Chapter I.2.6
Degradable and Resorbable Biomaterials
Introduction
Clinical circumstances often require the application of implants that must serve a temporary rather than a permanent purpose (Table I.2.6.1). In these situations degradable polymers are of interest because the implants fabricated from these materials do not need to be surgically removed. The surgical removal of an implant with a temporary purpose is undesirable, as the process creates another wound with the possibility of surgical complication and infection. Additionally, the use of degradable implants can sometimes circumvent problems related to the long-term safety of permanent implants, such as long-term immune rejection, chronic inflammation at the implant–tissue interface, and failure of the device. However, degradable implants are not without their own safety concerns, such as the toxicity of their degradation products, and the degradation-related, premature failure of the implant. Therefore, designing a degradable implant requires careful testing for potential toxicity of its degradation products and careful consideration of the implant’s mechanical integrity during the required service life of the implant. To facilitate a better understanding of the complex decisions that have to be made during the design of a degradable implant, this chapter covers basic definitions relating to the process of degradation and/or erosion, the most prominent types of synthetic, degradable polymers available today, a classification of degradable medical implants, and considerations specific for the design and use of degradable medical polymers.
TABLE I.2.6.1 Some “Short Term” Medical Applications of Degradable Polymeric Biomaterials
| Application | Comments |
| Sutures | The earliest successful application of synthetic, degradable polymers in human medicine |
| Drug delivery devices | One of the most widely investigated medical applications for degradable polymers. Commonly used in clinical devices for the delivery of chemotherapeutics for the treatment of cancer |
| Orthopedic fixation devices | Requires polymers of exceptionally high mechanical strength and stiffness |
| Adhesion prevention | Requires polymers that can form soft membranes or films |
| Temporary vascular grafts and stents made of degradable polymers | Only investigational devices are presently available. Blood compatibility and fouling, as well as device fragmentation upon degradation, are major concerns |
| Tissue engineering or guided tissue regeneration scaffold | Attempts to recreate or improve native tissue function using degradable scaffolds. Cells may be seeded with the scaffolds before implantation. Bioactives (e.g., growth factors, short selective peptide sequences, etc.) are sometimes included in the scaffolds to modulate cellular responses |
Definitions Relating to the Processes of Degradation Versus Biodegradation, and Erosion Versus Bioerosion
The term “degradation” refers to a chemical process resulting in the cleavage of covalent bonds. Hydrolysis is the most common chemical process by which polymers degrade, but degradation can also occur via oxidative, photodegradative, and enzymatic mechanisms. In contrast, the term “erosion” refers to physical changes in size, shape or mass of a device, which could be the consequence of degradation, dissolution, ablation or mechanical wear. Thus, it is important to realize that erosion can occur in the absence of degradation, and degradation can occur in the absence of erosion. For example, erosion without degradation occurs when a sugar cube is placed in water: the sugar cube loses its mass through dissolution, e.g., it erodes, but the sugar does not degrade as its chemical structure remains unchanged. Degradation without erosion is observed when an implant made of poly(lactic acid) (PLA) is first placed into the body of a patient: PLA clearly starts to degrade, as evidenced by the reduction of its molecular weight, but the initial degradation of the polymer is not associated with erosion, as evidenced by the lack of mass loss of the implant. Only when the degradation of PLA is well advanced do the first signs of implant erosion become evident.
It is important to distinguish between the many different terms that describe the conversion of a solid biomaterial into a solution of the biomaterial or its molecular fragments. The list of terms includes degradation, biodegradation, bioabsorption, bioresorption, erosion, bioerosion, surface erosion, and bulk erosion. These terms are often used in an inconsistent fashion in the biomaterials literature; the reader is encouraged to read this section carefully and also the Glossary of Key Terms at the end of the chapter.
All polymers undergo some degree of degradation. Therefore, the practice of designating certain polymers as “degradable” and others as “non-degradable” represents a certain amount of arbitrariness. For the purposes of this chapter, we will use the criteria as described by Göpferich that “degradable polymers” are those that degrade within the time scales of their expected service life or shortly thereafter (Göpferich, 1996). Conversely, “non-degradable” polymers have degradation times that are substantially longer than their service life.
