressure. Cured pCD coatings were rinsed several times to terminate crosslinking, and stored immersed in sterile PBS (for subsequent cell/bacterial culture) or deionized water (for XPS) to keep samples hydrated before use.
153 Table 4-1. Pre-polymer mixtures for different formulations of pCD *HDI volumes (density 1.047 g/mL) are calculated based on estimation that epichlorohydrin linkages and water comprise a negligible weight fraction of bCD. pCD Formulation bCD (mg) DMF (µL) HDI (µL)*
0.08 1000 3000 80
0.16 1000 3000 160
0.32 1000 3000 320
0.64 1000 3000 640
4.3.4. Effects of HDI-Crosslinking on pCD Physicochemical Properties
4.3.4.1. Determination of pCD Swelling Ratio
Crosslinking density impacts the ability of network polymers to absorb water, and the degree of swelling may influence the biofouling resistance of ABF polymers 810. For this reason, effects of HDI crosslinking on pCD swelling ratio were examined. pCD pre- polymer mixtures (scaled to a DMF volume of 1 mL) were cast into 30 mm diameter PTFE evaporating dishes, covered, and allowed to cure for 7 weeks at ambient temperature and pressure. Disks were punched out 6.0 mm in diameter using a biopsy punch, extracted from dishes, and hydrated in deionized water for 3 days. Hydrated disks (n = 10-12 per group) were gently blotted on filter paper to remove surface water, and weighed after 60 seconds to the nearest 0.1 mg (wet weight) on an analytical balance (Mettler Toledo ME104TE).
Samples were then placed onto sheets of Parafilm and dried overnight under vacuum at ambient temperature, and weighed again to the nearest 0.1 mg (dry weight) on the same balance. Swelling ratio for each individual disk was calculated as the absorbed water weight (wet weight minus dry weight) divided by the dry weight. Results shown represent
154 findings from one experiment, with the same trends having also been observed in 4 similar independent experiments.
4.3.4.2. Unconfined Compression Testing of pCD
Crosslinking directly alters network polymer rigidity, and biomaterial stiffness is recognized as a factor that impacts adhesion, spreading, and phenotype of mammalian cells
811,812. Therefore, effects of HDI crosslinking on pCD mechanical properties were examined. pCD pre-polymer mixtures (scaled to a DMF volume of 4.5 mL) were cast into pristine 63 mm diameter PTFE evaporating dishes, covered, and allowed to cure for either
4 days or 4 weeks at ambient temperature and pressure. Disks were then punched out 8.0 mm in diameter using a biopsy punch (1.38-2.26 mm thickness, n = 5-9 per group) and placed in deionized water, with which specimens were kept hydrated before and throughout testing. For preparation of drug-loaded specimens, disks were incubated for 24 h in 5 mL
DMF containing 5% w/v rifampicin (RIF), then rinsed numerous times and stored for >24 h in beakers of deionized water. Specimen dimensions were carefully measured with digital calipers to the nearest 0.01 mm before each test and crosshead speed was adjusted accordingly to ensure a consistent strain rate of -0.005 s-1. Unconfined compression tests were performed on a Rheometrics RSA II Solids Analyzer (Rheometric Scientific) up to a load limit of 10 N (50 kPa) using a data acquisition rate of 4 Hz. Elastic modulus was determined through linear regression of the stress-strain curve over all possible 2.5% strain ranges within each given test prior to reaching the load limit (or the onset of yielding for some pCD 0.08 specimens), with the maximum slope found being defined as the modulus.
Results shown represent pooled findings from 2 independent experiments.
155 4.3.4.3. Contact Angle Goniometry
Because HDI crosslinks in pCD can be expected to be nonpolar and wettability is known to affect biofouling resistance, the effect of HDI crosslinking on pCD wettability was examined. pCD pre-polymer mixtures (scaled to a DMF volume of 4.5 mL) were cast into pristine 63 mm diameter PTFE evaporating dishes, covered, and allowed to cure for 4 days at ambient temperature and pressure. Disks were punched out 12.0 mm in diameter.
Specimens were kept hydrated with deionized water until contact angle measurement. For preparation of drug-loaded specimens, disks were incubated for 24 h in 5 mL DMF containing 5% w/v RIF, then rinsed numerous times and stored for >24 h in beakers of deionized water. Surfaces were evaluated for wettability by static contact angle measurement using a KSV Instruments CAM 200 Optical Contact Angle Meter. Sample surfaces were gently blotted to remove surface moisture on dry filter paper followed by
KimWipes, and then deionized water droplets of 8 µL volume were dispensed onto each sample surface and allowed to equilibrate for 30 s prior to photographing and measurement.
Two unique droplets per sample were measured for 3 separate samples per group, and the measurement for each droplet reflects the average of the angles on the left and right sides.
Measurements were performed using KSV CAM 2008 software. Results shown represent findings from one experiment, with the same trends having also been observed in 2 similar independent experiments.
4.3.4.4. Attenuated Total Reflectance FTIR Spectroscopy
FTIR was performed on pCD surfaces to confirm the extent to which reactants are consumed in polymerization and verify that HDI crosslink ratio is reflected in polymer composition. After cured pCD disks were punched and extracted from PTFE dishes, the
156 polymer film remnants were re-covered, allowed to cure at ambient temperature until 4 weeks had elapsed since casting, then removed from the dishes and dried for 2 days under vacuum. pCD films were characterized using an Excalibur FTS 3000 FTIR spectrometer
(BioRad, Hercules, CA) equipped with a Pike MIRacle single-reflection attenuated total reflectance (ATR) accessory, germanium crystal, and flat-tipped pressure anvil (Pike
Technologies). Scans were collected (4 cm-1 resolution, 800-4000 cm-1 range, 5 kHz speed, sensitivity 16, open aperture, 100 co-added scans, and Boxcar apodization function) on 3 samples per group while dry nitrogen gas was continuously flowed to purge the system of
CO2 and water vapor. Background scans were collected before the first sample, and absorbance values relative to background were converted to % transmittance. In Microsoft
Excel, spectra were averaged to include all samples within each group, shifted to set the average value between 1900-2200 cm-1 to 100%, and normalized such that the magnitude of the 1037 cm-1 peak attributed to stretching of the CD ether (C-O-C) functionality 676,813 was held constant across all CD-based groups to account for differences in crystal-sample contact. The curve for HDI was normalized such that the peak at 2926 cm-1 corresponding to alkane C-H stretch was equal in magnitude to that in the curve for 0.64 pCD.
4.3.5. Effects of HDI-Crosslinking on pCD Anti-Biofouling Performance
4.3.5.1. Evaluation of Protein Adsorption
Given the important role of protein adsorption in subsequent biofouling processes, we attempted to characterize the resistance of pCD to protein adsorption. Protein adsorption was assessed using subtractive X-ray Photoelectron Spectroscopy (XPS) for the following surfaces: bare TCPS, bare untreated PP, and bare plasma-treated PP (included as
157 controls), and plasma-treated PP surfaces coated with 0.08, 0.16, 0.32, and 0.64 pCD. Two samples were prepared per group. PP sheet stock was cut to dimensions of ~7.5 mm x 7.5 mm, gently sanded to expose fresh surface using a graded series (1200, 2500, then 5000 grit) of SiC sandpaper, and rinsed thoroughly with deionized water. Plasma treatments and pCD coatings were then applied to the appropriate samples. After curing of pCD, all samples were hydrated in deionized water. One sample per group was then incubated in a sterile undiluted solution of heparinated bovine plasma for 24 h at 37°C on a rotisserie shaker, while the other sample per group remained in a bath of deionized water. After incubation, protein-exposed samples were then washed thoroughly with deionized water.
All samples were evaluated for elemental content using a PHI Versaprobe 5000 Scanning
X-Ray Photoelectron Spectrometer equipped with Al Kα source (hν = 1486.6 eV). Survey scans were collected using 200 µm spot size, 45 W power, 15 kV acceleration voltage,
117.40 eV pass energy, 0.40 eV step size, 25 ms/step, 8 cycles, 44.7° take-off angle, and
0-1100 eV range. The C1s peak was auto-shifted to 284.8 eV, and the ratios of carbon, nitrogen, and oxygen were analyzed. Peak areas were taken with background set using a
Shirley function from 280-292 eV for C1s, 396-404 eV for N1s, and 526-538 eV for O1s.
Analysis was performed using MultiPak software version 9.8.0.19 (Physical Electronics,
Inc.). The amount of nitrogen detected was considered indicative of both the amount of adsorbed protein, as well as the degree of HDI-crosslinking of pCD polymers on coated samples. Therefore, protein adsorption onto surfaces was measured indirectly based on the difference between the percentage of N/(N+C+O) after incubation in bovine plasma versus without incubation in bovine plasma. Results shown represent findings from one experiment.
158 4.3.5.2. Investigation of Mammalian Cell Attachment and Viability
Attachment of host cells to an implant surface determines the extent of tissue ingrowth around the biomaterial, and it is critical that the material be safe for these cells so as to promote a favorable host response. Attachment of mammalian cells and viability of attached cells were assessed using hemacytometry and flow cytometry for the following surfaces: bare TCPS, bare PS, bare untreated PP, and bare plasma-treated PP (included as controls), and plasma-treated PP surfaces coated with 0.08, 0.16, 0.32, and 0.64 pCD.
NIH/3T3 fibroblasts were suspended in sterile DMEM with 10% FBS and 1% penicillin- streptomycin, and seeded onto culture surfaces at a supra-confluent density of 150,000 cells/cm2. Cells were allowed to attach for either 3 h or 21 h at 37°C prior to rinsing twice with 750 µL sterile media, then once with 750 µL sterile PBS. Cells were then enzymatically detached from surfaces using 200 µL 0.25% Trypsin-EDTA followed by 2 further rinses of 500 µL sterile media. Rinsates and trypsinates were retained in separate tubes and analyzed using flow cytometry on a Millipore Guava EasyCyte using Guava
ViaCount as a viability indicator according to manufacturer instructions. Results shown represent pooled findings from 5 independent experiments for 3 h hemacytometry, and 2 independent experiments each for 21 h hemacytometry, 3 h viability, and 21 h viability (1 sample per experiment).
