Contents
Cover
Half Title page
Title page
Copyright page
Preface
Part 1: Introduction
Chapter 1: Historical Perspectives on Biomedical Coatings in Medical Devices
1.1 Introduction
1.2 Improving Physical Properties of Biomaterials: Hydrophilic, Lubricious Coatings
1.3 Modulating Host-Biomaterial Interactions: Biologically Active Coatings
1.4 Bioinert Coatings Redressed? Nonfouling Coatings
1.5 Future Biomedical Coatings
References
Part 2: Coating Applications
Chapter 2: Antimicrobial Coatings and Other Surface Modifications for Infection Prevention
2.1 Introduction
2.2 Genesis of Device-Related Infections
2.3 Antimicrobial Coatings
2.4 Non-Eluting Antimicrobial Surfaces
2.5 Coating and Surface Modification Technologies
2.6 Regulatory Considerations
2.7 Future Challenges
References
Chapter 3: Drug Delivery Coatings for Coronary Stents
3.1 Introduction
3.2 Polymer Coatings for DES
3.3 Biostable (Non-Bioabsorbable) Polymers
3.4 Bioabsorbable Polymers
3.5 Concluding Remarks
References
Chapter 4: Coatings for Radiopacity
4.1 Principles of Radiography
4.2 Use of Radiopaque Materials in Medical Devices
4.3 Radiopaque Fillers
4.4 Types of Radiopaque Fillers
4.5 Other Radiographic Materials and Coating Systems
4.6 Radiopaque Coatings by Physical Vapor Deposition
4.7 Challenges in Producing Radiopaque Coatings Using PVD
4.8 Gold Radiopaque Coatings
4.9 Tantalum Radiopaque Coatings
4.10 Summary
References
Chapter 5: Biocompatibility and Medical Device Coatings
5.1 Introduction
5.2 Challenges with Medical Devices
5.3 Examples of Products Coated to Improve Biocompatibility
5.4 Types of Biocompatible Coatings
5.5 Commercialization
5.6 Summary
References
Chapter 6: Tribological Coatings for Biomedical Devices
6.1 Introduction
6.2 Hard Thin Film Coatings for Implants
6.3 Binary Carbon-Based Thin Film Materials: Diamond, Hard Carbon and Amorphous Carbon
6.4 Progress of DLC, ta-C and a-C:H Films for Hip and Knee Implants
6.5 Wear-Resistant Coatings for Stents and Catheters
6.6 Wear-Resistant Coatings for Angioplasty Devices
6.7 Scalpel Blades and Surgical Instruments
6.8 Multifunctional, Nanostructured, Nanolaminate, and Nanocomposite Tribological Materials
References
Part 3: Coating and Surface Modification Methods
Chapter 7: Dip Coating
7.1 Description and Basic Steps
7.2 Equipment and Coating Application
7.3 Coating Solution Containers
7.4 Coating Parameters and Controls
7.5 Role of Solution Viscosity
7.6 Coating Problems
7.7 Process Considerations
Chapter 8: Inkjet Technology and Its Application in Biomedical Coating
8.1 Introduction
8.2 Inkjet Background
8.3 Equipment Used
8.4 Capabilities
8.5 Limitations and Ways around Them
8.6 Manufacturing Advantages and Future Directions
8.7 Conclusions
References
Chapter 9: Direct Capillary Printing in Medical Device Manufacture
9.1 Introduction
9.2 Fundamental Elements of Direct Capillary Printing
9.3 Practical Operational Considerations
9.4 Manufacturing Considerations
9.5 Medical Device Examples
9.6 Conclusions
Acknowledgments
References
Chapter 10: Sol-Gel Coating Methods in Biomedical Systems
10.1 Introduction
10.2 Overview of Sol-Gel Coatings in Biomedical Systems
10.3 The Sol-Gel Process
10.4 Coating Methods and Processes
10.5 Factors Influencing Coatings Characteristics/Performance
10.6 Summary and Concluding Remarks
References
Chapter 11: Chemical Vapor Deposition
11.1 Introduction
11.2 Process Description
11.3 Process Mechanism
11.4 Technology Advances
11.5 Future Outlook
References
Chapter 12: Introduction to Plasmas Used for Coating Processes
12.1 Introduction
12.2 DC Glow Discharges
12.3 RF Glow Discharges
12.4 RF Diode Glow Discharges
12.5 Ionization in RF Diode Glow Discharges
12.6 Inductively Coupled RF Discharges
12.7 Mid-Frequency AC Discharges
12.8 Pulsed DC Discharges
12.9 Comparison of Plasma Properties
12.10 Plasma Species
12.11 Summary
References
Chapter 13: Ion Implantation: Tribological Applications
13.1 Introduction
13.2 Applications
13.3 Nanocrystalline Diamond
Reference
Chapter 14: Plasma-Enhanced Chemical Vapor Deposition
14.1 Introduction
14.2 Process Description
14.3 Plasma Effects on Materials Deposition
14.4 Future Outlook
References
Chapter 15: Sputter Deposition and Sputtered Coatings for Biomedical Applications
15.1 Introduction
15.2 Overview of Sputter Coating
15.3 Characteristics of Sputtered Atoms
15.4 Sputtering Cathodes
15.5 Relationship between Process Parameters and Coating Properties
15.6 Biased Sputtering
15.7 Adhesion and Stress in Sputtered Coatings
15.8 Sputtering Electrically Insulating Materials
15.9 Recent Developments
15.10 Summary and Conclusions
References