Currently, at least four different terms (biodegradation, bioerosion, bioabsorption, and bioresorption) are used to indicate that a given material or device will eventually disappear after having been introduced into a living organism. However, within the literature, no consistent distinctions in the meaning of these four terms are evident. Likewise, the meaning of the prefix “bio” is not well established, often leading to the interchangeable use of the terms “degradation” and “biodegradation”, or “erosion” and “bioerosion.” Although efforts have been made to establish generally applicable and widely accepted definitions for all aspects of biomaterials research (Williams, 1987), there is still confusion even among experienced researchers in the field as to the correct terminology of various degradation processes.
In the context of this chapter, we follow the usage suggested by the Consensus Conference of the European Society for Biomaterials (Williams, 1987), and refer to “biodegradation” when we wish to emphasize that a biological agent (enzyme, cell or microorganism) is causing the chemical breakdown of the implanted device. After extensive discussion in the literature, it is now widely believed that the chemical degradation of the polymeric backbone of poly(lactic acid) (PLA) is predominantly controlled by simple hydrolysis and occurs independently of any biological agent (Li, 1999). Consequently, in these instances, the degradation of PLA to lactic acid should not be described as “biodegradation” but rather as “degradation.” (Only under rare conditions may the enzyme proteinase K catalyze the hydrolytic degradation of PLA (Williams, 1981; Tsuji and Miyauchi, 2001); that situation is so rare that we will not consider it further here.)
In agreement with Heller’s suggestion (Heller, 1987), we define a “bioerodible polymer” as a water-insoluble polymer that is converted under physiological conditions into water-soluble material(s) without regard to the specific mechanism involved in the erosion process. “Bioerosion” includes, therefore, both physical processes (such as dissolution) and chemical processes (such as backbone cleavage). Here the prefix “bio” indicates that the erosion occurs under physiological conditions, as opposed to other erosion processes caused, for example, by high temperature, strong acids or bases, UV light, mechanical stresses or weather conditions. The terms “bioresorption” and “bioabsorption” are used interchangeably and often imply that the polymer or its degradation products are removed by cellular activity (e.g., phagocytosis) in a biological environment. While commonly found within the literature, these terms have not been consistently defined.
An understanding of the distinct definitions of degradation and bioerosion are necessary to facilitate the subsequent discussion of degradable polymers and devices. The mechanisms and of degradation and bioerosion are discussed more thoroughly on pages 189–190: The Process of Bioerosion, Mechanisms of Chemical Degradation, and Factors that Influence the Rate of Bioerosion.
Overview of Currently Available Degradable Polymers
The development of highly stable materials has always been a major research challenge. Today, many polymers are available for clinical purposes that are virtually non-destructible in biological systems, e.g., Teflon™ (poly(tetrafluoroethylene)), Kevlar™ (poly(paraphenylene terephthalamide)), poly(ether-ether-ketone) (PEEK), or poly(ethylene terephthalate) (PET). On the other hand, the development of degradable biomaterials is a relatively new area of research. The variety of available, degradable biomaterials is still too limited to cover a wide enough range of diverse material properties. Thus, the design and synthesis of new, degradable biomaterials is currently an important research challenge. Within the context of tissue engineering, new biomaterials that provide predetermined and controlled cellular responses are a critically needed component of most practical applications (Anderson et al., 2004).
Degradable materials must fulfill more stringent requirements, in terms of their biocompatibility, than non-degradable materials. In addition to the potential problem of toxic contaminants leaching from the implant (residual monomers, stabilizers, polymerization initiators, emulsifiers, sterilization byproducts), one must also consider the potential toxicity of the degradation products and subsequent metabolites. The practical consequence of this consideration is that only a limited number of non-toxic, monomeric starting materials have been successfully applied to the preparation of degradable biomaterials.
Over the last decade a number of hydrolytically unstable polymers have been suggested as degradable biomaterials. However, in most cases, no attempts have been made to develop these new materials for specific medical applications. Thus, detailed toxicological studies in vivo, investigations of degradation rate and mechanism, and careful evaluations of the physico-mechanical properties have so far been published for only a small fraction of those polymers. An even smaller number of synthetic, degradable polymers have so far been used in medical implants and devices that have gained approval by the Food and Drug Administration (FDA) for use in patients. The FDA does not approve polymers or materials per se, but only specific medical devices and drug delivery formulations. As of 2006, only seven distinct synthetic, degradable polymer classes have been approved for use in a narrow range of clinical applications (Kohn et al., 2007). These polymers are polyesters containing lactic acid, polyesters containing glycolic acid, polyesters containing dioxanone, polyesters containing caprolactone, poly(trimethylene carbonate)s, polyanhydrides containing sebacic acid, and tyrosine-derived polyarylates. Various other synthetic, degradable biomaterials currently in clinical use are blends or copolymers of these base materials, such as a wide range of copolymers of lactic and glycolic acid. This listing does not include polymers derived from animal sources such as collagen, gelatin or hyaluronic acid.