4.3.5.3. Measurement of S. aureus Attachment
Attachment of bacteria to a biomaterial surface is a key step in prosthetic infection, therefore we investigated the ability of both gram-positive (S. aureus) and gram-negative
(E. coli) bacteria to adhere to pCD. S. aureus attachment was assessed using colony- forming unit (CFU) counts. This was done for bare untreated PP control, and plasma-
159 treated PP coated with 0.16 pCD. This was the only pCD formulation tested in this experiment given the low-throughput nature of CFU analysis, and was chosen for its demonstrated ABF performance in mammalian cell attachment and protein adsorption experiments. PP rod stock with a ¼” diameter was cut into cylinders of lengths between 2-
3.5 mm (n = 4-5 per group). Cut cylindrical samples were sanded smooth (to ensure comparable surface topography between samples) using a graded series of SiC sandpaper.
The length and diameter of each individual sample were recorded to the nearest 0.01 mm using digital calipers. Samples were submerged in 70% ethanol and allowed to dry, then plasma treatments and coatings were applied to the appropriate samples. After coatings had been allowed to cure, all samples were hydrated in sterile deionized water. Xen30 S. aureus bacteria were thawed from frozen stock, inoculated into a 14 mL round-bottom tube of sterile BBL broth, and expanded in suspension for 24 h in a dedicated 37°C incubator while the tube lid was vented. The tube was then removed from the incubator and stored at 4°C to maintain bacteria in the stationary phase. Prior to seeding onto experimental surfaces, the bacterial suspension was diluted to an OD600 of 0.72 relative to sterile BBL broth. The bacterial concentration at this OD600 is estimated as 2x1012 mL-1 based on CFU counts of dilutions spread onto agar plates. The bacterial suspension was then directly seeded at 1 mL per sample onto non-coated and coated cylinders in 1.5 mL tubes, then incubated at
37°C for 24 h on a rotisserie shaker. Following incubation, cylinders were rinsed by dipping in 5 sequential 2 mL baths of sterile PBS and then suspended in 1 mL sterile BBL broth and ground at 30,000 revolutions per minute with a tissue homogenizer using separate sterile blades for each individual sample. The bacterial concentration in each homogenized suspension was then quantified by spreading dilutions onto surfaces of agar-
160 coated plates and performing CFU counts after 18 h at 37°C. The number of adherent bacteria per square mm of sample surface area was then calculated. Results shown represent findings from one experiment, with the same trend having also been observed in one similar independent experiment.
4.3.5.4. Assessment of E. coli Attachment
E. coli attachment was assessed using bioluminescence measurements for the following 96-well plate surfaces: bare untreated PP, bare plasma-treated PP, bare TCPS, untreated PP coated with Pluronic F108, pHEMA-coated untreated PP, and plasma-treated
PP coated with 0.08, 0.16, 0.32, and 0.64 pCD. The additional controls Pluronic F108 and pHEMA were included solely in this experiment because bacterial attachment is involved in the majority of biofouling challenges encountered in both medicine and industry, and because bioluminescence measurement is high-throughput, thus facilitating their study. All coatings (pCD pre-polymer mixtures, sterile-filtered 2% w/v pHEMA in 95% ethanol, and sterile-filtered 2% w/v Pluronic F108 in PBS) were applied at 42 µL/well. pHEMA coatings were left uncovered in a sterile hood and allowed to dry overnight before being covered with Parafilm and stored. Pluronic coatings were covered with Parafilm for 2 days, then rinsed several times and stored immersed in sterile PBS. Bioluminescent ilux pGEX
(-) E. coli were thawed from frozen stock, inoculated into a 14 mL round-bottom tube of sterile lysogeny broth (LB), and expanded in suspension for 24 h in a dedicated 37°C incubator with the tube lid vented. The tube was then removed from the incubator and stored at 4°C to maintain bacteria in the stationary phase. Prior to seeding onto experimental surfaces, the bacterial suspension was diluted to an optical density at 600 nm
(OD600) of 0.50 relative to sterile LB broth. The bacterial concentration at this OD600 is
161 estimated as 1011 mL-1 based on CFU counts of dilutions spread onto agar plates. The bacterial suspension was then directly seeded at 100 µL/well, cultured under static conditions for 24 h at 37°C, and following incubation, wells were rinsed 5 times with 200
µL/well sterile LB broth, then emptied and filled a final time with 100 µL/well sterile LB broth. Bioluminescence measurement was then performed using a Biotek Synergy H1 plate reader (~23°C read temperature, luminescence endpoint scan, 5 s integration time, 1 mm read height, full light emission, 135 gain, top optics, 100 ms delay, extended dynamic range). A subset of non-rinsed wells for each type of surface were also seeded in duplicate with several known concentrations of bacteria, included for creation of standard curves.
Results shown represent pooled findings from 2 independent experiments (n = 16-32 total wells per group).
4.3.6. Statistical Analysis
All data is presented as mean ± standard deviation. Statistical significance was defined for all analyses as p<0.05. Bivariate correlations were evaluated using Spearman’s rho (rS) in Minitab 2019. All other data comparisons were performed using 2-tailed 2- sample t-tests with unequal variance in Microsoft Excel 2016.
162 4.4. Results
4.4.1. Effects of HDI-Crosslinking on Physicochemical Properties of pCD
Given that the overarching goal of this work was to explore the usefulness of pCD for ABF applications, we first sought to broadly characterize the material properties of several different CD polymer formulations that might potentially be applied towards ABF purposes. The effects of crosslinking on swelling ratio, elastic modulus, contact angle, and chemical composition of pCD were thus explored. Additionally, considering that these polymers might even be used for simultaneous passive ABF and active ABF or drug delivery, we also assessed effects of drug-loading on elastic modulus and contact angle of pCD. To this end, the antibiotic rifampicin (RIF), was chosen as a model drug, due to its poor water-solubility, and known ability to form inclusion complexes with CD subunits.
Swelling ratio describes the ability of network polymers to absorb water, and the degree of swelling may reflect polymer chain mobility, therefore influencing ABF performance 810. The swelling ratio of CD polymers was found to decrease with increasing
HDI crosslinking ratio (rS = -0.968, p < 0.001) (Figure 4-2A). This could be a result of the decreasing mobility of the polymer network upon further crosslinking.
Biomaterial stiffness (i.e. elastic modulus) is considered to be a factor which impacts adhesion and spreading of cells. In agreement with findings of decreasing swelling ratio, elastic moduli of CD polymers were found to increase with increasing HDI crosslinking ratio (Figure 4-2B). This was true for empty pCD after 4 days (rS = 0.802, p
< 0.001) or 4 weeks (rS = 0.969, p < 0.001) of crosslinking, and for RIF-filled pCD after 4 days of crosslinking (rS = 0.796, p < 0.001). Increasing the duration of crosslinking from 4 days to 4 weeks led to a doubling in elastic modulus at the 0.64 crosslink ratio (p < 0.001),
163 reflecting increased formation of covalent crosslinks within the pCD network over this time frame. Drug-loading of RIF had no meaningful effect on elastic modulus at any crosslink ratio (p > 0.07). All pCD samples except for 43% of those in the 4-day 0.08 pCD empty group, and 13% of those in the 4-day 0.08 pCD-RIF group, were able to sustain loads up to the 10N load limit without apparent yielding or failure. This indicates possible fragility of the polymers at the lowest crosslinking ratio. Incorporation of RIF into pCD was previously shown to increase compressive properties of pCD microparticle-laden bone cement composites on the order of MPa, an effect thought to be in part due to RIF complexation hindering structural collapse of molecular pockets within pCD 814. Results here, however, suggest that RIF has little direct effect on the compressive moduli of isolated pCD materials.
Surface wettability is recognized as an important property that affects biofouling resistance of a material. Static water contact angles of CD polymers increased with increasing HDI crosslink ratio (Figure 4-2C), likely as a consequence of incorporation of nonpolar hexamethylene spacers with further crosslinking. This was apparent for both empty (rS = 0.825, p < 0.001) and RIF-filled pCD (rS = 0.473, p < 0.017). This indicates increasing hydrophobicity of pCD with further incorporation of nonpolar HDI crosslinks.
Interestingly, incorporation of the poorly water-soluble RIF into pCD led to decreased static water contact angles at most crosslink ratios (p < 0.015) except for 0.16 (p = 0.753).
This result may reflect diffusion of RIF into the water droplet, or increased presentation of polar functionalities (which conceal nonpolar HDI crosslinks) at the surface as RIF molecules complex with CD subunits.
164 A
B
C
Figure 4-2. Physical properties of 0.08, 0.16, 0.32, and 0.64 pCD formulations: A) swelling ratio, B) elastic modulus, C) static water contact angle. Points in plots for elastic modulus and contact angle are intentionally staggered along the abscissa to improve readability.
165
Normalized FTIR spectra (Figure 4-3) revealed that increased crosslinking ratio led to relatively larger peak sizes at 1256cm-1, 2856 cm-1, and 2926 cm-1 corresponding to alkane C-H stretch 676,813,815–817, at 3332 cm-1 corresponding to both hydrogen-bonded alcohol O-H and secondary amide N-H stretch 815,818, at 1576 cm-1 corresponding to urethane N-H stretch 676,819, at 1624 cm-1 corresponding to urethane C=O bending 676, at
1278 cm-1 corresponding to urethane C-N stretch 815, and at 1460 cm-1 corresponding to C-
H bending 817. Together, these peak size increases confirm increased abundance of HDI crosslinks with increased crosslink ratio. Additionally, for pCD spectra, disappearance of the peak seen in the HDI spectrum at 2270 cm-1 indicates that isocyanates were fully converted (most likely to urethanes) in crosslinking reactions 815,816,819,820.
166
Figure 4-3. Normalized ATR-FTIR spectra of pCD materials and components. Spectra are intentionally shifted along the ordinate to improve readability. Going from left to right, arrows are drawn at 2926 cm-1 for the alkane C-H stretch peak, 2270 cm-1 for the isocyanate peak only seen in the HDI spectrum, and 1576 cm-1 for the urethane N-H stretch peak which appears only after pCD crosslinking.
4.4.2. Effects of HDI-Crosslinking on Protein Adsorption to pCD
Having characterized the effects of HDI-crosslinking on the physical properties of pCD, we next sought to evaluate non-specific protein adsorption on the surfaces of these polymers. Subtractive XPS was used to measure protein adsorption onto polymer surfaces.
167 This method was chosen over colorimetric or fluorimetric protein binding assays because several of our preliminary studies indicated that the tags and labels (e.g. FITC, Coomassie brilliant blue) used to visualize proteins in such experiments had inherent affinity for CD, thus making such strategies insufficiently specific and sensitive for detection of protein adsorption (data not shown). Prior literature corroborates these observations 821,822.
