Chapter 14 Cardiovascular Biomaterials

Source: STANDARD HANDBOOK OF BIOMEDICAL ENGINEERING AND DESIGN

CHAPTER 14 CARDIOVASCULAR BIOMATERIALS

Roger W. Snyder L. VAD Technology, Inc., Detroit, Michigan

Michael N. Helmus Boston Scientific Corporation, Boston, Massachusetts

14.1 INTRODUCTION 14.1 14.5 MATERIAL REPLACEMENT 14.10 14.2 MATERIALS 14.2 REFERENCES 14.11 14.3 TESTING 14.6 14.4 MATERIAL PROCESSING AND DEVICE DESIGN 14.10

14.1 INTRODUCTION

Numerous definitions for biomaterials have been proposed. One of the more inclusive is “any substance (other than a drug) or combination of substances synthetic or natural in origin, which can be used for any period of time, as a whole or part of a system which treats, augments, or replaces tissue, organ, or function of the body,” proposed by a Biomaterials Consensus Committee meeting at the National Institutes of Health (NIH).1 This definition must be extended because biomaterials are currently being utilized as scaffolds for tissue-engineered devices (hybrids of synthetic or biologic scaffolds and living cells and tissue for vessels, heart valves, and myocardium). Completely resorbable scaffolds can result in new organs without a trace of the original biomaterial. Efforts are also underway to use coatings that release bioactive agents to prevent thrombosis, infection, and hyper-plastic reactions (excessive tissue formation). The cardiovascular system consists of the heart and all the blood vessels. Cardiovascular biomaterials may contact blood (both arterial and venous), vascular endothelial cells, fibroblasts, and myocardium, as well as a number of other cells and acellular matrix material that make up all biological tissue. This chapter will consider a wide range of biomaterials that interact with the heart, blood, and blood vessels. Biomaterials used in the cardiovascular system are susceptible to a number of failure modes. Like all materials, mechanical failure is possible, particularly in implants. Although typical loads are low (as compared with orthopedic implants, for example), implant times are expected to exceed 10 years. At a typical heart rate of 90 beats per minute, 10 years of use would require more than 470 million cycles. Thrombosis is a unique failure mode for cardiovascular biomaterials. The resulting clots may occlude the device or may occlude small blood vessels, resulting in heart attacks, strokes, paralysis, failures of other organs, etc. On the other hand, devices can also damage blood cells. Hemolysis can

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14.2 BIOMATERIALS

diminish the oxygen-carrying capacity of the blood. Hemolysis can occur as a reaction to the material or its degradation products or as a result of shear due to the relative motion between the material surface and the blood. Cardiovascular biomaterials are also in contact with other tissues. Another common failure mode of these devices is excessive growth of the tissues surrounding the device. This can be caused by reaction to the material (the natural encapsulation reaction to any foreign body), stresses on surrounding tissues caused by the device, or reaction to material degradation products. Vascular grafts (in particular, smaller-diameter grafts) are subject to anastomotic hyperplasia, which reduces the diameter of the graft at the anastomosis. A similar hyperplastic response occurs around endovascular stents used to keep vessels open after angioplasty or as a de novo treatment. Heart valves can fail if tissue grows into the space occupied by the moving disk. Finally, tissue surrounding a device can die. As in hemolysis, this can be as a result of reaction with the material or its degradation products or as a result of continuous micromotion between the device and the tissue. Biomaterials that have been used in the cardiovascular system include processed biological substances, metals, and polymers (see Table 14.1 for typical materials and applications). Materials of biologic origin include structures such as pericardia, arteries and veins, and heart valves. Devices can also include biological substances, e.g., as coatings, such as collagen and heparin. Metals such as titanium, stainless steel, nitinol, cobalt-chrome alloys, etc., are used in many devices. Generally, these are metals with passive surfaces or surfaces that can be passivated. Silver has been used as a coating designed to resist infection. Glassy carbons have also been used as coatings to render surfaces thromboresistant. Pyrolytic carbon structures or coatings on graphite have been utilized in the fabrication of bileaflet heart valves. These are the most popular mechanical valves in use today. Polymeric materials that have been used in the cardiovascular system include polytetrafluorethylene, polyethylene terephthalate, polyurethane, polyvinyl chloride, etc. Textiles based on polytetrafluorethylene and polyethylene terephthalate are used extensively as fabrics for repair of vasculature and larger-vessel replacement (greater than 6 mm in diameter). Stent-grafts are hybrid stent grafts placed by catheter to treat aortic aneurysms nonsurgically and are fabricated of the same metallic alloys used in stents and textiles similar to those used in vascular grafts. Table 14.1 lists many of the biomaterials currently used in the cardiovascular system. Biomaterials are used throughout the cardiovascular system in both temporary and permanent devices. Cardiovascular devices can be divided into three categories: temporary external devices, temporary internal devices, and permanent internal devices. These categories are useful in determining the type of testing required. Temporary external devices range from simple tubing (for bypass or hemodialysis) to more complicated devices such as oxygenators, arterial filters, and hemodialysis equipment. For the purposes of this chapter, we will consider devices that contact blood only as external devices. Temporary internal devices include a wide range of catheters used for diagnostics and treatment. These also include guidewires and introducers for use with catheters and cannulae for use in bypass circuits. Vascular grafts and patches, as well as heart valves, are among the oldest of cardiovascular implants. More recently, permanent internal devices include pacemakers, defibrillators, stents, left ventricular assist devices, and artificial hearts.