It is important to note that no polymer will degrade or erode by either hydrolytic or enzymatic processes unless water molecules can reach and “interact with” the molecular groups of the polymer chains. Thus, the “degree of hydrophobicity,” the presence of crystallites, and the resultant absorption of water by the polymer are critical factors for all degradation/resorption processes.
Recent research has led to several well-established investigational polymers that may find practical applications as degradable implants within the next decade. This chapter will concern itself mostly with synthetic degradable polymers, since natural polymers (e.g., polymers derived from animal or plant sources) are described elsewhere in this book. Furthermore, the chapter on Hydrogels (Chapter I.2.5) contains a section on degradable hydrogels, and therefore, that subject won’t be covered in this chapter.
Table I.2.6.2 provides an overview of some representative degradable polymers. For completeness, some of the natural polymers have also been included here. Structural formulas of commonly investigated synthetic degradable polymers are provided in Figure I.2.6.1. A large proportion of the currently investigated, synthetic, degradable polymers are polyesters. It remains to be seen whether some of the alternative backbone structures such as polyanhydrides, polyphosphazenes, polyphosphonates, polyphosphoesters, polyamides, or polycarbonates will be able to challenge the predominant position of the polyesters in the future.
TABLE I.2.6.2 Degradable Polymers and Applications Under Investigation
FIGURE I.2.6.1 Chemical structures of widely investigated degradable polymers.
Poly(glycolic acid) and poly(lactic acid) and their copolymers are currently the most widely investigated, and most commonly used, synthetic, bioerodible polymers. In view of their importance in the field of biomaterials, their properties and applications will be described in more detail.
Poly(glycolic acid) (PGA) is the simplest linear, aliphatic polyester (Figure I.2.6.1). Since PGA is highly crystalline, it has a high melting point and low solubility in organic solvents. PGA was used in the development of the first totally synthetic, absorbable suture (Herrmann et al., 1970). PGA sutures have been commercially available under the trade name “Dexon™” since 1970. A practical limitation of Dexon™ sutures is that they tend to lose their mechanical strength rapidly, typically over a period of two to four weeks after implantation. PGA has also been used in the design of internal bone fixation devices (bone pins). These pins have become commercially available under the trade name “Biofix™.”
To adapt the material properties of PGA to a wider range of possible applications, copolymers of glycolic acid with the more hydrophobic lactic acid were intensively investigated (Gilding and Reed, 1979). The hydrophobicity of lactic acid limits the water uptake of poly(lactic acid) (PLA) thin films to about 2%, and reduces the rate of backbone hydrolysis as compared to PGA. Sutures composed of copolymers of lactic acid and glycolic acid, poly(lactide-co-glycolide) (PLGA), were developed as an alternative to PGA sutures. The suture and the 90:10 PLGA copolymer it is composed of are distributed under the trade names “Vicryl™” and “Polyglactin 910™.”
It is noteworthy that there is no linear relationship between the ratio of glycolic acid to lactic acid and the physico-mechanical properties of the corresponding copolymers. Whereas PGA is highly crystalline, crystallinity is rapidly lost in copolymers of glycolic acid and lactic acid. These morphological changes lead to an increase in the rates of hydration and hydrolysis. Thus, 50:50 PLGA copolymers degrade more rapidly than either PGA or PLA.
Since lactic acid is a chiral molecule, it exists in two stereoisomeric forms which give rise to four morphologically distinct polymers: the two-stereoregular polymers, D-PLA and L-PLA; and the racemic form D,L-PLA. A fourth morphological form, meso-PLA, can be obtained from D,L lactide but is rarely used in practice.
The polymers derived from the optically active D and L monomers are semicrystalline materials, while the optically inactive D,L-PLA is always amorphous. Generally, L-PLA is more frequently employed than D-PLA, since the hydrolysis of L-PLA yields L(+) lactic acid, which is the naturally occurring stereoisomer of lactic acid.
The differences in the crystallinity of D,L-PLA and L-PLA have important practical ramifications. Since D,L-PLA is an amorphous polymer, it is usually considered for applications such as drug delivery, where it is important to have a homogeneous dispersion of the active species within the carrier matrix. On the other hand, the semicrystalline L-PLA is preferred in applications where high mechanical strength and toughness are required, such as sutures and orthopedic devices.