Chapter 16: Cathodic Arc Vapor Deposition
16.1 Introduction
16.2 Medical Uses of Cathodic Arc Titanium Nitride Coatings
16.3 Brief History and Commercial Advancement of Cathodic Arcs
16.4 Review of Arc Devices
16.5 Description of PVD Coating Manufacturing
16.6 Macroparticle Generation and Mitigation
16.7 Considerations for Coating Success
16.8 Materials Used in Biomedical PVD Coatings
References
Part 4: Functional Tests
Chapter 17: Antimicrobial Coatings Efficacy Evaluation
17.1 Introduction
17.2 In-Vitro Methods
17.3 In-Vivo (Animal) Methods
17.4 Equipment and Laboratory Resources
17.5 Human Clinical Trial Considerations
17.6 Regulatory Considerations
References
Chapter 18: Mechanical Characterization of Biomaterials: Functional Tests for Hardness
18.1 Introduction
18.2 Basic Principles of Hardness and Indentation Testing
18.3 Depth-Sensing Indentation Testing
18.4 Dynamic Indentation Testing: A More Advanced Hardness Measurement Technique for More Complex Material Behavior
18.5 Special Case of Coatings Configuration Under Indentation Testing
18.6 Conclusions
References
Chapter 19: Adhesion Measurement of Thin Films and Coatings: Relevance to Biomedical Applications
19.1 Introduction
19.2 Mechanical Test Methods of Adhesion Measurement
19.3 Summary and Remarks
Appendix
References
Chapter 20: Functional Tests for Biocompatibility
20.1 Introduction
20.2 Inflammation
20.3 Blood Compatibility
20.4 Wound Healing
20.5 Encapsulation
20.6 Tissue Integration
20.7 Vascularization
20.8 Toxicity
20.9 Infection
20.10 When to Move In Vivo?
References
Chapter 21: Analytical Requirements for Drug Eluting Stents
21.1 Introduction
21.2 Instrumentation
21.3 API and Excipient Characterization
21.4 Analytical Methods
21.5 Conclusion
References
Part 5: Regulatory Overview
Chapter 22: Regulations for Medical Devices and Coatings
22.1 Introduction
22.2 Types of Regulated Products
22.3 Scope of Regulation
22.4 Marketing Clearance of Medical Devices
22.5 Comparison to EU Regulation
22.6 Good Manufacturing Practices
Part 6: Future of Coating Technologies
Chapter 23: The Future of Biomedical Coatings Technologies
23.1 Introduction
23.2 The Continuing Evolution of Biomaterials
23.3 Tissue Engineering and Regenerative Medicine
23.4 Coating Process Development
References
Index
Medical Coatings and Deposition Technologies
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Preface
Medical devices sit at the crossroads of diverse disciplines such as chemistry, materials science, mechanical and biomedical engineering not to mention biology and medicine. Coatings technology evolved via empirical trial and error, but the past few decades have shown a vast increase in the understanding of the applicable principles and elucidation of the underlying science and engineering. The successful development of a coated medical device involves a coordinated effort amongst practitioners in these fields and usually is undertaken to address an unmet clinical need. A coating is provided on a device to either enhance surface properties or impart functionality to better serve the rationale for the design of the device. The point of this book is to showcase examples of how coatings are employed in the enhancement and functioning of medical devices and also survey the manufacturing techniques deployed to produce them. Each chapter focuses on an example of how coatings impart critical functionality to the device or espouses upon the technology that makes them possible. The various coating methods are also covered in detail so that a reader with an interest in design of medical devices may learn the options available to provide a coating and gain understanding from the commercialized examples detailed within the text.
The scope of medical coatings is so great that one of the most significant challenges we faced was deciding what to include and what to leave out. Our goal was to give the reader enough background in the most widely used types of coatings, coating processes and test methods to provide an understanding of the fundamentals and enough references to guide further reading. We also wanted to include some of the context within which medical devices are produced and have done so in the chapters on historical perspectives and regulatory issues. Finally, we included a perspective on the future of coatings technologies as applied to medical devices in particular. We hope this book is of interest to scientists and engineers in the medical device arena as well as students of biomedical engineering and science who would like to learn more about industrial applications of their art.