14.2 MATERIALS

14.2.1 Metals

Metals are utilized for applications requiring high strength and/or endurance, such as structural components of heart valves, endovascular stents, and stent-graft combinations. Commonly used alloys include austenitic stainless steels (SS), cobalt-chrome (Co-Cr) alloys including molybdenum- based alloys, tantalum (Ta), and titanium (Ti) and its alloys. Elgiloy, a cobalt-nickel-chrome-iron

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TABLE 14.1 Biomaterials Used in the Cardiovascular System

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14.4 BIOMATERIALS

TABLE 14.1 Biomaterials Used in the Cardiovascular System (Continued)

alloy, has been used in fine-wire devices such as self-expanding endovascular stents. The shape memory or superelastic properties of nickel-titanium alloys are used in stents. Noble metals such as platinum-iridium are also utilized in pacemaker electrodes. In addition to the noble metals, stainless steel and tantalum can also be used in sensing (nonpacing) electrodes. Stainless steel has also been used as wire braids and reinforcements in catheters, particularly in high-pressure catheters, such as those used for radiopaque dye injection. Fatigue life is critical in many of the applications for which metallic alloys are used. Therefore, processing and joining methods must be such that crack initiation and propagation are mitigated.

14.2.2 Carbons and Ceramics

Carbons and glassy carbons have been widely used as heart valve components, particularly as pyrolytic carbon in the leaflets and housings of mechanical valves.10 They demonstrate good biocompatibility and thromboresistance, as well as high lubricity and resistance to wear, in this application. Graphite is used as the substrate for many of the pyrolytic carbon coatings. Strength and durability are imparted by the pyrolytic coatings. The use of a graphite substrate reduces residual stresses that become significant in thick pyrolytic coatings. The substrate has the potential to act as a barrier to crack propagation within the pyrolytic coating. Low-temperature isotropic coatings (LTI) can be used to coat more heat-sensitive polymeric substrates. Sapphires have also been utilized as bearings in high-rpm implantable rotary blood pumps. Ceramics have had limited application in cardiovascular devices except for hermetic seals on pacemakers and for insulation in radioablation catheters. Potentially, bioactive ceramics and glasses could have uses for enhanced cell and tissue adhesion. Experimental heart valves have been fabricated from ceramics, such as single-crystal sapphire leaflets for heart valves. Ceramic coatings of heart-valve components to improve their wear properties, particularly by chemical vapor deposition methods, e.g., diamondlike coatings, are another potential application.