PLA and PGA and their co-polymers have been investigated for more applications than any other degradable polymer. The high interest in these materials is not based on their superior material’s properties, but mostly on the fact that these polymers have already been used successfully in a number of approved medical implants and are considered safe, non-toxic, and biocompatible by regulatory agencies in virtually all developed countries. Therefore, implantable devices prepared from PLA, PGA or their copolymers can be brought to market in less time and for a lower cost than similar devices prepared from novel polymers whose biocompatibility is still unproven.
Currently available and approved products include: sutures; suture reinforcements; GTR membranes for dentistry; orthopedic fixation devices; injectable fillers for restoration of facial volume; skin replacement materials; dura mater substitutes; hormone delivery systems; and implantable drug delivery systems. The polymers are also being widely investigated in the design of vascular and urological stents, and as scaffolds for tissue engineering and tissue reconstruction. In many of these applications, PLA, PGA and their copolymers have performed with moderate to high degrees of success. However, there are still unresolved issues: first, in tissue culture experiments, most cells do not attach to PLA or PGA surfaces and do not grow as vigorously as on the surface of other materials, indicating that these polymers are actually poor substrates for cell growth in vitro. Second, the degradation products of PLA and PGA are relatively strong acids (lactic acid and glycolic acid). When these degradation products accumulate at the implant site, a delayed inflammatory response is often observed months to years after implantation (Bergsma et al., 1995; Ignatius and Claes, 1996; Törmälä et al., 1998).
Poly(ethylene glycol) (PEG) has been used to initate lactide and/or glycolide ring opening, producing A–B–A triblock copolymers, which are degradable, producing LA, GA, and PEG as byproducts. They have been made into a variety of PEG-based hydrogels and used in drug delivery, and as tissue engineering scaffolds. They are discussed in detail in the chapter on Hydrogels (Chapter I.2.5).
Polydioxanone (PDS) is a poly(ether-ester) made by a ring opening polymerization of p-dioxanone monomer. PDS was first introduced in the early 1980s and has gained increasing interest in the medical and pharmaceutical fields due to the low toxicity of its degradation products in vivo. The degradation time of PDS in vivo is approximately 60 days, with complete mass loss within 9 to 12 months (Maurus and Kaeding, 2004). This degradation time is comparable to PGA. PDS has a lower modulus and glass transition temperature than PLA or PGA. PDS became the first degradable polymer to be used to make a monofilament suture (Ray et al., 1981). PDS has also been introduced to the market as a suture clip, a degradable ligating device, as well as a bone pin marketed under the name OrthoSorb® in the USA and Ethipin® in Europe.
Poly(hydroxybutyrate) (PHB), poly(hydroxyvalerate) (PHV), and their copolymers represent examples of resorbable polyesters that are derived from microorganisms. PHB and PHV are intracellular storage polymers providing a reserve of carbon and energy to microorganisms similar to the role of starch in plant metabolism. Therefore, although these classes of polymers are examples of natural materials (as opposed to synthetic materials), they are included here because they have similar properties and similar areas of application as the widely investigated poly(lactic acid). The polymers can be degraded by soil bacteria (Senior et al., 1972), but are relatively stable under physiological conditions (pH 7.4, 37°C). PHB is currently available as either poly(3-hydroxybutyrate) (P3HB) or poly(4-hydroxybutyrate) (P4HB). P3HB and its copolymers with up to 30% of 3-hydroxyvaleric acid are commercially available under the trade name “Biopol®” (Miller and Williams, 1987). P4HB is likewise available for clinical applications under the trade name “PHA4400.” Within a relatively narrow window, the rate of degradation can be modified slightly by varying the copolymer composition between PHB and PHV. Members of the P3HB family of polymers require several years for complete resorption in vivo, while P4HB members require 8–52 weeks. In vivo, P3HB and P4HB degrade to D-3-hydroxybutyric acid and D-4-hydroxybutyric acid, respectively, which are natural metabolites present in the brain, heart, lung, liver, kidney, and muscle (Nelson et al., 1981). The low toxicity of PHBs may at least in part be due to this fact.