A work like this, of course, is a product of the contributors. We would like to thank the many experts in their fields who devoted their time and energy to creating this book. We are grateful to Martin Scrivener (our long suffering publisher) for his patience and steady encouragement that made this project possible – we largely underestimated the time commitment needed to herd authors, juggle careers and of course families over the course of several months. We would also like to acknowledge and thank Aparna Bhave for her early work on the project especially in its infancy.
Shrirang V. Ranade & David A. Glocker April 10th, 2016.
In this chapter we will discuss the evolution of biomedical coatings from the perspective of three generations of biomaterials: From the initial emphasis on “do no harm” bioinert coatings to second and third generation biomedical coatings having increasingly more activity and the ability to obtain a beneficial response through interaction with the biological environment.
After a half century of effort, the arc of progress in biomedical coatings in medical devices bends towards positive clinical outcomes. Indisputable demonstration that better management of the material’s biointerface yields positive impact on clinical outcome remains the most challenging goal and need for the future.
We conclude that for biomedical coatings to truly deliver such evidence of clinical outcome improvement it is of quintessential importance to step away from the tendency to focus on individual aspects and rather focus on the total picture of the clinical problem. In other words, to bring together a multifactorial approach to surface science and the cellular and molecular pathophysiology of implanted devices, so as to optimally calibrate the clinical impact that surface modification of medical devices will have.
Keywords: Medical coatings, surface modification, biocompatibility, blood compatibility, biomaterials, medical devices, heparin, antimicrobial
The use of biomedical materials has a long and fascinating legacy characterized by creativity, innovation and positive medical outcomes. Since the dawn of civilization mankind has been exploring ways in which natural materials might replace or enhance the natural functions of the human body. Archaeological evidence suggests, for example, that the ancient Egyptians used seashells to replace missing teeth and linen to close wounds, creating what may have been the first dental implants and sutures. Natural materials have been the key source of biomedical materials throughout much of the history of their use, from coconut shells used to close holes in skulls to elephant ivory that was used to create the first recorded hip implant in 1891.
The real revolution in biomedical materials began in the 20th century, with the introduction of synthetic materials that enabled medical device makers to break free from many of the limitations and risks associated with relying solely on natural materials. For example [1], polymethylmethacrylate (PMMA) was used in dentistry in the 1930s and cellulose acetate was used in dialysis tubing in the 1940s. Dacron was used to make vascular grafts; polyetherurethanes were used in artificial hearts; and PMMA and stainless steel were used in total hip replacements. Characteristically, these materials were brought in from other areas of science and technology without substantial redesign for their clinical use purpose. Their choice was based on achieving a suitable combination of physical properties to match those of the replaced tissue with a “biopassive,” minimal toxic response in the host. However, while these materials did enable the development of new medical treatments, critical problems including biocompatibility, thrombogenicity, fibrous encapsulation, infectious complications and biostability—oxidative and hydrolytic degradation—remained. A quest for the perfectly inert material, harmless to the host tissue environment ensued. This “do no harm” paradigm is best illustrated by the way professor David Williams defined biocompatibility [2]: “The ability of a material to perform with an appropriate host response in a specific application.”
Scientific efforts started to be focused on engineering the materials’ surfaces, both involving physical and chemical methods. Baier disclosed a number of surface property concepts that were hypothesized to be beneficial [3]. The majority of surfaces he listed are in essence attempts to create surfaces with minimal effect, termed either low protein or platelet adhesion. Despite a tremendous amount of scientific and technological effort, the quest for the holy grail of inert materials turned out to be “20 years of frustration” [4].
The mid-70s’ emergence of molecular biology analytical tools, like RIA and ELISA, opened up the possibilities of looking at the molecular and cellular aspects involved at the surface. With the consequent increasing understanding of the pathophysiology of host-material interactions at the cell and molecular level, the field of biomaterials moved toward emphasis on improved management of the material’s biointerface. Rather than trying to exclusively achieve the bioinert response, it instead moved to pursuit of strategies aimed at optimizing the biological interactions with the synthetic material. An example is the incorporation of bioactive components that could elicit a controlled action and reaction in the physiological environment. Very prominent examples of these second-generation “bioactive” biomaterials are heparin coatings for improved blood compatibility, and drug-eluting stent coatings for prevention of vascular restenosis.
Now third-generation biomaterials are being designed to stimulate specific cellular responses at the molecular level. Taking contemporary understanding of molecular and cell biology further, biology is incorporated into materials design: molecular modifications of polymer systems elicit specific interactions with cell surface integrins and thereby direct cell proliferation, differentiation, and extracellular matrix production and organization. These third-generation “bio-interactive” biomaterials stimulate regeneration of living tissues.