14.2.3 Polymers

In the late 1800s, autologous venous grafts and homologous grafts were used to close arterial defects.4 However, the supply of these materials was limited. Long-term results were not promising, with many of the grafts developing aneurysms. In the early 1900s, solid-wall tubes of glass, methyl methacrylate, and various metals were tried. These were largely unsuccessful due to thrombosis and anastomotic aneurysms.

CARDIOVASCULAR BIOMATERIALS 14.5

During World War II and the Korean War, great progress was made in vascular surgery. Based on observations of sutures placed in the aorta, a textile was shown to have the ability to retain a fibrin layer, which then organized into a fibrous tissue layer. A number of materials were tested. Selection of these materials was based on two criteria: (1) minimal tissue reactivity and (2) availability in a textile form. Materials such as polyvinyl alcohol, polyamide, polyacrylonitrile, polyethylene terephthalate, and polytetrafluorethylene were all tried. As long as these textile tubes were implanted in the aorta, there was little clinical difference among the materials. Most of the differences in results were due to the different textile structures used. However, biostability and availability of commercial yarns did depend on the polymer chosen. Polyvinyl alcohol was soon abandoned due to excessive ruptures. Polyamide and polyacrylonitrile were discovered to be biodegradable, although it took 12 to 24 months to occur. Thus polyethylene terephthalate (polyester) and polytetrafluorethylene (PTFE) became the polymers of choice. Both these materials have demonstrated their longevity as an implant.5 Clinical results are excellent when the devices are implanted in high-flow, large-diameter arteries such as the aorta. However, patency rates decrease significantly when these devices are implanted below the aortic bifurcation. Thus other materials and structures have been investigated for low-flow, small-diameter arteries. In the mid-1970s, PTFE in a different form was introduced. Expanded PTFE is formed by compressing PTFE with a carrier medium and extruding the mixture. This is termed a paste extrusion because PTFE is not a thermomelt polymer. The resulting extrudate is then heated to near the glass transition temperature and stretched. Finally, the stretched material is sintered at a higher temperature. The resulting structure is microscopically porous with transverse plates of PTFE joined by thin PTFE fibers. This form of PTFE was indicated for use in smaller arteries with lower flow rates. However, the patency results obtained with this vascular graft were not significantly higher than those obtained with a polyester textile. The PTFE textile graft is no longer commercially available. In general, the handling characteristics of that device were not as good as those of the polyester textile because commercially available PTFE fibers were larger in diameter than polyester fibers. However, polyester and PTFE textiles, as well as expanded PTFE, are available as flat sheets. The textile materials are available as knits, weaves, and felts. These materials are used for patches and suture buttresses. Silicone is a rubberlike polymer. It is normally cross-linked in a mold or during extrusion. The most common silicone used is room-temperature-vulcanizing (RTV) silicone. In general, tissue does not adhere to silicone. The first commercially viable heart valve used a silicone ball in a cage. However, when first used, these silicone balls absorbed lipids and swelled, causing premature failure of the valves. These problems were corrected, and a small number of these valves are still implanted today. It is not uncommon that explants are recovered 30 years after implantation showing no degradation, minor wear, and only discoloration of the silicone ball. In current cardiovascular devices, silicone may be used as percutaneous drive lines. Synthetic bioresorbable materials have had a wide application as suture materials, although they have not generally been used in vascular anastomoses. They are being investigated for scaffolds for tissue-engineered heart valves and blood vessels. They are also being investigated as drug-release coatings on vascular prostheses and stents (to prevent thrombosis, infection, and excessive tissue formation) and as nanoparticles to deliver drugs to prevent restenosis.

14.2.4 Biological Materials

Materials of biological origin are used as cardiovascular devices and as coatings. Most of the devices commercially available rely on collagen as the structural material. Collagen is a macromolecule that exists as a triple helical structure of several amino acids. Procollagen is expressed by cells. The ends of the procollagen molecule are enzymatically trimmed, allowing the trimmed helical strands to self