XPS measurements suggested that protein adsorption onto pCD polymers was lower than that onto TCPS or PP (Figure 4-4). Protein adsorption onto pCD was observed to decrease with decreasing crosslink ratio, possibly as a result of increasing hydrophilic character. The coating for the 0.08 pCD surface was found to have delaminated during a rinse step, so no data is shown for that sample. Plasma treatments were observed to reduce protein adsorption onto PP surfaces, in agreement with previous findings 642, again possibly a result of increasing wettability.
Figure 4-4. Protein adsorption onto surfaces as measured using subtractive XPS. Increased nitrogen content reflects increased protein adsorption.
168 4.4.3. Effects of HDI-Crosslinking on Fibroblast Adhesion to pCD
Following protein adsorption experiments, mammalian cell (NIH/3T3 fibroblast) adhesion and viability were studied using hemacytometry and flow cytometry, respectively. These methods were chosen over optical or colorimetric in situ measurements because our preliminary studies suggested that indicator dyes for cell metabolism (e.g.
Alamar blue) had inherent affinity for CD, thus interfering with assay results (data not shown). Prior literature corroborates this observation 823. Furthermore, while certain CD- based polymers can be made optically transparent 824, the translucent PP substrates and CD polyurethane coatings in this study were not optimized for cell observation.
Cell adhesion for pCD polymers was found to range from 0-20% of the level of cell adhesion observed for TCPS (p < 0.001) (Figure 4-5A). Among pCD polymers, the most crosslinked, 0.64 formulation was found to possess the least resistance to cell adhesion (p
< 0.003). This indicates that lightly-crosslinked pCD polymers resisted stable adhesion of fibroblasts, likely due to low levels of protein adsorption. Cell adhesion was equally low for untreated PP as for most pCD polymers (p > 0.293). Plasma treatment of PP increased cell adhesion (p = 0.01), consistent with prior findings 642. Likewise, cell adhesion for
TCPS (commercially plasma-treated PS) was higher than that for untreated PS (p < 0.003).
The low cell adhesion to untreated PP may indicate that the identities and/or conformations of the serum proteins which adsorb onto the PP surface are unfavorable for cell adhesion.
Additionally, despite relatively lower apparent levels of protein adsorption on plasma- treated PP, the adsorbed proteins on this material seem to be more favorable for cell adhesion than those on untreated PP. The high standard deviations for plasma PP and untreated PS appeared to a result of the fibroblasts forming monolayers and either
169 detaching all at once, or remaining attached altogether, resulting in an apparent bimodal distribution for cell adhesion in these groups.
Viability of the cells that did attach to pCD polymers was comparable to that of controls (p > 0.08) (Figure 4-5B), with all values >75%, suggesting that pCD polymers did not exert cytotoxic effects on cells. This result is encouraging, as previous studies have indicated that at high concentrations, non-polymeric CD can reduce cell viability by abstracting cholesterol from cell membranes 825. Polymerization of CD monomers may help mitigate such cytotoxic effects. Similar trends for cell counts and viability were observed at both 3h and 21h.
170 A
B
Figure 4-5. Fibroblast A) cell counts, and B) viability, following 3h or 21h of adhesion to control and pCD surfaces. Data reflect only the cells that attached to pCD, so as to exclude effects of anoikis-induced cell death. *Denotes significant difference to all other groups. ⸕Denotes significant difference to all pCD groups and untreated PP. ⸙Denotes significant difference to untreated PP and .08, .16, and .32 pCD. No differences were observed between groups in terms of viability.
171 4.4.4. Effects of HDI-Crosslinking on Bacterial Attachment to pCD
Having evaluated protein adsorption and mammalian cell adhesion onto pCD, bacterial attachment was next investigated. This was done in two separate sets of experiments, the first for the gram-positive species S. aureus, in this case the methicillin- resistant Xen30 strain, and the second for the gram-negative species E. coli, specifically the bioluminescent ilux pGEX(-) strain.
S. aureus attachment to .16 pCD after 24h was significantly lower than attachment to untreated PP (p = 0.041) (Figure 4-6A). Likewise, E. coli attachment to plasma-treated
PP, .08 pCD, .16 pCD, .32 pCD, and PP coated with Pluronic F108 after 24h was significantly lower than attachment to untreated PP (p < 0.001) (Figure 4-6B). E. coli attachment to surfaces coated with Pluronic F108, chosen a control because of its ease of application and highly effective passive ABF properties 826, was roughly half that seen for
.08 and .16 pCD (p < 0.014). Surprisingly, pHEMA did not resist attachment of E. coli as well as TCPS or PP (p < 0.001), despite its common use for creating coatings on cultureware that prevent adhesion of mammalian cells. This could be a reflection of differences in the mechanisms by which E. coli and anchorage-dependent mammalian cells adhere to surfaces. In support of this explanation, plasma treatment suppressed attachment of E. coli despite enhancing attachment of NIH/3T3 fibroblasts.
172 A
B
Figure 4-6. Bacterial attachment of A) S. Aureus and B) E. coli. *Represents significant difference (p<0.05) to PP.
173 4.5. Discussion
This study has demonstrated that CD polymers, in addition to their ability to actively release biocidal compounds, have potential to passively resist biofouling by proteins, mammalian cells, and bacteria, without apparent cytotoxicity. Increased HDI crosslinking of CD polymers was found to decrease swellability, increase rigidity, and increase hydrophobicity, but at the same time tended to reduce pCD resistance to protein, cell, and bacterial attachment. This reduction in ABF properties with increased HDI crosslinking might be expected given that less wettable materials have a stronger tendency to promote protein adsorption 222–224,756,776–778 (along with subsequent biofouling events), that higher substrate rigidity promotes mammalian cell adhesion and spreading 811,812, and that more abundant crosslinks should restrict chain mobility and limit presentation of polar domains at the network polymer surface. Although this study design did not allow for independent determination of the degree to which each material property individually affected biofouling resistance of pCD, results here provide insights into the rational design of pCD for ABF purposes and set the stage for further exploration.
Carbohydrate surface coatings, such as dextran, cellulose, or mannose, have previously been investigated for ABF applications 790–792, but unlike CD, these sugars lack the ability to form affinity-based inclusions of small molecule drugs. Affinity properties give pCD materials unique advantages for active ABF, such as more sustained drug release
635,636,640 that is relatively independent of material dimensions 827, and more efficient biocide refilling 637. PEG-based coatings may be the most popular materials studied for
ABF purposes. Results from E. coli repulsion experiments suggest that lightly-crosslinked pCD formulations may perform nearly as well in terms of resistance to bacterial adhesion
174 as a freshly PEGylated surface. However, PEG can be prone to oxidation in the presence of oxygen and metal ions found in physiological solutions 828, and many human patients have shown antibodies to PEG following its increasing use in food, cosmetic, and pharmaceutical products 829, potentially limiting its utility for long-term or medical ABF applications. Cyclodextrins, conversely, do not elicit immune responses in mammals 797, and the CD subunits and urethane bonds in pCD are in theory quite stable. Zwitterionic polymers are another class of polymers that have become popular for ABF research given outstanding performance, their main drawback being that they are difficult and costly to synthesize 787. pCD on the other hand is easy to make, and inexpensive.
Apart from ABF, a unique advantage of pCD is its ability to complex with and deliver small-molecule drug cargo over long periods of time. It may seem counterintuitive that a polymer having inherent affinity for small-molecule drugs could be useful for prevention of protein adsorption, cell adhesion, and bacterial attachment. One possible explanation is the much smaller size of drug compounds – which must be able to fit in a molecular pocket having a diameter of 5.7Å, 7.8Å, or 9.5Å for the native cyclodextrins α-
CD, β-CD, and γ-CD, respectively 830 – than proteins, bacteria, and cells. Most proteins, which are typically at least one order of magnitude larger than a small-molecule drug (yet still much smaller than bacteria or cells), are likely too bulky to diffuse through the pCD polymer matrix let alone complex with CD subunits. Additionally, unlike small-molecule drugs, proteins in aqueous solution exist in conformations that inherently shield their nonpolar domains against hydrophobic interactions with CD subunits. Similarly, polyurethane materials such as pCD undergo surface restructuring in aqueous solution through microphase separation, which enriches the polymer surface with hydrophilic
175 domains 831, further deterring protein displacement of water and adsorption. With increased
HDI crosslinking, the proportion of hydrophobic segments in pCD increases, while their mobility (i.e. ability to be shielded from protein solutes at the interface) decreases, with the consequence being increased protein adsorption and subsequent biofouling.
Interestingly, incorporation of the poorly water-soluble antibiotic RIF into pCD increased wettability and had minimal effect on material rigidity. Considering that increased wettability might generally be expected to improve passive ABF properties, these results together may suggest that drug loading with biocidal agents, like RIF, for purposes of active ABF does not negatively influence the passive ABF properties of pCD.
Nonthermal plasma treatments of PP in this study, just as in a prior study 642, were shown to reduce non-specific protein adsorption and E. coli attachment, while enhancing adhesion of mammalian fibroblasts, all of which could be favorable outcomes for a plasma- treated biomedical implant. However, the effects of plasma treatment are known to diminish (i.e. age) over time. Literature suggests that hydrophobic recovery of plasma- treated PP normally occurs over the course of several weeks, the rate being dependent on environmental factors and polymer crystallinity 701–704. The presence of a coating may act to protect or constrain the activated PP substrate surface.
One limitation of this study is that the effects of different crosslinkers or subunits were not examined. A more hydrophilic crosslinker might in theory produce polymers that remain resistant to biofouling regardless of crosslink ratio. In this study, HDI was chosen as a crosslinker to allow investigation of neutrally-charged polymers having different overall hydrophobicity. In preliminary studies, a different neutrally-charged crosslinker, ethylene glycol diglycidyl ether, was found to produce polymers that swelled to the point
176 of rupture and delamination upon hydration post-curing. The polymers produced by HDI- crosslinking were more robust, therefore we considered HDI a more useful crosslinker for preparation of pCD coating materials here. There are many other possible crosslinkers that could be considered 815,832–835. Similarly, different CD subunits were not evaluated in this study. Thus, we were not able to comprehensively assess the generality of the passive ABF performance of pCD. β-CD was considered the subunit of choice among the native cyclodextrins given that from a drug delivery standpoint, it is the most widely used and generally has the most versatility in terms of the variety of guest molecules that it can complex with 836–839. However, each of the native cyclodextrins possess different dimensions and water solubilities, thus changing the subunit could foreseeably alter polymer affinity properties and hydrophilicity, as well as ABF performance. For instance,
Dos Santos et al. showed that incorporation of β-CD and γ-CD decreased lysozyme and albumin adsorption onto pHEMA hydrogels, while incorporation of α-CD decreased lysozyme adsorption but increased albumin adsorption 840.