Circling back to the definition of biocompatibility, some 20 years after his original definition, the same professor, David Williams, revised the original definition of biocompatibility [5] to now read: “The ability of a biomaterial to perform its desired function with respect to a medical therapy, without eliciting any undesirable local or systemic effects in the recipient or beneficiary of that therapy, but generating the most appropriate beneficial cellular or tissue response in that specific situation, and optimizing the clinically relevant performance of that therapy.” Clearly this reflects the evolution of biomaterials into second and third generations; those materials having increasingly more activity and interaction with the biological environment. With regard to biocompatibility, next generations of biomaterials not only focus on “do no harm,” but actually have the ability to obtain a beneficial response.
The use of hydrophilic coatings on medical devices has probably the longest clinical use history. More recently their increasing use appears to go hand in hand with medical devices getting smaller and smaller. Various clinical fields increasingly turn to percutaneous therapies or minimally invasive procedures. The reasons are obvious: Smaller devices help to reduce trauma, speed recovery time, and shorten hospital stays, all contributing to reducing the burden on our healthcare systems, thus at the end of the day reducing costs.
The majority of substrates found in medical devices are hydrophobic with surface energies generally spanning the range between 42 and 20 mN/m [6]. For a surface to wet perfectly and thus assist lubrication, its surface energy should be above that of the wetting fluid. Knowing that the surface tension of water at room temperature is around 72 mN/m and that of blood at body temperature around 52 mN/m [7], most medical materials are thus poorly suited to many medical applications in the absence of suitable surface modification [8]. With poor wetting and self-lubrication, an invasive device can compromise the host’s health due to tissue trauma and subsequent infection. Self-lubrication of medical devices, such as catheters and guidewires, can be imparted by applying a hydrophilic coating. Hydrophilic coatings are beneficial in reducing trauma damage. The coating’s increased surface energy facilitates wetting and provides the ability to tune lubricity (reduce coefficient of friction) in an aqueous environment. These features can be leveraged to improve the usability or performance of a device.
Polyethylene glycol (PEG) and polyvinyl pyrrolidone (PVP), along with various PEG derivatives and PVP copolymers, are the most common polymers used in hydrophilic lubricious coatings. Further, derivatives of phosphoryl choline, such as 2-methacryloyloxyethylphosphorylcholine (MPC) in acrylic copolymers and water-soluble polyvinylethers, are used, as well as the natural polymer hyaluronic acid [6].
Having been extensively tested and used as a plasma substitute, PVP has an extensive safety record [9]. It has demonstrated to be well suited for short-term device applications and is the key ingredient in market-leading hydrophilic lubricious coatings of Surmodics (Harmony® advanced lubricity coatings) and DSM (ComfortCoat® hydrophilic coating). Both use photo-initiated crosslinking, grafting and entanglement as the method of producing their PVP-based coatings, taking advantage of the proton abstraction capability of type II photoinitiators and in particular the labile alpha-hydrogen found on the backbone of PVP. While the technological approach differs, both aim to immobilize a film of a hydrophilic polymer on a substrate, whereby the resulting coating swells in water to fulfill its function as a hydrogel, yet remains sufficiently well bound to the substrate for the application in question. The latter is an important aspect. Good adhesion of the hydrophilic coating is essential. Vascular devices, for instance, travel through narrow, tortuous blood vessels. Poor adhesion can lead to shedding of the coating and particles breaking off, thus increasing the risk of embolism. This potential risk has most recently come under increased scrutiny by the FDA, which has lead to the issuing of a new ASTM standard guide for coating inspection and particulate characterization [10]. This has in turn triggered coating manufacturers to put more emphasis on developing improved coating formulation chemistries with the emphasis on providing more durable, “low particulates” hydrophilic coatings.
Near the end of the 20 years of frustration, in the early 1980s, Larm introduced covalently bound heparin surfaces [11, 12]. The biological activity of this surface was convincingly demonstrated in direct comparison to the biological activity of rabbit aorta by Boris Pasche MD in his PhD thesis in 1989 [13–15]. At this time heparinized surfaces were understood mainly in terms familiar to Gott’s 1963 ionic heparin surfaces [16]. It was believed that heparin surfaces could only work by releasing the “drug” into the blood. It took some time for the scientific community at large to accept that the covalently bound surface heparin was truly bioactive. The concept and approach of bioactive surfaces was generally accepted by 1996 [17]. However, the problem remained that, for the most part, coagulation was the focus of bioactive surfaces; and even though the covalently bound heparin surfaces demonstrated decrease in inflammatory complications in cardiopulmonary bypass (CPB) [18], its value was still measured in its ability to prevent thrombin formation in the clinical setting [19]. This “natural scientific tendency to focus on individual aspects of the whole problem rather than considering the various interactions …” is discussed by Gorbet and Sefton in a critically important review article that should reset many clinical approaches to determining efficacy of surface modifications of biomedical devices [20]. This increased understanding of the interaction of coagulation and inflammation together with additional multifactorial variables, such as priming volume, closed circuits, separation of cardiotomy suction, inflammatory inhibitors, as well as biocompatible coatings, shows promising results for improving cardiopulmonary bypass procedures [21].