14.6 BIOMATERIALS

assemble into the collagen molecule. At least 19 different types of collagen have been identified,5,13 with types I and III predominating in cardiovascular structures. The collagen molecule can be cross-linked by a number of techniques to improve its structural integrity and biostability. An early example of this is the tanning of skin to make leather. The use of formaldehyde to preserve biological samples is another example. In 1969, porcine aortic valves were cross-linked with gluteraldehyde and used to replace human aortic valves. Gluteraldyhde cross-linking was shown to yield a more biostable structure than cross- linking with formaldehyde.5 These valves have been very successful in older patients and do not require the anticoagulation regimen needed for mechanical heart valves. Significant effort has been made to reduce the calcification of bioprosthetic heart valves, both porcine aortic valve prostheses and bovine pericardial valve prostheses. Calcification and degradation mechanisms limit the use of these devices in young patients and children. Reduction of calcification entails modification of the surface, e.g., binding amino oleic acid or treatments with surfactants to remove lipids and other agents that can be a nidus for calcification.7 Their use in patients under 60 years of age is increasing and will continue to increase with the development of new treatments to reduce calcification. Cross-linked bovine arteries and human umbilical veins have been used as vascular replacements. Cross-linked pericardium has been used as a patch material, primarily between the myocardium and the pericardial sac to prevent adhesions. Biological materials can also be used as coatings. Textile vascular grafts are porous and must be sealed (preclotted) prior to use. Research6 suggests that a four-step procedure, including a final coat with heparinized blood, can improve patency results. However, surgeons typically take nonheparinized blood and coat the graft in one or two applications. Precoated grafts are commercially available. Cross-linked collagen and gelatin (a soluble form of collagen), as well as cross-linked albumin, can be used to seal porous materials. The rate of degradation of the coating will depend on the material chosen, as well as the degree of cross-linking. Significant effort is now focusing on tissue-engineered vessels and heart valves. The history of this effort is found in the seeding or culturing of endothelium on synthetic vascular prostheses. The clinical outcomes did not justify continued development. However, new technology allows vascular tissue to be formed on scaffolds of either synthetic or resorbable materials. The awareness that endothelium alone was not suitable has led to the evolution of techniques to recreate the vascular tissue utilizing multiple cells types.8,9 This approach utilizes the cell types expected in final structure, e.g., endothelium, smooth muscle cells, and fibroblasts, or pluripotential cells such as stem cells.

14.3 TESTING

The testing program for any medical device can be divided into five phases: (1) biocompatibility, (2) short-term bench (or in vitro) tests, (3) long-term bench tests, (4) animal (or in vivo) studies, and (5) human clinical studies.11 For each of these five phases, the type of device and length of time it will be used must be considered in developing test protocols.

14.3.1 Biocompatibility

Biocompatibility testing must measure the effects of the material on blood and tissue, as well as the effects of the organism on the material. International standards exist for demonstrating biocompatibility. These standards prescribe a series of tests, the selection of which depends on the length of time that the device will be in contact with the body. In these standards, any use less than 30 days is considered short term. Whether or not the device is external and will only contact blood or will be internal and in contact with tissue and blood also dictates which tests are necessary. Biocompatibility will be affected by surface contamination. Surface contamination can occur as a result of processing. Process aids, cleaning agents, finger oils, etc., can all have an impact on compatibility. Residues can also result from sterilization. Residual sterilants or sterilant by-products