We envision that pCD might be applied as coatings for medical devices, such as textile implants. Specifically, CD-based polymer coatings have previously been investigated for use as surgical mesh coatings that could enable controlled antibiotic/drug release 639–641,659,676,841, but it would also be noteworthy to explore the potential of these materials to prevent post-surgical adhesions. The formation of tissue adhesions on a surgical fabric surface is preceded by two events that CD-based polymers are shown to be capable of mitigating: 1.) non-specific adsorption and denaturation of protein (especially fibrinogen), which promotes coagulation, inflammation, and impaired fibrinolysis, and 2.) adhesion of fibroblasts that remodel persistent coagulation products into fibrous
177 connections between the implant surface and adjacent tissue structures. Alternatively, pCD coatings may be useful for application onto PP water treatment membranes. Such coatings might simultaneously resist bacterial biofouling and remove small-molecule pollutants from water 842. This investigation provides a basis for the use of CD-based polymers for such applications. Future studies should seek to directly evaluate field or in vivo performance of pCD coatings for ABF purposes.
4.6. Conclusions
This study has demonstrated across several different readouts the potential of cyclodextrin-based polymers to passively resist biofouling by proteins, mammalian cells, and bacteria, without overt cytotoxicity. Additionally, the effects of HDI crosslinking on pCD swellability, rigidity, and wettability were characterized. Lower HDI-crosslinking densities yielded pCD with superior resistance to biofouling, an effect that could be attributed to increased chain mobility, decreased rigidity, and/or decreased hydrophobicity.
Cyclodextrin-based coating materials may be appropriate in industrial or medical applications for which biofouling-resistant and/or drug-delivering surfaces are required.
178 4.7. Acknowledgments
The authors gratefully acknowledge support from National Institutes of Health:
NIH R01GM121477 (HvR), and NIH NIAMS Ruth L. Kirschstein NRSA T32 AR007505
Training Program in Musculoskeletal Research (GDL). Core facility services provided by the Swagelok Center for Surface Analysis of Materials, and the Cytometry & Microscopy core at Case Western Reserve University are also appreciated. The authors also thank
Joseph Mansour for use of the Rheometrics instrument, the Advincula group for use of the contact angle goniometer, Kevin Abbassi for expertise and assistance with XPS, David
Jasen WuWong for assistance with S. aureus bacterial attachment experiments, Katherine
Yan for assistance with sample preparation, and Erika Cyphert and Ali Ansari for writing suggestions.
179 5. CHAPTER 5: PNEUMOPERITONEAL COMPUTED TOMOGRAPHY AS A
PRE-CLINICAL APPROACH FOR VISUALIZING INTRA-ABDOMINAL
ADHESIONS TO PROSTHETIC MESH
Authors: Greg D. Learn, Emerson J. Lai, Erika L. Cyphert, Aldo Fafaj, Michael J.
Rosen, Horst A. von Recum
5.1. Abstract
Peritoneal adhesion formation is a challenging, undesirable consequence associated with intraperitoneal placement of surgical meshes, as performed in hernioplasty. One obstacle impeding progress in the study of mesh adhesions (and their prevention) is that neither adhesions nor meshes are readily visible using conventional non-invasive imaging approaches. Additionally, standard methods for evaluating adhesions rely heavily on subjective scoring. There is an unmet need for approaches that can permit non-invasive and/or objective characterization of mesh adhesions. In this study, a computed tomography
(CT)-based imaging strategy was explored for observing adhesions to implanted mesh in rats. Bare polypropylene (PP) mesh, known to trigger adhesions, and PP mesh covered with polymerized cyclodextrin (pCD), an experimental adhesion-resistant barrier material, were implanted into the abdominal cavity of one animal each. After one month, animals were sacrificed, and adhesions were studied using pneumoperitoneal CT, as well as direct visualization to confirm findings. The pCD-covered PP mesh demonstrated fewer and less severe adhesions than bare PP mesh, and adhesions for the former were restricted to implant borders. This was apparent with both the imaging and direct visualization methods.
180 This preliminary study suggests that pneumoperitoneal CT can be used to detect differences in distribution and extensiveness of mesh adhesions.
5.2. Introduction
Peritoneal adhesions are the most common complication following abdominal and pelvic surgery 15–17, reportedly occurring in up to 93% 18 and 55-100% 19 of patients, respectively. These bands of fibrous internal scar tissue are detrimental to both patient health and healthcare resource utilization. Peritoneal adhesions can cause complications ranging from bowel obstruction 17,62–64 to infertility 66–68, and can increase the complexity and hazards of future surgeries 17,71,73,79,80.
Polymeric mesh implants are frequently utilized in conjunction with abdominal and pelvic surgery, often for the repair of hernias. Over 700 thousand inguinal hernia repairs
22–25, 350 thousand ventral hernia repairs 21, and 4 thousand parastomal hernia repairs 26 are performed each year in the US. Most of these procedures involve surgical mesh 23,27,36,39,40, as mesh prostheses decrease hernia recurrence rates when compared with suture-only reconstruction 27,31–36.
Unfortunately, meshes are often associated with peritoneal adhesions and related complications. This is especially the case with intraperitoneal placement 44, which becomes necessary in minimally invasive hernia repair 45. Formation of adhesions is particularly prevalent with mesh devices made from polypropylene (PP), the most commonly used material in hernia repair meshes 46–50. PP meshes are often selected for desirable qualities such as their rapid incorporation with the abdominal wall, adequate strength, and low cost
51,52. However, the very characteristics of PP mesh that encourage incorporation with the abdominal wall also tend to promote peritoneal adhesion formation 51,52.
181 There is immense interest in approaches that can mitigate adhesions, both generally, and in association with mesh devices. Commercial efforts have sought to address the problem of mesh adhesions by introducing mesh products with adhesion-resistant barriers affixed to the visceral face. However, information regarding the comparative effectiveness of these devices is quite limited, as mesh-related adhesions are inherently difficult to study.
The standard methods for evaluating adhesions in animal and human studies involve scoring, during direct visualization at necropsy or during second-look surgery, respectively. These approaches are both invasive and subjective. The invasiveness of these approaches limits the ability to monitor adhesions over time (i.e. longitudinal studies), and their inherently subjective nature (given the process of scoring) may reduce comparability between studies. One potential reason for the lack of less invasive or more objective methods to study mesh adhesions is that neither adhesions 58 nor most mesh devices 57 are easily seen using conventional imaging techniques. This hinders the study of mesh adhesions in both animals and humans.
Pneumoperitoneal MRI is a technique that was recently introduced for visualizing mesh adhesions in a pre-clinical rat model, though the method has been described in just one study 628. This approach was demonstrated in live animals, and entails insufflation of the abdomen during MRI scanning to provide high-contrast separation between tissues, thus facilitating visualization of peritoneal adhesions. In the pre-clinical setting, pneumoperitoneal MRI is elegantly simple, but potential drawbacks include that MRI systems may be inaccessible to most labs, that certain mesh materials, such as PP 57, are indistinguishable from adjacent soft tissue with MRI, and that MRI cannot achieve the same spatial resolution as certain other methods, such as micro-CT. Additionally, with
182 regards to a potential clinical scenario, MRI may be contra-indicated for certain patient populations, such as those with metallic implants. However, the principle of insufflation of the abdomen during imaging might be extended to other non-invasive imaging modalities.
This could permit visualization of mesh adhesions with a wider variety of implant materials and spatial resolutions.
In this work, a CT-based approach involving insufflation of the abdomen, pneumoperitoneal CT, is introduced and preliminarily explored in efforts to address the need for alternative non-invasive strategies that permit evaluation of adhesions to surgical meshes. The hypothesis of this study is that pneumoperitoneal CT enables visualization and detection of large differences in adhesion characteristics between two diverse mesh types: bare PP mesh as a high-adhesion control, and PP mesh covered with an experimental adhesion barrier material, polymerized cyclodextrin (pCD), as a low-adhesion control. The pCD material is a polysaccharide-based network polyurethane that has previously demonstrated ability to passively repel proteins, mammalian fibroblasts, and multiple strains of bacteria, without any signs of overt cytotoxicity 638,824. Investigating resistance of this polymer to tissue adhesions was therefore a secondary objective of this work. High- radiodensity barium-containing tissue markers were incorporated into PP mesh to facilitate identification of mesh borders in the presence of soft tissue. One month after mesh implantation in rats, animals were sacrificed and peritoneal adhesions were visualized using both pneumoperitoneal CT and direct visualization for comparison.
183 5.3. Materials and Methods
5.3.1. Materials
Prolene mesh (Ethicon) and 4-0 Prolene Suture (Ethicon) were purchased from eSutures. FibermarX Radiopaque Tissue Marker (Viscus Biologics LLC, Cleveland, OH) was a gift from Dr. Kathleen Derwin. Soluble, lightly epichlorohydrin-crosslinked β-CD polymer precursor (bCD) (#CY-2009) was purchased from CycloLab R&D (Budapest,
Hungary). Hexamethylene diisocyanate (HDI) crosslinker (#52649) was purchased from
Sigma Aldrich. N,N-dimethylformamide (DMF) solvent (#D119-4) was purchased from
Fisher Scientific.
5.3.2. pCD Preparation
bCD was desiccated for 1h at 85°C under a vacuum pressure of 50 kPa (~380 torr), then weighed into a 50 mL polypropylene tube. DMF was added to the bCD at a ratio of
33% w/v, and the mixture was vortexed. After the bCD had completely dissolved, HDI was added at a ratio of 160 µL per 1 g bCD, and the mixture was again vortexed. This pCD pre- polymer mixture was applied to the visceral side of PP mesh within 20 min of mixing, then allowed to cure to form a pCD adhesion barrier.
5.3.3. Mesh Preparation and Sterilization
Prolene meshes were cut to 1 inch x 1 inch squares. Two 1 inch pieces of FibermarX
Radiopaque Tissue Marker, which is a barium sulfate infused polypropylene suture indicated for radiographically marking soft tissue, were cut for each mesh square. These tissue markers were placed parallel to the mesh longitudinal direction, one adjacent to each
184 border, loosely woven into the mesh structure, and 3 interrupted surgeon’s knots (each made using 3 square throws) of 4-0 Prolene were used to secure each in place (Figure 5-1).