Early visionary researchers, such as Robert Eberhart, clearly demonstrated the importance of identifying the impact of biomedical devices on activation of leukocytes and the consequences downstream on tissues and organs [22]. Recent studies have further added to the ability to account for activated molecular and cellular species and their compartmentalization [23]. Almost a 400% increase in leukocyte extravasation to the skin was seen in patients post-CPB. This level of potential inflammation is not measured by looking for activated cells in the blood or urine. Still further expansion of our knowledge concerning variation in clinical results, and thus a better prospective ability to control clinical outcomes with new technologies, comes from studies showing that the “cytokine storm” witnessed in many procedures may be the result of a biomedical device triggering a release of built-up stores of TNF-∝ and IL-8 in mucosal mast cells [24]. This build up of inflammatory cytokines can be the result of a pre-existing condition. An old study comes to mind where it was suggested that an indwelling catheter elevates the circulating levels of TNF-∝, and can interact with transient infections, exacerbating the response to the level of causing organ dysfunction [25]. Martin’s suggestion in 1991 that biomaterials can activate destructive cascades in the blood and tissue shows up again in the concluding questions of Gorbet and Sefton [20]:
“What happens over hours to days as leukocytes synthesize TF (tissue factor) and thrombotic deposits become remodeled is almost beyond current experimental capacity. Whether thrombosis leads to passivation or embolization or some other long-term consequence is still largely unknown.”
The phrase “beyond current experimental capacity” is echoed by Edmund’s statement regarding coatings [26]:
“Unfortunately, after a half century of effort, the complexity of biomaterial surface chemistry and physics, when exposed to blood, remains daunting; thus, no biomaterial or coating yet discovered reduces thrombin generation in extracorporeal perfusion systems. Our prospects are better by controlling the blood side of the interface.”
Interestingly, Gorbet and Sefton similarly concluded that the solution to blood and materials interactions associated with cardiovascular devices “may be achieved by preventing the pathway responsible for the inherent thrombogenicity of the materials.” Even more pessimistically, they state, “Rather than minimizing non-specific biomaterial-associated activation, active inhibition may be the only recourse” [20].
This last statement appears to suggest that coatings have minimal effect and they “only” generate “non-specific biomaterial-associated activation.” What happens on biomaterial surfaces is no more “daunting” than the other three impacts of CPB, namely; ischemia-reperfusion injury, endotoxemia, and operative trauma [27]. Modifying biomaterial surfaces to minimize production of activated species, if nothing else, will reduce the acute and chronic load for which the reticuloendothelial system needs to compensate. Inhibition of a pathway cannot solve all problems. There are simply too many examples of why we need to avoid completely blocking one pathway in a physiological system that depends on feedback mechanisms.
Rather than repeating the “tendency to focus on individual aspects of the whole problem,” it might be better to follow Ranucci’s [21] suggestion and take the multifactorial approach. From the clinical side of the ongoing discussion of biocompatible coatings, important studies and articles continue to provide data suggesting that heparin-coated circuits have benefit. Included among the number of consistent and dedicated contributors in this area are von Segesser [28], Lennartz [29], Ovrum [30], and Aldea [31]. They have carefully explored the potential of lowering systemic heparin levels with coated circuits with positive outcomes. Reducing systemic heparin greatly diminishes the bleeding risk and can be an important part of strategies to improve clinical protocols to include new inflammatory inhibitors.
It is well known that in the US the Centers for Medicare and Medicaid Services will stop reimbursement to hospitals for the cost of treating certain so-called hospital-acquired infections [32]. The consequences of a device-associated infection can be dramatic. For the patient it more often than not means an intervention to remove the infected device and a concomitant intense treatment with antibiotics to eradicate the infection [33]. Dependent on the clinical indication, at a later stage, when the infection has been eradicated, another intervention may yet be necessary to replace the medical device. Clearly, costs associated with a device-associated infection can be dramatic. The literature has reported that treatment of a device-infection can lead to substantial costs, often a multiple of the cost of the original device placement [34–36].
The pathophysiology of implant-associated infections is very complex. There are a number of host and microbial determinants that may contribute to the increased risk of implant-associated infections, and most probably the occurrence of an implant-related infection is the effect of interplay between them. Of note, The International Journal of Artificial Organs recently published a focus issue [37] on implant infections to which the interested reader is referred.
It has been unequivocally shown that implanted biomaterials strongly potentiate the susceptibility of tissues to infection [38, 39]. Most often bacterial adhesion to the implant surface is suggested to be predisposed to the pathogenesis of implant-related infections [40–42]. Foreign surfaces irreversibly attach microorganisms, after which a glycocalyx, an extracellular polysaccharide matrix synthesized by the microorganism, may be formed. The glycocalyx assists the bacteria in firmly adhering themselves to the inert surface. Cell division inside the glycocalyx results in the formation of microcolonies. Microcolonies coalesce to form biofilms as the size and number of adherent microcolonies increase. The adherent biofilm plays an important role in sustaining bacterial growth by trapping nutrients from the environment and resisting the effects of natural host defenses and antibiotics [43–45].