CARDIOVASCULAR BIOMATERIALS 14.7

(such as ethylene oxide), modification of the material surface from radiation, and toxic debris from microbes can have an impact on biocompatibility. Materials such as solvents, plasticizers, unreacted monomers, and low-molecular-weight polymers can diffuse from polymers. Certain cleaning or thermal processes can accelerate diffusion. Therefore, all samples for biocompatibility testing should be from completed devices that have seen the complete process, including sterilization. If there is a chance that contaminates could continue to diffuse from the material, testing samples after storage should be considered. Since the cost of doing some of these tests (including the cost of the samples) can be significant, screening tests can be performed on any new material (or process). These tests are subsets of the standard tests. Some materials manufacturers provide biocompatibility data of this type. Once the materials used in the device have been shown to be biocompatible, consideration must be given to the function of the device. For example, devices subjected to flowing blood must be tested to document the lack of damage to blood components. Devices that rely on tissue ingrowth must be tested for this feature. These types of tests could be part of animal testing, which will be discussed later. The Blue Book Memo2 issued by the U.S. Food and Drug Administration (FDA), tabulates the tests required to demonstrate biocompatibility. These tables are based on an International Standards Organization (ISO) Document ISO-10993.3 For implant devices contacting blood for more than 30 days, the following tests are required: cytotoxicity, sensitization, irritation or intracutaneous reactivity, acute system toxicity, subchronic toxicity, genotoxicity, implantation, and hemocompatibility. For devices in contact with blood for less than 24 hours, subchronic toxicity and genotoxicity are not required. The international standard also has a category for devices that are in use for between 24 hours and 30 days. This standard does not require the subchronic toxicity testing. The FDA may require this type of testing, however. The tests required for implanted cardiovascular devices that do not contact blood require the same type of testing program except for the hemocompatibility requirement and for the implantation tests for devices in use for less than 24 hours. The tests for external devices contacting blood are also the same as for implanted devices, although implantation tests are noted as “may be applicable.” For long-term devices, either external or implants, chronic toxicity and carcinogenicity testing may also be required. Many manufacturers can provide biocompatibility data either in their literature or as an FDA Master File. Often material manufacturers will advertise that a material meets class IV biocompatibility requirements. Class VI requirements are an old set of tests published in the U.S. Pharmacopeia that were developed for testing food packaging. They are similar to the cytotoxicity, acute toxicity, and subchronic toxicity tests. However, the data provided by a materials manufacturer are on samples that have not seen the processing and storage of the device. The data simply are an indication that the material can pass the initial set of biocompatibility tests if processed appropriately. There are a wide variety of tests in the literature addressing these various requirements. Protocols for many of these tests have been issued as ISO standards. The American Society for Testing and Materials (ASTM) has also developed protocols for demonstrating biocompatibility. Since these standard protocols are recognized by many regulatory agencies, their use will often aid in the device approval process. Collaborating with a laboratory that specializes in these types of tests and is familiar with the regulatory requirements will generally produce the best data to demonstrate biocompatibility.

14.3.2 Short-Term Bench Testing

Short-term bench (in vitro) testing includes material identification, surface characterization, mechanical properties, etc. Material identification tests characterize the bulk properties of the material. Tests chosen depend on the type of material. Chemical formula, molecular weight, percentage crystallinity, melting or softening point, and degree of cross-linking may all be important to characterize a polymer. Composition, grain size, and contamination levels may define metallic materials. Composi- tion, molecular weight, cross-linking, shrinkage temperature, and purity may define materials of a biological origin.

14.8 BIOMATERIALS

Surface properties will affect the reaction between the material and tissue. Material composition at the surface may differ from the bulk composition. Coatings, either deliberately applied or as contaminates, will modify the biological response. Extraction studies, perhaps performed as part of the development of the cleaning process, could identify any inadvertent contamination. The coating should have an identification test. The geometry of the surface, such as surface roughness, will also modify this response. The dimensions of the surface roughness can be measured microscopically. For smaller features, scanning electron microscopy or atomic force microscopy can be utilized to characterize the surface roughness. Mechanical properties of the material will determine if a device will suffer an early failure. Tensile strength and elastic modulus can be measured by a simple tensile test. If the device could be subjected to impact loading, an impact type test can be performed. Tear tests are important for materials in sheets such as fabrics and films. This is particularly true for tear-sensitive materials such as silicone. The ASTM has published protocols for mechanical tests and for operating test equipment.