All knots were placed on the same side of the mesh. Next, each mesh receiving a pCD adhesion barrier was activated with Ar/H2O nonthermal plasma (500 mTorr, 50 W, 13.56
MHz) for 10 min using a Branson/IPC Model #1005-248 Gas Plasma Cleaner, then laid flat on a piece of Parafilm, and within 2 h of plasma treatment 700 µL pCD pre-polymer mixture was pipetted through the mesh interstices. Being denser than the PP mesh, the pCD pre-polymer mixture formed a liquid layer between the Parafilm and the mesh (Figure
5-2A), which was allowed to polymerize for 5 days at room temperature (Figure 5-2B).
Meshes were then inserted into individual 50 mL PP tubes. For pCD-covered mesh, 25 mL
PBS was added to the tube, so as to immerse the mesh and minimize warping or drying of the pCD layer during subsequent autoclave sterilization treatment. Caps were loosely screwed onto tubes and meshes were autoclaved in the tubes at 121°C for 20 min with slow exhaust. Similar pCD materials have withstood autoclave sterilization at 121°C for up to
40 min, previously shown to increase their mechanical robustness in terms of Young’s modulus 843. Additionally, PP surgical meshes in prior studies demonstrated no significant change in mechanical properties following a single autoclave re-sterilization at 134°C for
~60 min 844.
185
Figure 5-1. Prolene mesh with radiopaque tissue markers. Scale bar = 1 cm.
Figure 5-2. Mesh with pCD barrier applied A) before polymerization, and B) after polymerization.
5.3.4. Animals
All animal procedures were conducted in accordance with standard policies at the
CWRU Animal Resource Center and an IACUC-approved animal protocol (#2018-0071).
Adult male Sprague Dawley rats (n = 1 bare PP, n = 3 pCD-covered PP) were purchased from Charles River and acclimated to the research facility for a minimum of 5 days prior to any animal manipulation. Animals used in surgery weighed 250-300 grams.
186 5.3.5. Mesh Implantation and Animal Recovery
Animals were anesthetized using isoflurane, shaved over their abdomens, then their skin was wiped with antiseptic solutions. A local subcutaneous injection of lidocaine (10 mg/kg at 20 mg/mL) was given, then the rats were sterile draped. A longitudinal incision
(~1.5 cm long) of the skin and abdominal wall muscle layers was made through the linea alba. Meshes were folded and passed through the incision to implant them within the peritoneum. Meshes were placed so as to cover the incision and affixed to the abdominal wall with Prolene suture, with one interrupted knot per mesh corner. Next, the incision in the muscle layer was closed using a running Prolene suture, and the skin incision was closed using wound clips. For post-operative analgesia, animals were given subcutaneous injections of meloxicam (1.25 mg/mL) immediately post-surgery (2 mg/kg) then 24 h and
48 h later (1 mg/kg), and buprenorphine (0.03 mg/kg at 0.03 mg/mL) immediately post- surgery then 5-6 h later. Rats were monitored at least once daily until post-operative day
14 at which point wound clips were removed under isoflurane anesthesia.
5.3.6. Animal Sacrifice and Pneumoperitoneal CT
One month following surgery, rats were euthanized using CO2 inhalation followed by bilateral thoracotomy. Pneumoperitoneal CT was performed within a few hours of sacrifice. Prior to insufflation, a drop of cyanoacrylate glue was applied to seal the puncture produced in the chest wall and allowed to cure briefly, so as to minimize possibility of subsequent air leakage. To establish pneumoperitoneum, air (60 mL per animal) was delivered via intraperitoneal injection into rat abdominal cavities (Figure 5-3). Trial attempts using rat carcasses had suggested that no deflation of the peritoneal cavity could be visibly detected over a 30 min period following this insufflation procedure (data not
187 shown). Within 5 min after establishing pneumoperitoneum, rats were positioned in a
Siemens Inveon positron emission tomography (PET) / CT scanner. CT scanning was performed using Siemens Inveon Acquisition Workspace software. Scan settings included:
Rotation = Continuous, Total Rotation = 360°, Rotation Steps = 512, Voltage = 80 kV,
Current = 500 µA, Settle Time = 2000 ms, Exposure Time = 160 ms, Effective Pixel Size
= 97.1 µm, Binning Factor = 4, Magnification = Low. After scanning was complete, rats were removed from the CT scanner to commence dissection.
Figure 5-3. Depiction of rat abdomen insufflation as preparation for pneumoperitoneal CT. A) Injection of air via syringe into the peritoneal space. B) Animal abdomen becomes distended. Figure created with BioRender.com 5.
188 5.3.7. Animal Dissection and Adhesion Grading
The lead author performed all animal dissections. For each animal, the skin over the ventral abdominal wall was carefully removed. A video recording was then begun and continued for the remainder of the dissection. A large U-shaped incision was made through the abdominal wall muscle, taking care to avoid the mesh implant and any adhesions. The abdominal wall muscle layer was then lifted up and held in position. Mesh adhesions were observed from several angles, then efforts were begun to divide adhesions, and as this dissection proceeded, the operator verbally described any difficulties encountered (e.g. inability to avoid bowel perforation). Video recordings were reviewed at a later time by the lead author and mesh adhesions were graded on a semi-quantitative scale according to a consensus scoring system 603 (Table 5-1), in terms of their coverage on the mesh visceral side, tenacity, morphology, and organ involvement.
189 Table 5-1. Explanations of the META score components. Adapted with permission (see APPENDIX) from Springer Nature Customer Service Center GmbH: World Journal of Surgery 603. Score Description of Observed Adhesions Assigned Percentage of mesh surface area covered with adhesion tissues 0% 0 1–25% 1 26–50% 2 51–75% 3 76–100% 4 Tenacity of adhesions No adhesions 0 Loose adhesions easily released by traction only 1 Adhesions require sharp dissection, no organ/serosal damage 2 Adhesions require sharp dissection, with unavoidable organ/serosal damage 3 Morphology of adhesions No adhesions 0 Single thin, filmy adhesion 1 Multiple thin, filmy adhesions 2 Single dense adhesion with or without filmy adhesions 3 Multiple dense adhesions with or without filmy adhesions 4 Organs involved with adhesions No adhesions 0 Adhesions between mesh and omentum or a solid organ 1 Adhesions between mesh and part(s) of the intestinal tract 2 Adhesions between mesh and part(s) of the intestinal tract with enteric 3 fistulas or bowel erosion
5.3.8. Three-Dimensional Image Segmentation and Analysis
To derive quantitative metrics regarding mesh adhesion coverage from pneumoperitoneal CT, image segmentation was performed on DICOM file datasets in 3D
Slicer software (version 4.10.2). First, the analysis volume was restricted to the region of interest using the Crop Volume module, so as to reduce file size and computational
190 processing requirements. Second, radiopaque tissue markers were visually aligned with the rendering planes using the Transforms tool. This was done such that the transformed sagittal and transformed axial planes were approximately parallel with the lines connecting the mesh corners (as identified based on the tissue markers), and the transformed coronal plane was roughly parallel with the plane intersecting the mesh corners. Third, a square
(the “mesh footprint square”) was drawn on one slice in the transformed coronal plane bounded by the lines connecting the mesh corners (i.e. coincident with the mesh footprint on the abdominal wall). Fourth, an approximate 3D mesh surface was manually segmented in the transformed axial and transformed sagittal planes using the Segment Editor module and with the help of dissection photos and the mesh tissue markers as boundaries. Fifth, among slices in the transformed coronal plane (i.e. those parallel to the mesh footprint square), the slice containing the point on the approximate 3D mesh surface closest to the viscera was identified (the “interface plane”). Sixth, starting on the slice immediately adjacent and parallel to the interface plane, adhesion tissues were segmented using the
Threshold Paint feature on every slice over a ~1 mm thickness (as determined from the difference between the initial and final slice positions), moving in the direction toward the viscera. Seventh, any threshold volumes that fell outside of a projection (perpendicular to the transformed coronal plane) of the bounds of the mesh footprint square were excluded using the Scissors feature. Eighth, the mesh footprint square and the thresholded adhesion volumes were rendered, the former was oriented parallel to the computer screen, and screenshots were taken (Figure 5-4). Ninth, ImageJ software (version 1.53e) was used to measure the relative total areas of the mesh footprint square and the thresholded adhesion
191 volume in the screenshot images, and these were then used to calculate the percentage of mesh visceral surface covered by adhesions.
Figure 5-4. Screenshots, subsequently used for image analysis in ImageJ, showing renderings of mesh footprint square (yellow) and thresholded adhesion volume (red) for PP mesh (left) and pCD-covered PP mesh (right).
5.4. Results
All animals tolerated surgeries and survived for the entire month. During dissection, it was noted that in two rats from the pCD-covered mesh group, the implants were completely covered in a distinct fibrous capsule. Dissecting into this capsule revealed yellow-green pus (resemblant of guacamole), highly suggestive of prosthetic infection.
These infections are considered to possibly reflect inadequate sterilization (as the meshes were immersed in saline), or contamination introduced during surgery. Note that animals were not given antibiotics (prophylactically or otherwise). More rigorous sterilization and prophylaxis measures will need to be implemented in future studies to prevent such infections. Given the possibility that these infections could impact inflammation and subsequent adhesion formation, findings from these animals are not considered here.
192 Peritoneal adhesions from the rat implanted with bare PP mesh were found to be dense, completely covering the mesh surface, and a loop of intestine was adherent to the
PP mesh. These adhesion features were qualitatively apparent using both pneumoperitoneal CT (Figure 5-5, Figure 5-6) and direct visualization (Figure 5-7).
Additionally, the bowel was unable to be cleanly (i.e. without inadvertent perforation) separated from the mesh, even with meticulous dissection. The severe characteristics of adhesions to the PP mesh, as observed upon direct visualization, were reflected by maximal scores for coverage, tenacity, and morphology (Table 5-2).
Figure 5-5. Adhesions to bare PP mesh. Approximate mesh position is marked with yellow dashed line. Blue arrows delineate radiopaque tissue markers. Red arrows delineate loops of intestine adhered to mesh. A) Axial slice. B) Sagittal slice.
193
Figure 5-6. Pneumoperitoneal CT 3D renderings of adhesions to bare PP mesh. Yellow = Mesh. Red = Adhesions. Pink = Abdominal Wall. Turquoise = Tissue Markers. A) Mesh viewed straight on. B) Mesh viewed at an angle of 45°.