Obviously, protein adsorption has an influence on the bacterial adhesion process. Protein adsorption is the first event that takes place after contact of the implant material with the body(-fluids); subsequent phenomena are to a large extent determined by interactions of the body(-fluids) with this adsorbed protein layer [46]. The deposited protein layer may be pro- or anti-coagulant, pro- or anti-thrombotic, and pro- or anti-adhesive for bacteria [47].
Fibronectin was found to play a predominant role in promoting staphylococcal adhesion [48]. Besides fibronectin, however, there are a number of other host proteins, such as fibrinogen [49], laminin, and collagen [50], that have been suggested to be involved, independently or concomitantly, in the mediation of staphylococcal surface adhesion. The complexity of the entire interplay between adsorbed host proteins and bacteria-bearing binding sites for one or several of these proteins is obvious, but has not been fully unraveled yet.
In addition to offering bacteria a substrate to adhere to, implanted materials may also simply offer bacteria protective sites by means of their pores, crevices, interstices and/or lumen. Merritt et al. demonstrated that microorganisms can evade host defense mechanisms if they are able to enter the interstices of the implant [51].
In a discussion on implant-associated infections, apart from the role of the implant material, the host tissue response itself should also be included. Under normal circumstances, host defense mechanisms can cope with tissue injury by effecting repair and limiting the establishment of infection even when challenged by large numbers of bacteria [38]. The presence of a foreign body, however, provides a hindrance to appropriate healing, which may lead to an impaired antibacterial defense mechanism, thereby facilitating the establishment of infection [39, 52, 53]. Generally, an implant is not phagocytized because the implant is far greater than the size of phagocytic cell; because of this size disparity, “frustrated phagocytosis” may occur [54]. In an attempt to degrade the biomaterial, extracellular release of phagocytic products takes place, which may account for the decreased capacity to generate a respiratory burst, and consequently, a locally acquired phagocytic-bactericidal defect. This was elegantly demonstrated by Zaat et al. [55], who reported that the tissue surrounding an implant must be considered a niche for bacteria that cause biomaterial-associated infection. They found bacteria in the tissue surrounding the implanted material, actually in higher numbers than on the biomaterial itself. Bacteria were found in large numbers within macrophages, which suggests that these macrophages allowed intracellular survival and possibly replication of the bacteria. Tissue-residing Staphylococcus epidermidis bacteria resisted rifampin/vancomycin treatment, whereas bacteria were cleared from the implanted biomaterial itself [56].
Another rationalization for the implant-associated potentiation of infection is the generally observed dense fibrous encapsulation [47]. This fibrous tissue capsule is primarily composed of collagen, and forms a dense, poorly vascularized layer. The lack of vascularization presents a barrier to phagocytic cell migration to the implant site, thus providing for the development of relatively protective sites, where bacterial survival is favored.
It will not come as a surprise that the aforesaid change in reimbursement policy triggered an increased interest and demand for antimicrobial material technologies, with efforts directed at modification of the material surface to protect it from being colonized by bacteria. This “surface protection” approach can be roughly separated into the three distinct routes described below.
Two distinct routes for development of antimicrobial-releasing materials can be identified: inclusion of antimicrobial agents in or on the material. The resultant material eventually can be used as a vehicle for the prolonged delivery of the antimicrobial agent at the site. The ability of a sustained antimicrobial-releasing device to selectively deliver optimal amounts of the agent to the surrounding tissues offers an alternative, if not, an addition to conventional prophylactic antimicrobial therapy in minimizing postoperative complications of infection.
Catheters impregnated with chlorhexidine and silver sulfadiazine or with minocycline and rifampicin are the best studied as well as the most commercialized and frequently used antimicrobial-impregnated catheters [57, 58]. Silver or silver nanoparticles, due to their good antimicrobial action and low toxicity, have also been exploited [59]. Various other active release strategies for reducing the incidence of implant-associated infection have been reported on as well, on which Hetrick and Schoenfisch provide a good tutorial review [60].
If bacterial adhesion, mediated by deposited host proteins, is indeed the predominant step in the pathogenesis of implant-related infections, the inhibition of the adhesion to the implant surfaces would be an effective approach. Various surface modification approaches have been reported that significantly reduce the bacterial adhesion level [61–64]. However, little information is available on the susceptibility to bacterial adhesion of suchlike surfaces in an in-vivo environment. While in the in-vivo situation bacterial adhesion will be mediated by deposited proteins, contrarily, most studies were performed in vitro in the absence of proteins. Implant data proving that suchlike surfaces indeed may be beneficial in reduction of implant-related infections is sparsely available [65]. Considering that the surrounding tissue may provide a niche for biomaterial-associated infections [55, 56], the fact that a nonadhesive surface strategy does not provide for any (active) protection may limit the efficacy of this strategy after all. The interested reader is further referred to a recent review on antifouling coatings [66].