14.3.3 Long-Term Bench Testing

For long-term devices, endurance (or fatigue) testing is required. In general, simple tensile or bending tests can be performed on the basic material. Frequently, however, the device itself is tested in some simulated loading condition. Such a test includes the effects of processing, sterilization, and shelf life on the material. It also allows the designer to calculate reliability. There are test and reliability standards for some cardiovascular devices. Vascular grafts, heart valves, stents, and left- ventricular assist devices, among others, have reliability requirements and recommendations for the types of tests that can be employed. However, since there is a wide variation in these types of devices, tests that fit the device must be developed. Materials can be tested in tension, compression, or bending. Using the material in a sheet form, biaxial loading can be applied. For larger stress or strain ranges (and thus a lower number of cycles), the same equipment used to test for material strength can be used. However, for smaller loads and higher numbers of cycles, specialized equipment is required to complete the tests in a reasonable time. Loading devices using rotating cam shafts will apply fixed strain ranges. Fixed stress ranges can be applied using pressure-actuated devices. For very small stress or strain loads approaching the endurance limit, electronic mechanisms similar to those used to drive audio speakers can be used to drive a material at very high speeds. If one plots a variable such as stress range or strain range versus number of cycles, the resulting curve will approach a limit known as the endurance limit. Below this limit, the number of cycles that a material can withstand is theoretically infinite. Above this limit, the number of cycles that a material can withstand under a variable load can be calculated from Miner’s rule:

where ni is the number of cycles for a given load and Ni is the total number of cycles to failure under that load. Thus, if the stresses on a device can be calculated, the fatigue life of a device can be estimated. However, regulatory agencies prefer that the life of a device, or its reliability, be measured rather than calculated. Therefore, it is common practice to perform reliability testing on the device as it will be used in a patient. Although it is usually not possible to use blood or other biological fluids in a long-term test setup due to the difficulty in preserving the fluid, if the environment will affect the material, then a reasonable substitute must be found. In general, saline has been accepted as a substitute test media. If necessary, the saline can be buffered to an appropriate pH. Since cardiovascular devices, particularly implants, are expected to function for multiple years, tests to demonstrate reliability must be accelerated. At the same time, the test setup should apply loads that are as close to the actual usage as possible. Although some forms of degradation can be

CARDIOVASCULAR BIOMATERIALS 14.9

accelerated by an increase in temperature, it is common practice to test materials and devices at normal body temperature. Thus reliability tests are normally accelerated by increasing the cyclic rate. The upper limit of this type of acceleration is determined by several factors. First, the normal range of load and motion must be duplicated. For larger device parts, such as heart-valve disks, inertia will limit the cyclic rate. Second, any increase in temperature due to accelerating bending must not change the material. Internal temperatures of materials with glass transition temperatures above body temperature must remain below these transition points. Finally, the rate of loading must not affect, either negatively or positively, the amount of creep (or viscoelastic deformation) experienced by the material in actual use. Reliability can be calculated by several methods. One method used in other industries (such as the automotive industry) is to test 12 samples to the proposed life of the device. If all 12 samples survive, then the sample size is adequate to demonstrate a reasonable risk of failure using a binomial model. To determine failure modes, the stress or strain on the device can be increased by 10 percent for 10 percent of the proposed life of the device. If there are no failures at 110 percent of the proposed life, the stress or strain range is increased another 10 percent. This stair-step method continues until a failure mode is demonstrated. A more common method for medical devices is to run the life test until failure occurs. Then an exponential model can be used to calculate the percentage survivability. Using a chi-square distribution, limits of confidence on this calculation can be established. These calculations assume that a failure is equally likely to occur at any time. If this assumption is unreasonable (e.g., if there are a number of early failures), it may be necessary to use a Weibull model to calculate the mean time to failure. This statistical model requires the determination of two parameters and is much more difficult to apply to a test that some devices survived. In the heart-valve industry, lifetime prediction based on S-N (stress versus number of cycles) or damage-tolerant approaches is required. These methods require fatigue testing and ability to predict crack growth.10,11,14 Another long-term bench test required for these devices is shelf life. Some cardiovascular biomaterials degrade on the shelf. Thus typical devices, having seen the standard process, are packaged and aged before testing. Generally, aging can be accelerated by increasing the temperature of the storage conditions. As a general rule, it is accepted that the rate of degradation doubles for every 8°C increase in temperature. Some test labs also include variations in humidity in the protocol and may include a short period of low-temperature storage. Products that have been packaged and aged can then be tested to determine if they still meet the performance criteria. At the same time, these products can be tested for sterility, thus demonstrating that the packaging material and packaging process also yield the appropriate shelf life. Polymeric and biologic-based devices may also need to be evaluated on the basis of the biostability of the materials. This could include hydrolytic and enzymatic stability, requiring a combination of testing that examines hydrolytic stability under simulated physiologic stresses as well as evaluation in animals. Stress tends to accelerate many of these degradative mechanisms, and materials that look stable under static conditions may not perform well when stressed. Soft grades of polyether polyurethane are an example of a material than can undergo oxidative degradation when stressed due to the presence of oxidative enzymes present in biologic systems.