Figure 5-7. Adhesions to bare PP mesh. Mesh borders are outlined in yellow dashed line. White arrow points to loop of intestine adhered to mesh. A) As initially found, adhesions are anchored over the entire mesh surface. B) After dissecting away some of the less challenging adhesions, the bowel cannot be cleanly separated from the mesh even with applied traction force.
For the pCD-covered PP mesh, adhesions were mostly flimsy, and were entirely restricted to the mesh borders. Interestingly, a loop of bowel had become trapped between the mesh and the abdominal wall, where it had become moderately adherent to the uncoated
PP side and the tissue (both interfaces required blunt dissection to separate). This may have resulted from the absence of fixation elements along the center of each edge of the implant,
194 as meshes were secured with sutures only at the corners. However, no adhesions were seen to form directly on the surface of the pCD barrier, which provided coverage over most of the mesh area in contact with the viscera. These adhesion features were visible using both pneumoperitoneal CT (Figure 5-8, Figure 5-9) and direct visualization (Figure 5-10).
Localization of adhesions to the implant periphery and parietal surface suggest preferential involvement of adhesions with abdominal wall tissues and bare PP face, in comparison with the pCD adhesion barrier. Additionally, the adhesions not associated with the mesh parietal face were divided easily with applied traction. As determined through direct visualization, the mild characteristics of adhesions to the pCD-covered PP mesh were reflected by low scores for coverage, tenacity, and morphology (Table 5-2).
Figure 5-8. Adhesions to pCD-covered PP mesh. Approximate mesh position is marked with yellow dashed line. Blue arrows delineate radiopaque tissue markers. Red arrows delineate loops of intestine. A) Axial slice. B) Sagittal slice.
195
Figure 5-9. Pneumoperitoneal CT 3D renderings of adhesions to pCD-covered PP mesh. Yellow = Mesh. Red = Adhesions. Pink = Abdominal Wall. Turquoise = Tissue Markers. A) Mesh viewed straight on. B) Mesh viewed at an angle of 45°.
Figure 5-10. Adhesions to pCD-covered PP mesh. Mesh borders are outlined in yellow dashed line. A) As initially found. Adhesions are localized to the mesh periphery. B) After dividing and removing the flimsy adhesions. A loop of bowel (white arrow) was observed to have migrated between the mesh and the abdominal wall, becoming adhered to the bare PP side and the tissue. However, no adhesions were found to attach to the smooth pCD-covered area.
196 Table 5-2. Adhesion characteristics as determined using scoring during direct visualization. Group PP Mesh pCD-Covered PP Mesh Scores Scores Adhesion Characteristic Percent Coverage 4 1 Tenacity 3 1 Morphology 4 2 Organ Involvement 2 2 Total 13 6
In agreement with mesh coverage scores determined through direct visualization, image segmentation of pneumoperitoneal CT datasets revealed a decrease in visceral surface coverage for the pCD-covered PP mesh as compared to the bare PP implant (Table
5-3). The measured coverage for the PP mesh using image segmentation was lower on an absolute basis than that determined using direct visualization (i.e. the score of 4 correlates to >75% apparent coverage). However, this likely reflects differences between visualizing a single, thin, essentially 2D, slice of adhesion tissue at a prescribed distance from the mesh surface for CT image segmentation, versus attempting to view a mesh obscured behind 3D adhesion tissues with direct visualization (contrast Figure 5-4 with Figure 5-6A and
Figure 5-9A).
Table 5-3. Adhesion characteristics as determined using image segmentation of pneumoperitoneal CT data. Group PP Mesh pCD-Covered PP Mesh Measurements Measurements Adhesion Characteristic Percent Coverage 34.7% 6.4%
197 Another noteworthy observation from this study was that during implantation of the pCD-covered PP meshes, all pCD barriers cracked along the crease line upon folding for introduction through the incision into the abdominal cavity. Importantly however, upon explantation, no evidence was found to suggest that any portion of the pCD barriers had delaminated from these PP meshes. Despite the presence of a linear crack in each pCD barrier, the pCD layer maintained apparent physical continuity, with no portions of the visceral surface of the PP mesh becoming exposed, and with no obvious missing pieces of pCD.
5.5. Discussion and Conclusions
The primary objective of this study was to preliminarily assess the utility of a new imaging technique, pneumoperitoneal CT, for detection of peritoneal adhesions to surgical mesh implanted within the abdominal cavity. The hypothesis was that pneumoperitoneal
CT enables visualization and detection of large dissimilarities in adhesion characteristics between two diverse mesh prostheses. Findings from this preliminary investigation supported this hypothesis, as pneumoperitoneal CT results were in agreement with those obtained using direct visualization, revealing qualitative and quantitative differences in severity and distribution of adhesions for bare PP versus pCD-covered PP mesh. A secondary purpose of this work was to examine the impact of an experimental adhesion barrier material, pCD, on mesh adhesions. Results suggested that pCD-based adhesion barriers may attenuate adhesions to PP mesh. Overall, further study is needed, however.
In comparison to pneumoperitoneal MRI, the pneumoperitoneal CT method has advantages of improved spatial resolution and compatibility with metallic implants or instruments. One drawback of pneumoperitoneal CT would be radiation exposure of the
198 study subject. Given unique pros and cons for each method, this study does not intend to suggest that either method has greater utility for studying or diagnosing adhesions. Rather, each of the two methods could present unique value for studying (or one day diagnosing) adhesions and experimental goals/logistics (or clinical circumstances) should ultimately dictate selection of one or the other. For instance, some implant materials may produce superior contrast with soft tissue on CT than on MRI, and conversely. Likewise, some researchers may have access to only one type of scanner (CT or MRI). Introducing pneumoperitoneal CT as an alternative option to pneumoperitoneal MRI expands opportunities to study adhesions beyond having one method alone.
The radiopaque tissue markers used in this study enabled unambiguous detection of mesh borders, despite the fact that PP mesh ordinarily cannot be distinguished from soft tissue on CT radiographs. The markers used herein are not the only materials that could be valuable for delineating mesh edges, however. Other materials that might potentially be useful could include ePTFE sutures, or metal tacks used for mesh fixation (though the latter may be impractical in a rodent model given their relatively large size).
This study has a number of limitations. Most importantly, animal sample sizes were extremely small, so results should be interpreted in this context. This low sample size was due to the exploratory nature of this work, and future studies should seek to verify, challenge, or extend beyond findings here, using larger animal sample sizes. A second key shortcoming was that animals were imaged post-mortem rather than alive. Imaging in a live animal would be advantageous in that it would permit longitudinal monitoring of adhesions, possibly enabling more detailed insights into the dynamics of mesh adhesion formation and prevention. However, these advantages come with important technical
199 considerations. For instance, anesthesia and introduction of the needle for insufflation carry risks for animal morbidity and mortality, while gas should be sterile filtered prior to administration. Likewise, for longitudinal studies, desufflation after imaging may be desired as CO2 gas instilled in the rat abdomen has been shown to persist for days and directly impact adhesion formation 178. Gases with lower solubility in blood/tissue, such as nitrogen and oxygen (and air by extension) 845, would persist even longer. Additionally, although live animal imaging has been performed previously with pneumoperitoneal MRI
628, there may be challenges involved in extending this to pneumoperitoneal CT. For example, peristaltic and breathing motions generally produce imaging artifacts, and it might be more difficult to limit the deleterious effects of these artifacts on scan quality with
CT in comparison to MRI. These challenges will need to be overcome in future investigations on adhesions using pneumoperitoneal CT. A third limitation was that this study used only male animals. Differences could be expected in characteristics and features of intra-abdominal adhesions between males and females, whether in rats or humans, given disparities in abdominopelvic anatomy. To ensure thorough pre-clinical evaluation of safety and efficacy of experimental adhesion barrier materials such as pCD, it will be critical to test their efficacy in both genders regardless of animal species. As a final limitation, it should be noted that the relationship between adhesion characteristics measured in animal models (whether by direct visualization or non-invasive imaging), and symptoms in humans, has never been fully established or clarified.
Future investigations should seek to extract additional quantitative metrics regarding adhesions with similar imaging strategies, to include other meshes having adhesion barriers as clinically-used control groups, to enhance extensibility of pCD
200 adhesion barrier materials to reduce likelihood of crack formation, to perform longitudinal monitoring of adhesions in vivo, and to translate such imaging procedures towards clinical application. Regarding the first point, additional quantifiable adhesion metrics that might be derived with similar imaging strategies moving forward could pertain to the total number of adhesions, the detailed composition of adhesion tissue (e.g. vascularity, adiposity), and the impact of insufflation pressure on relative local displacements of the abdominal wall, adhesion tissues, and/or viscera. Concerning the second point, clinically- used meshes that are considered to be resistant to adhesions, such as Bard Sepramesh or
Gore Dualmesh, will be important as control groups in future studies on experimental adhesion barrier materials, such as pCD, so that their efficacy in adhesion prophylaxis can be determined relative to a clinical standard. In terms of the third point, reformulating pCD to provide for increased extensibility would be beneficial for its application as a surgical mesh adhesion barrier. Surgical meshes tend to be highly flexible, and they may undergo substantial deformations, for instance during introduction through a trocar in laparoscopic hernia repair. Such large deformations of a mesh substrate may cause cohesive failure of applied pCD adhesion barriers. This could predispose the material to delamination from the mesh and subsequently result in detrimental outcomes, such as migration of the detached barrier (or its fragments) into anatomical locations where it might cause problems, or adhesions to uncovered portions of the mesh. With regards to the fourth point, longitudinal monitoring is discussed in the paragraph above. Concerning the final point, although this work was very preliminary and only performed in a pre-clinical model, translation of similar imaging strategies towards improving the study or diagnosis of mesh adhesions in humans would seem a worthwhile goal that may one day be possible.
201 Abdominal insufflation is now routine within the realm of surgery (i.e. laparoscopy), so perhaps this technique could be extended towards purposes of visualizing adhesions in clinical radiology.
5.6. Acknowledgements
The authors acknowledge support through National Institutes of Health: NIH R01
GM121477 (HvR) and NIH Ruth L. Kirschstein NRSA T32 AR007505 Training Program in Musculoskeletal Research (GDL). Valuable core facility services were provided by the
Case Center for Imaging Research. The authors also thank Jonah Thomas for technical assistance with animal surgeries, Michael Kavran for training on operation of the Inveon
PET/CT scanner, and Kathleen Derwin for helpful writing feedback and donation of the
Viscus Biologics Radiopaque Tissue Marker.