Gristina coined the term “race for the surface” [67] to describe the fate of biomaterial implants in relation to the development of biomaterials-associated infection: a race between microbial adhesion and biofilm growth on an implant surface vs. tissue integration. If the race is won by tissue cells, then the surface is covered by tissue and less vulnerable to bacterial colonization. On the other hand, if the race is won by bacteria the implant surface will rapidly become covered by a biofilm and tissue cell functions will be hampered by bacterial virulence factors and toxins. This concept of the race for the surface has been embraced by many researchers in the field. However, new biomaterials or functional coatings are evaluated either for their ability to resist bacterial adhesion and biofilm formation [68–71], or for their ability to support tissue cell adhesion and proliferation [68, 69, 72, 73]. Busscher and coworkers recently were the first to describe an in-vitro experimental methodology to investigate the race for the surface between bacteria and tissue cells in a single experiment [74].
Modification of surface properties in order to modulate the tissue response and to promote tissue integration can be classified into three categories: (a) physicochemical, (b) morphological, and (c) biochemical.
The physicochemical approach is based on the control of surface free energy, wettability, or surface charge [75]. Several studies have demonstrated that optimal cell adhesion occurs with surfaces showing a moderate wettability [76]. Additionally, cationic surface charge appears to accelerate cell adhesion, based on electrostatic interactions [77].
Cellular behavior is affected by the topographical morphology of the underlying surface. Surface texture in a micro-range was found to significantly modulate the cellular response. Experimental results showed that the histological response is affected more by the implant surface texture than by differing surface chemistry or surface energy [78, 79]. When in an optimal range, texture is apparently responsible for completion of the wound healing process and for the absence of chronic inflammatory cells at the interface. Apparently a properly microtextured surface either seems to offer cells the three-dimensional architecture they need for attachment and proliferation [80], or it seems to offer cells the possibility of creating one.
The biochemical approach comprises enhancing cell adhesion and proliferation by pre-coating of surfaces with cell adhesion proteins such as collagen [81], laminin [82], and fibronectin [83], and peptides such as RGD [84].
The tissue integration approach is very forceful: Okada and Ikada, for instance, reported on the application of a collagen coating on a percutaneous device that improved tissue adhesion and thereby prevented epidermal downgrowth; consequently, a significant reduction in device-related infections was observed [85]. An implanted material that promotes rapid colonization and integration of the appropriate tissue cells at its surface may be the best strategy for decreasing infectious complications. However, cell adhesive extracellular matrix proteins, for instance fibronectin and collagen, also have bacterial recognition sites. Therefore, during the initially vulnerable period before the surface is stabilized and when random colonization by bacteria may occur, active release strategies may be used protectively [67].
Stenting has undisputedly revolutionized the field of interventional cardiology. With that the biomaterials technology of stenting has improved rapidly in relation to both materials of construction of the device and surface modification of the stent. The initial focus for stent coatings was on heparinization of stents as an approach to stem high incidence of complications involving thrombotic occlusion of the stent and significant bleeding complications resulting from the use of intensive anticoagulation after implantation [86]. For example, the Carmeda BioActive Surface (CBAS) was coated onto Cordis (Johnson & Johnson) stents [86], and Medtronic introduced its HepaMed coating [87].
Clinically heparin-coated stents demonstrated a reduction of subacute thrombosis compared to bare metal stents, but through stent therapy optimization this soon was clinically managed pharmaceutically. However, the publication in 2001 of the first-in-man results showing zero restenosis after sirolimus-eluting stent implantation produced enormous excitement in the cardiological community. It was not too long thereafter that paclitaxel-eluting stents also demonstrated their capability to decrease restenosis. Today, sirolimus-, other ~limus- and paclitaxel-eluting stents have been shown in randomized trials to reduce restenosis as compared with conventional metallic stents (the reader is referred to Chapter 3, “Drug Delivery Coatings for Coronary Stents,” in this book for further details).