14.3.4 Animal Studies

There are few standard protocols for animal studies. Each study is typically designed to take into account the function and dimensions of the device. There are two approaches to developing a protocol. First, one could use a model of the condition being treated to demonstrate the function of the device. For example, vascular grafts can be implanted as replacements for arterial segments. The second approach is to design a test that will demonstrate the functioning of the device but not treat an abnormal condition. For example, a left ventricular assist device can be implanted in normal animals. The protocol would then consist of operating the device and monitoring the effect of the device on the blood. In addition, the effect of the biological environment on the device could be documented.

14.10 BIOMATERIALS

The first step in developing an appropriate protocol is to determine the purpose of the test. Is the purpose to demonstrate the functionality of the device, to demonstrate that the device is effective in treating a medical condition, or to test long-term biocompatibility and biostability of the device? The purpose of the test and the proposed use of the device (i.e., biological environment and length of use) will determine the model and length of time required for the test. If the device can be miniaturized, or if nonfunctioning devices are appropriate for the purpose of the test, then smaller animals can be used. Of course, the life span of the animal must be adequate for the implant time required. If, however, the purpose of the test requires a full-sized, functioning device, a larger animal model will have to be selected.

14.4 MATERIAL PROCESSING AND DEVICE DESIGN

Processing methods can have a major impact on the success or failure of a cardiovascular biomaterial. As described previously, surface features (either deliberately introduced or as the result of machining or tool imperfections), residues (from cleaning, handling, or sterilization), or process aids (either as surface residues or as bulk material diffusing from the biomaterial) can change the biological results. There are three methods currently used to decrease thrombosis: (1) use of a smooth surface to prevent thrombi from adhering, (2) use of a rough surface (usually a fiberlike surface) to encourage the formation of a neointima, and (3) use of a coating to prevent coagulation or platelet adherence. All these methods have been used in cardiovascular medical devices with varying degrees of success. As a general rule, the slower the flow of blood, the more likely thrombi will form. Thus the design of a device should avoid areas of stasis or low flow. A suitable surface must be either smooth, avoiding any small features that might cause microeddies in the flow, or of sufficient roughness so as to allow the resulting coagulation products to securely anchor to the surface. The thickness of the resulting layer is limited by the need to provide nutrients to the underlying tissue. Without the formation of blood vessels, this is generally about 0.7 mm. Should the material itself cause the formation of a thicker layer or should parts of the underlying structure move or constrict the tissue, the tissue will die. If this occurs continuously, the tissue will calcify or the underlying biomaterial will remain unhealed. Cleanliness of a material will also affect the biologic outcome. All processing aids should be completely removed. This includes any surfactant used in the preliminary cleaning steps. Surfactants can cause cell lysis or pyrogenic reactions. Solvents can diffuse into plastics and diffuse out slowly after implantation, causing a local toxic reaction. Some plastics may retain low-molecular-weight polymer or even monomers from their formation. These can also be toxic to cells. These may also leach out slowly after implantation. Plasticizers (used to keep some polymers pliable) can also damage blood components. The oxidation by-products of some metals can also be locally toxic. Thus it is important to establish a processing method prior to final evaluation of a cardiovascular biomaterial for use in a medical device.

14.5 MATERIAL REPLACEMENT

A number of major suppliers of materials no longer sell their products for long-term (more than 30 days) implantation in humans. In the past, the market for such materials was small (compared with the commercial market) and the liability high. Even if the material manufacturer could demonstrate that it was not liable for a particular use of its material, the legal cost was still quite high compared with possible sales. Even after Congress passed the Biomaterials Access Assurance Act of 1998 that shielded manufacturers of these materials from liability, many manufacturers remained reluctant to allow their materials to be used for long-term implants. In addition, materials may be totally withdrawn from the market if other commercial uses decline significantly.