202 6. CHAPTER 6: CONCLUSIONS AND FUTURE DIRECTIONS
6.1. Conclusions
This body of work has explored the application and potential utility of pCD materials as novel adhesion barriers for PP surgical mesh devices. Key findings are summarized below:
6.1.1. Utility of Plasma for Improving pCD Coatings on PP Substrates
In Chapter 2, nonthermal plasma activation was determined to be a useful approach for enhancing the uniformity (Figure 2-4, Figure 2-5, Table 2-2), mechanical adherence (Table 2-3, Figure 2-6), and covalent attachment (Figure 2-8, Table 2-4) of hydrophilic polymer coatings (in this case pCD) upon otherwise poorly compatible polymer substrates such as PP. For application of pCD as an adhesion barrier for PP meshes, it is key that the pCD not only forms a continuous layer, but also that it makes good contact with, and is firmly connected to, the PP mesh material. Poor coverage or connection with the PP may lead to detachment of the pCD layer from the mesh, which could cause problems in terms of adhesion formation (e.g. to uncovered portions of the PP) or migration of the barrier or its fragments (which could possibly get lodged in anatomical locations that promote adverse consequences). Plasma treatment represents an attractive strategy to improve the quality of the connections between a pCD adhesion barrier and a
PP surgical mesh, as plasma circumvents need for hazardous chemicals as adhesion promoters, which could otherwise present safety concerns for biomedical/biomaterial applications. Future studies should extend the use of plasma toward application of pCD onto other unreceptive polymeric substrates for which it might be useful (Section 6.2.1).
203 6.1.2. Versatility of CD-Based Polymers
Beyond usage as a potential adhesion barrier material, this body of work has also highlighted the incredible versatility of CD-based polymer materials for a wide variety of applications. Importantly, in Chapter 4, results indicated that CD-based polymers may be used to mitigate events of biofouling, such as protein adsorption (Figure 4-4), mammalian cell attachment (Figure 4-5), and bacterial attachment (Figure 4-6). This is a new and meaningful finding as it could widen the application range of these materials substantially.
CD-based materials already find use in broad applications due to their unique affinity properties, which enable them to achieve sustained release or efficient sorption of small-molecule compounds, and even be recycled for these processes multiple times. The
ABF properties demonstrated by pCD herein further expands the potential uses for CD- based materials. Furthermore, it implies that there may be applications for which this class of materials may be uniquely suited, such as those involving combined release or uptake of small molecules in addition to passive biofouling resistance. Examples of this could include an adhesion barrier for surgical meshes that can simultaneously release antibiotics or analgesics while mitigating adhesions, or a coating for water treatment filters that concurrently scavenges pollutants while resisting accumulation of bacterial biofilms.
In order to support the pursuit and realization of such niche applications, future studies should seek to determine how complexation of various small molecule compounds into the pockets of CD subunits may impact the passive resistance of CD-based polymers to various biofouling events (Section 6.2.2).
204 6.1.3. Tradeoffs with pCD HDI Crosslinking: Biofouling Resistance versus Mechanical
Robustness
In Chapter 4, it was found that reduced HDI crosslinking tended to increase wettability of pCD (Figure 4-2), as well as the ability of pCD materials to resist biofouling events such as protein adsorption (Figure 4-4), mammalian fibroblast attachment (Figure
4-5), and attachment by two diverse species of bacteria (gram-negative E. coli and gram- positive S. aureus) (Figure 4-6). This suggests that for a pCD adhesion barrier, a lower degree of HDI crosslinking may be preferable for maximal resistance to adhesion formation. However, decreased crosslinking using HDI also produced CD-based polymers having lower resistance to physical deformation (i.e. lower rigidity) (Figure 4-2). This decrease in rigidity is not considered to be detrimental to the success of pCD as a mesh adhesion barrier, given that the purpose of such an adhesion barrier component is typically not structural (i.e. the mesh component satisfies the mechanical needs of the device).
However, there could be other potential applications for pCD in which a rigid ABF coating would be preferable to a softer one, such as an ABF paint on a ship hull that must withstand years of fluid drag and intermittent mechanical abrasion. Future studies are needed to identify alternative pCD formulations that can remain resistant to biofouling events regardless of the degree of crosslinking (Section 6.2.3).
6.1.4. Tradeoffs with Plasma Treatment: PP Surface versus Bulk Properties
In Chapter 2 and Chapter 3, increasing plasma treatment duration beyond the shortest non-zero exposure time (1 min) was generally found not to provide major benefits in terms of pCD spreading upon PP (Figure 2-4, Figure 2-5, Table 2-2), resistance of PP to fibrinogen adsorption (Figure 3-5) and E. coli attachment (Figure 3-6), and ability of
205 PP to support mammalian fibroblast attachment (Figure 3-7). This was at least the case among the range of exposure durations tested. At the same time, in Chapter 3, several mechanical properties of PP mesh, such as uniaxial tensile strength (Figure 3-11, Table
3-2), suture retention strength (Figure 3-12), and ball burst strength (Figure 3-14, Table
3-5), declined with increasing plasma treatment duration.
Though there are currently few generally-accepted requirements for the mechanical properties of surgical meshes (Section 3.5), it is important to recognize that these devices serve structural purposes. For instance, inadequate mesh suture retention strength could lead to suture pull-out at mesh fixation points upon sudden stress or strain of the abdominal wall (e.g. when the implant recipient coughs or twists), potentiating hernia recurrence and mesh migration. Thus, it is critical to be aware of, and seek to minimize, potential impacts of device surface modifications on mesh mechanical properties.
Altogether, the results from Chapter 2 and Chapter 3 imply that short plasma treatment durations (i.e. ~1 min using the materials/system described here) were most optimal for providing reasonable improvement in both pCD spreading upon PP and biofouling resistance of PP, with minimal impact on the substrate mechanical properties.
Given these findings with PP, future studies are needed to determine impacts of nonthermal plasma exposure duration on the surface and bulk properties of other mesh substrate materials, as well as the effects of different plasma carrier gases (Section 6.2.4).
6.1.5. Ability of Plasma Treatment to Directly Modulate Protein Adsorption, Fibroblast
Attachment, and Bacterial Attachment to PP Substrates
In Chapter 3 and Chapter 4, plasma treatment was shown to directly decrease protein adsorption (Figure 3-5, Figure 4-4), increase fibroblast attachment (Figure 3-7,
206 Figure 4-5), and decrease attachment of E. coli to PP substrates (Figure 3-6, Figure 4-6).
These findings are important as they suggest that there may be a possibility for plasma treatment to modulate events that can have a major impact on the ultimate success of an implanted surgical mesh, such as adhesion formation, mesh incorporation, and prosthetic infection. This highlights a need for future research in this area (Section 6.2.5).
6.1.6. Value of pCD as a PP Mesh Adhesion Barrier, and of Pneumoperitoneal CT as a
Method for Visualizing Adhesions to Intra-Abdominal Mesh
In Chapter 5, preliminary evidence suggested that pCD-based adhesion barriers were able to attenuate adhesions to PP mesh (Figure 5-5, Figure 5-6, Figure 5-7, Figure
5-8, Figure 5-9, Figure 5-10, Table 5-2, Table 5-3), and that pneumoperitoneal CT could permit qualitative visualization of mesh adhesions (Figure 5-5, Figure 5-6, Figure 5-8,
Figure 5-9) and quantitative measurement of adhesion characteristics (Table 5-3). Further, these measurements fell in agreement with those obtained using consensus scoring upon direct visualization (Table 5-2). However, this investigation had several limitations
(Section 5.5), and there is a great deal of room for future studies to compare pCD-based adhesion barriers to clinically-used standards and to refine the imaging method (e.g. to permit longitudinal monitoring and derivation of additional adhesion metrics) (Section
6.2.6). Additionally, pCD barriers sustained cracking upon folding of meshes when introducing them into rat abdominal cavities, highlighting a need to develop new pCD formulations that can withstand larger deformations without cohesive failure (Section
6.2.3).
207 6.2. Future Directions
This work has highlighted several unexplored or underexplored research directions that could represent worthwhile next steps. These could include any of the following:
6.2.1. Application of pCD Coatings to Other Polymeric Substrate Materials
As described in Section 6.1.1, future studies should aim to expand the use of nonthermal plasma toward application of pCD materials onto unreceptive polymeric substrates (other than PP) for which pCD might provide beneficial functions. This could include ePTFE, another common biomaterial used in surgical meshes. This material is inert and hydrophobic, making it a potentially difficult substrate for successful coating application in the absence of a process to activate the surface. Furthermore, meshes made from ePTFE are notorious for their propensity to need removal upon prosthetic infection, however pCD coatings may be useful for counteracting this drawback, given their prior success in resolving PETE mesh infections 639,640.
208 6.2.2. Evaluation of Drug-Loading Impacts on pCD Biofouling Resistance
As indicated in Section 6.1.2, one research area worth exploring entails investigating the effects of CD subunit complexation (e.g. with various small molecule compounds) on the passive ABF properties of pCD. Interestingly, in Chapter 4, incorporation of the antibiotic compound rifampin, which has poor water solubility, actually improved wettability of pCD materials (Figure 4-2). It is difficult to ascertain the exact mechanism for this phenomenon. Potential explanations include that the drug may have diffused into the water droplet and acted as a surfactant, or that the relatively nonpolar and polar regions of the rifampin were respectively localized to the interior and exterior of
CD pockets. Improved wettability of materials can potentially suggest enhanced passive resistance to non-specific protein adsorption, though this generalization may not always hold true. Regardless, further studies, which are well-designed, are needed to directly determine effects of complexation of various compounds on passive biofouling resistance of pCD. One potential difficulty in exploring this research topic would be that compounds may diffuse out of CD subunits during experiments. These free molecules might then affect the propensity of biofoulants to attach to pCD surfaces. For example, the released compounds may alter viability of biofoulants (e.g. cells, bacteria), making it difficult to ascertain the degrees to which ABF performance is attributed to passive versus active processes. However, assuming these effects are additive, if pCD can successfully combine passive resistance to biofouling events with active, sustained, and multi-window release of biocidal compounds, CD-based polymers might represent the ultimate ABF materials, attaining theoretically unmatched levels of biofouling resistance.
209 6.2.3. Reformulation of pCD to Achieve Crosslinking-Independent Resistance to
Biofouling, and/or Greater Extensibility
As suggested in Section 6.1.3 and Section 6.1.6 respectively, future work should attempt to develop pCD materials with mechanical properties that first, can be tuned without impacting passive ABF properties, and/or second, which have greater extensibility.