Biomaterials technology plays a key role in the design of a drug-eluting stent (DES) coating. In successfully designing a polymer-based drug delivery solution, many polymer performance requirements come together. Think of such design criteria as biocompatibility, drug elution rate, drug compatibility, film-forming properties, surface adhesion, and packaging requirements. While at first DES coatings utilized biostable polymers, more recently biodegradable polymers were introduced for the development of drug-eluting stent coatings as well [88]. Since the mid-2000s, the use of biodegradable DES coatings has gained interest with the emerging evidence of (very) late stent thrombosis with first generation DES. [89]. Another most interesting, “third generation biomaterials” approach (vide supra) to stent surface functionalization has aimed at accelerating stent endothelialization through coating of anti-CD34 antibodies on a stent surface to attract circulating endothelial progenitor cells to the stent, which then differentiate into endothelial-like cells [90]. A number of small-to-medium size registry and post-marketing studies confirmed the good safety profile of the EPC capture stent. However, large-scale randomized trials are lacking, so that its comparative effectiveness cannot be ascertained. Recent study of angiographic data showed in-stent late loss to compare unfavorably with that of drug-eluting stents [91]. Very recently, therefore, the Combo stent was designed [92]. This stent combines the EPC capture technology with abluminal elution of sirolimus. Its first-in-man results were presented at the Transcatheter Cardiovascular Therapeutics 23rd Annual Scientific Symposium, showing non-inferiority to the Taxus Liberté paclitaxel-eluting stent with respect to in-stent late lumen loss at 9-month angiographic follow-up.
As early as 1976 hydrogels grafted onto surfaces were considered to have excellent biocompatibility [93]. This led to several hydrophilic polymers such as poly(vinylpyrrolidone), poly(hydroxethylmethacrylate), poly(acrylamide), and poly(acrylic acid) being used to modify biomaterial surfaces. While these surfaces appeared to be clean and thrombus free, it became clear over time that they merely were non-adherent thrombus, and in fact generated thrombin that went downstream, creating emboli that were found in organs [94]. The next advance for biocompatible nonfouling or inert surfaces was polyethylene oxide (PEO). Early studies of Merill and Salzman showed low platelet adhesion to segmented polyurethanes rich in PEO surface molecules [95]. Numerous articles have been written on surfaces using this concept, which essentially creates a surface that is “repellant” to proteins. An excellent reference on protein and surface interactions for biocompatible polymer materials and coatings by Chen et al. [96] includes a comprehensive discussion and references for passivation of tethered surfaces, PEO protein resistance, biomimetic phosphorylcholine, zwitterionic surfaces by sulfobetaine or carboxybetaine polymers, heparin surfaces, and clot lysing surfaces. This well-balanced review is particularly valuable in pointing out potential pitfalls of the individual concepts. It also takes a strong position that whatever concept is attempted, having ligands attached to nonbiofouling surfaces will be the best approach for the future.
There are many hypotheses and strategies for modifying biomaterials, and with most methods careful process control and characterization are required. The best example of this need is for heparin coatings, where it has been demonstrated many times that not all heparin surfaces are the same [97, 98]. It is critically important to develop comprehensive in-vitro tests that look at coagulation and inflammatory response to materials, avoiding focus on single tests, if there is any hope of predicting clinical relevance [99]. For protein-resistant surfaces it is not sufficient to demonstrate minimal adhesion to the surface, as what is activated can go downstream to tissues and organs. With the new emphasis on extrinsic coagulation and the important role of complement, it is worth noting that in direct comparison some of these biopassive approaches are not as inert as hypothesized [100]. Perfection of surface modifications may be as elusive as the holy grail of inert surfaces. Nevertheless, the gain in understanding of blood and tissue interaction with materials has been impressive, especially in the last two decades. Most surface modifications improve the blood tissue response with respect to one or more categories of ISO 10993-4. The extent to which this can have a positive impact on clinical outcome remains the most challenging goal and need for the future. From Eberhart’s early demonstration of the importance of biomaterials impact on downstream tissues and organs [22] using labeled cells, to models that include broad molecular and cellular markers, a fairly good working model is developing that hopefully will make it more plausible to calibrate the impact that modification of the device will have in the total picture of clinical impactors [101].
The future of biomedical coatings lies directly with our increased understanding of the total impact from material surfaces to the glycocalyx surface. In the 1990s we had the decade of the brain; while not announced, the first decade of the third millennium could be considered the decade of the glycocalyx. Not only are we learning of the biochemical properties of this entity, but we continue to improve our understanding of cellular interactions and inflammatory sequelae associated with several diseases. Most recently, research has been revealing the ultimate missing link in tissue engineering and regenerative medicine, and that is the role of mechanotransduction in maintenance and repair of tissue [102].
A recent Dutch study best summarizes the current state-of-the-art by looking at the reactive oxygen (ROS) and reactive nitrogen (RNS) species damage to the glycocalyx and the far-reaching clinical implications. The majority of glycocalyx mechanisms in this study can be extrapolated to several disease states [103]. It is useful to recall the comments of Gorbet and Sefton [20] with respect to cellular activation by biomaterials:
“What happens over hours to days as leukocytes synthesize TF (Tissue Factor) and thrombotic deposits become remodeled is almost beyond current experimental capacity. Whether thrombosis leads to passivation or embolization or some other long-term consequence is still largely unknown.”
What is known and has been shown by Wagner et al.