CARDIOVASCULAR BIOMATERIALS 14.11

Many implant device manufacturers were forced to find replacement materials for their devices. Sometimes it was possible to locate an alternate supplier. At the same time, a number of material suppliers to the medical device industry have been started. In particular, new suppliers of silicones and polyurethanes have been established. Following the withdrawal of silicone from the breast implant market, the FDA modified its procedures, not requiring that replacement materials be tested as new materials but allowing manufacturers to demonstrate equivalency to previously used materials. Equivalency can be established through a series of data. Replacement materials can be compared with the original material in terms of chemical (structure, molecular weight, thermal properties, etc.) and mechanical properties (strength, tensile modulus, endurance, etc.). Data from actual devices fabricated from the new material should be compared with those from the original device. Biocompatibility data may be available from the material manufacturer, although the agency will request that additional data be supplied if the process used to fabricate the test samples differs from the process used to fabricate the device.

REFERENCES

1. Williams, D. F., ed., “Definitions in biomaterials,” Progress in Biomedical Engineering 4:67 (1987). 2. Blue Book Memorandum G#95-1, U.S. Food and Drug Administration, 1995. See http://www.fda.gov/cdrh. 3. Biological Evaluation of Medical Devices, Part 1: Evaluation and Testing, International Standards Organization (ISO) Document Number 10993. Geneva: Iso, 1997. 4. Weslowski, S. A., Evaluation of Tissue and Prosthetic Vascular Grafts. Springfield, IL: Charles C. Thomas, 1963. 5. Guidoin, R. C., Snyder, R. W., Awad, J. A., and King, M. W., “Biostability of vascular prostheses,” in Cardiovascular Biomaterials, Hastings, G. W. (ed.). New York: Springer-Verlag, 1991. 6. Greisler, H. P., New Biologic and Synthetic Vascular Prostheses. Austin, Tex.: R. G. Landes Co., 1991. 7. Schoen, F. J., and Levy, R. J., “Founder’s Award, 25th Annual Meeting of the Society for Biomaterials, Perspectives. Provi- dence, RI, April 28-May 2, 1999: Tissue heart valves: Current challenges and future research perspectives,” Journal of Biomedical Materials Research 15(47, 4):439–465 (1999). 8. Helmus, M. N., “Introduction/general perspective,” Frontiers of Industrial Research, International Society for Applied Cardiovascular Biology, (Abstract), Cardiovascular Pathology 7(5):281 (1998). 9. Helmus, M. N., “From bioprosthetic tissue engineered constructs for heart valve replacement,” in First International Sym- posium, Tissue Engineering for Heart Valve Bioprostheses, Satellite Symposium of the World Symposium on Heart Valve Disease, Friday, June 11, 1999, Westminster, London (Abstract), pp. 35–36. 10. Ritchie, R. O., “Fatigue and fracture of pyrolytic carbon: A damage-tolerant approach to structural integrity and life prediction in ‘ceramic’ heart valve prostheses,” Journal of Heart Valve Disease 5:(Suppl. 1): S9-S31 (1996). 11. Helmus, M. N., ed., Biomaterials in the Design and Reliability of Medical Devices. Georgetown, Tex.: Landes Bioscience, 2001. 12. von Recum, A. F., ed., Handbook of Biomaterials Evaluation: Scientific, Technical, and Clinical Testing of Implant Mate- rials, 2d ed. Philadelphia: Taylor & Francis, 1999. 13. Prockop, D. J., and Kivirikko, K. I., “Collagens: Molecular biology, diseases, and potentials for therapy,” Annual Revue of Biochemistry 64:403–434 (1995). 14. Kafesjian, R., and Schmidt, P., “Life analysis and testing: Short course, evaluation and testing of cardiovascular devices,” Society for Biomaterials, Course Notebook, Cerritos, Calif., 1995.