With regards to the first point, in Chapter 4, increased HDI crosslinking of pCD materials reduced their wettability, and decreased their resistance to fouling by proteins, mammalian cells, and bacteria. Crosslinking agents which are less nonpolar than HDI overall may permit decoupling of mechanical properties and wettability in future pCD materials.
Preliminary efforts had been attempted to formulate pCD with another crosslinking agent, ethylene glycol diglycidyl ether (EGDGE) (Figure 6-1), which is mainly used in our lab to produce pCD in the form of microparticles. However, these pCD materials were generally found to possess insufficient cohesive strength, as they would generally fragment upon rehydration with water. This was the case for EGDGE-crosslinked pCD applied as coatings for TCPS dishes (Figure 6-2) and PP meshes (Figure 6-3). Likewise, EGDGE- crosslinked pCD had also been cast in the form of cylinders, which swelled and ruptured when submerged in several polar solvents: water, DMF, and dimethylsulfoxide (data not shown). Though in no way definitive, these preliminary findings discouraged further exploration of EGDGE-crosslinked pCD materials for mesh adhesion barrier applications, narrowing selection to HDI. In support of our observations, others previously reported that
EGDGE-crosslinked pCD materials were mechanically fragile to the extent that tensile testing of them was deemed unfeasible 824.
210
Figure 6-1. Chemical structure of EGDGE. Figure created using ChemDraw Professional software (PerkinElmer Informatics, Inc).
Figure 6-2. EGDGE-crosslinked pCD coating on a 60 mm TCPS dish. The coating mixture was prepared by dissolving 250 mg epichlorohydrin-crosslinked β-CD polymer precursor in 1 mL of 0.2 M potassium hydroxide solution, and mixing with 38.4 µL EGDGE to achieve an approximate crosslink molar ratio of 0.16 (EGDGE per glucose residue). The mixture was pipetted and spread uniformly on the dish, then covered and allowed to cure for 1 day. A) As found. B) Two hours later, after pipetting water into the dish and allowing polymer hydration, the coating has fragmented into several pieces, which are suspended in the liquid. C) Dish tilted to demonstrate movement of suspended coating fragments.
211
Figure 6-3. EGDGE-crosslinked pCD coating on an untreated PP Prolene mesh. The coating mixture was prepared by dissolving 1 g epichlorohydrin-crosslinked β-CD polymer precursor in 3 mL of 0.2 M potassium hydroxide solution, and mixing with 154 µL EGDGE to achieve an approximate crosslink molar ratio of 0.16 (EGDGE per glucose residue). The PP mesh was then laid flat on a piece of Parafilm, and the mixture was pipetted through the mesh interstices, forming a liquid layer between the Parafilm and mesh. The coating was allowed to polymerize for 5 days at room temperature. A) Rehydrated coating on the mesh. Dashed lines of yellow box indicate approximate boundaries of zoomed-in region for panels B and C. B) Before dipping the mesh (black arrow) into a beaker of clean water. C) Immediately after dipping the mesh into the water, small coating fragments separate from the mesh, becoming suspended in the liquid (dashed red circle).
With regards to the second point, in Chapter 5, cracks formed in HDI-crosslinked pCD adhesion barriers when pCD-covered PP meshes were folded to introduce them into rat abdominal cavities. This mismatch in the flexibility between an HDI-crosslinked pCD layer and a PP mesh substrate had been observed in preliminary experiments as well
(Figure 6-4). While the present body of work has identified strategies to improve the
212 connection between pCD and PP materials, this unfortunately does not preclude pCD cohesive failure. Ideally, a pCD adhesion barrier would deform in parallel with the mesh substrate without sustaining fractures, as cohesive failure would be detrimental for a few reasons. For one, cracks which initiate in the pCD layer can easily propagate, extend along the pCD-PP interface, and lead to delamination of pCD from the PP substrate. For two, defects in the pCD layer may serve as sites where the PP substrate becomes unprotected, facilitating nucleation of adhesions at these locations. The most substantial deformation of meshes is likely to occur either during surgical introduction (e.g. being rolled up and pushed through a trocar) or potentially during certain abdominal movements of the patient/recipient (e.g. twisting). Adhesion barriers for surgical mesh devices placed within the peritoneum will ideally need to withstand these events without cohesive failure.
Figure 6-4. Tensile test of untreated PP Prolene mesh with HDI-crosslinked pCD layer on one side. A) At beginning of test. The dashed yellow line serves as a visual reference for the distance between the top and bottom grips. B) Very shortly after cohesive failure of the pCD layer. The dashed line is 14% longer than it had been in panel A, as measured using image analysis. C) Immediately following mesh rupture. The dashed line is 58% longer than in panel A, highlighting a large difference in flexibility for the HDI-crosslinked pCD layer versus the PP mesh substrate. Note that increases in dashed line length here are not equivalent to strain, given that photos were taken at a non-orthogonal angle to the sample, and that the sample width was not constant over the entire length between the grips given the ASTM dogbone geometry.
213 Strategies that could improve extensibility of pCD could include either replacing a fraction of the covalent crosslinks with non-covalent connections, such as linkages based on host-guest 846,847 or ionic 848 interactions, or increasing the spacer segment length of the covalent crosslinks 849. Potential downsides to incorporation of host-guest-mediated linkages include that this strategy may partly abolish the loading capacity and affinity properties of CD-based polymers as pockets of CD subunits become occupied, and conversely polymer structural integrity may be disrupted as small molecule compounds compete with these linkages.
Preliminary efforts had been attempted to enhance pCD extensibility through increasing spacer segment length of covalent crosslinks. This was done by a two-step synthesis procedure in which PEG (various molecular weights) was first reacted with HDI at a 1:2 ratio to produce HDI-PEG-HDI (Figure 6-5), then this was used as a crosslinking agent for soluble, lightly epichlorohydrin-crosslinked β-CD polymer precursor. PEG was chosen as a chain extender because its use in synthesis of HDI-PEG-HDI has been reported previously 849, and incorporation of PEG was considered to have potential to improve both extensibility and biofouling resistance of pCD. Some promising results had been attained using PEG at a number average molecular weight of 600 Da; specifically, in one experiment each, pCD crosslinked at equivalent molar ratios with HDI versus HDI-PEG-
HDI ruptured at tensile failure strains of <20% (Figure 6-6) versus >250% (Figure 6-7), respectively. Such increases in extensibility may reflect not only increased molecular length of crosslinks, but potentially also some degree of polypseudorotaxane formation, as
PEG-based crosslinker threading through CD subunits was previously shown to enhance elongation at break of pCD materials 824. However, with all our investigated HDI-PEG-
214 HDI crosslinked formulations, there were often challenges with polymer phase separation
(one region of the material was stretchy while another was not) and results between experiments tended to be highly variable (data not shown).
Figure 6-5. Synthesis of HDI-PEG-HDI as a long-chain covalent crosslinker for improving the extensibility of pCD. Figure created using ChemDraw Professional software (PerkinElmer Informatics, Inc).
Figure 6-6. Tensile test of pCD crosslinked using HDI at a molar ratio of 0.64 HDI per glucose residue. A) At beginning of test. B) Immediately prior to rupture. Image analysis using ImageJ indicated an ultimate strain value of 10.2%. 215
Figure 6-7. Tensile test of pCD crosslinked using HDI-PEG-HDI at a PEG Mn of 600 Da, and a molar ratio of 0.64 HDI-PEG-HDI per glucose residue. A) At start of test. B) Immediately prior to rupture. Image analysis using ImageJ indicated an ultimate strain value of 289.3%.
6.2.4. Investigation of Impacts of Plasma Exposure and Carrier Gas on Different Mesh
Materials
As mentioned in Section 6.1.4, future studies should seek to evaluate effects of exposure to nonthermal plasma of other polymers used for surgical meshes apart from PP, and effects of plasma carrier gases other than Ar/H2O, on the surface and bulk properties of these substrates. While PP is the most common surgical mesh material 46–50, PETE and
PTFE represent the second and third most frequently used polymers used in mesh devices.
Plasma could potentially be useful for directly or indirectly modifying surface properties
216 of either of these materials, and it remains to be determined whether plasma-induced embrittlement would be seen with these substrates as well. Likewise, it may be worthwhile to assess the impact of different plasma carrier gases, such as N2 or NH3, on the surface chemistry and mechanical properties of mesh devices. Plasma produced using these gases may permit engraftment of other potentially useful functional groups, such as amines, onto polymeric mesh surfaces. This could enable alternative strategies to covalently link coatings onto these substrates.
6.2.5. Exploration of the Impacts of Plasma Treatment on Mesh-Related Complications
As highlighted in Section 6.1.5, considering that findings from Chapter 3 and
Chapter 4 revealed the potential for nonthermal plasma to directly reduce protein adsorption, increase fibroblast attachment, and decrease bacterial attachment to PP, future investigations should assess the potential for plasma treatment to modulate mesh-related complications. Reduction in protein adsorption to PP mesh may reflect a decrease in protein denaturation, which might lead to a subsequent reduction in inflammation and mesh adhesion formation. Increased fibroblast attachment to PP mesh may predict improved incorporation of mesh with the abdominal wall, thus preventing mesh migration. Finally, reduced attachment of bacteria to PP mesh may reduce risk of, or simplify treatment of, prosthetic infection. To our knowledge, no animal or human studies to date have reported correlations between any of these biofouling events and a mesh-related complication, meaning each could represent an unexplored avenue for discovery.
217 6.2.6. Further Studies on Adhesion Prophylaxis Using pCD Barriers, and Refinement of
Pneumoperitoneal CT
As touched upon in Section 6.1.6, future work should more rigorously assess the effectiveness of pCD adhesion barriers in deterring mesh adhesions and further refine the imaging approach. In particular, pCD materials should be evaluated in animal studies having larger sample sizes (to allow for statistically meaningful comparisons), both animal genders, and clinically-used mesh adhesion barrier products (e.g. Sepramesh) included as control groups. With regards to improving upon the imaging method, longitudinal monitoring in live animals and/or potential derivation of additional adhesion metrics (such as detailed composition of adhesion tissues) should be attempted with pneumoperitonal
CT. Additionally, larger sample sizes should be used to validate/compare quantitative measurements of adhesion characteristics made through the imaging technique to consensus scoring metrics obtained through direct visualization.
218 APPENDIX
Permission for Reprint of Chapter 2
Permission for Reprint of Chapter 3
219 Permission for Adaptation of Table 5-1
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