The mechanical properties of PHBs are dependent on the morphology considered. P3HB homopolymer is a highly crystalline and brittle material, while the copolymers of PHB with hydroxyvaleric acid are less crystalline, more flexible, and more readily processible (Barham et al., 1984). In contrast, P4HB is a strong, pliable thermoplastic material that is more flexible than PLA and PGA with an elongation break around 1000% (Martin and Williams, 2003). These polymers have been considered in several biomedical applications, such as: controlled drug release devices; sutures; artificial skin; surgical patching materials for congenital cardiovascular defects; heart valves; vascular grafts; and also in industrial applications such as medical disposables (Martin and Williams, 2003). P3HB is especially attractive for orthopedic applications due its slow degradation time. The polymer typically retains 80% of its original stiffness over 500 days upon in vivo degradation (Knowles, 1993).
Polycaprolactone (PCL) was first synthesized via ring-opening polymerization in the 1930s (Van Natta et al., 1934). However, its degradability was identified in the 1970s and it became available commercially following efforts at Union Carbide to identify synthetic polymers that could be degraded by microorganisms (Huang, 1985). Based on a large number of tests, ε–caprolactone and polycaprolactone are currently regarded as non-toxic and tissue-compatible materials. It is a semicrystalline polymer having a low glass transition temperature. The high solubility of polycaprolactone, its low melting point (59–64°C), and exceptional ability to form blends has stimulated research on its application as a biomaterial. Polycaprolactone degrades at a slower rate than PLA and has a high permeability to many drugs, therefore prompting its use in drug delivery devices that remain active for over one year (Sinha et al., 2004). The release characteristics of polycaprolactone have been investigated in detail by Pitt and his co-workers (Pitt et al., 1979). The Capronor™ system, a one-year implantable contraceptive device, has been commercially available in Europe and the USA. The toxicology of polycaprolactone has been extensively studied as part of the evaluation of Capronor™. In Europe, polycaprolactone is already in clinical use as a degradable staple (for wound closure). A monofilament suture derived from copolymers of polycaprolactone and PGA is widely used clinically and is available under the brand name Monocryl™ (Bezwada et al., 1995). Caprolactones have also been used in block copolymers with glycolide, lactide, and poly(ethylene glycol) for a commercially available drug delivery device under the brand name SynBiosys™.
Polyanhydrides were explored as possible substitutes for polyesters in textile applications, but failed due to their pronounced hydrolytic instability. It was this property that prompted Langer and his co-workers to explore polyanhydrides as degradable implant materials (Tamada and Langer, 1993). The main advantages of polyanhydrides are that: their production requires a single-step synthesis scheme with inexpensive reagents; they have controllable molecular weights and predictable degradation/drug release profiles; they are easily processable; and they are relatively non-toxic upon degradation (Kumar et al., 2002).
The polyanhydride library includes a large number of synthetic materials that all contain degradable anhydride bonds within the backbone of the polymer chain. To date, polyanhydrides have been synthesized with aliphatic monomers, unsaturated monomers, aromatic monomers, and linear fatty acid monomers (Nair and Laurencin, 2007). Aliphatic polyanhydrides degrade within days, whereas some aromatic polyanhydrides degrade over several years. Thus, aliphatic–aromatic copolymers are usually employed to achieve intermediate rates of degradation which are dependant on the monomer composition (Göpferich and Tessmar, 2002).
Polyanhydrides are among the most reactive and hydrolytically unstable polymers currently used as biomaterials. The high chemical reactivity is both an advantage and a limitation of polyanhydrides. Many polyanhydrides degrade by surface erosion without the need to incorporate various catalysts or excipients into the device formulation. While this property makes them good candidates for drug delivery applications, their hydrolytic instability limits shelf life. To maintain their integrity, polyanhydride materials must be stored under anhydrous and low temperature conditions. Additionally, polyanhydrides can react with drugs containing free amino groups or other nucleophilic functional groups, especially during high temperature processing (Leong et al., 1986). The potential reactivity of the polymer matrix toward nucleophiles limits the type of drugs that can be successfully incorporated into a polyanhydride matrix by melt processing techniques.
A comprehensive evaluation of the toxicity of polyanhydrides, as reviewed by Katti et al. demonstrates that, in general, the polyanhydrides possess excellent in vivo biocompatibility (Katti et al., 2002). The most immediate applications for polyanhydrides have been in the field of drug delivery. Drug loaded devices made of polyanhydrides can be prepared by compression, injection or melt molding, solvent casting or microencapsulation (Kumar et al., 2002). Depending on the monomer composition, polyanhydrides are able to release drugs with zero order kinetics over time periods ranging from days to years (Nair and Laurencin, 2007). A wide variety of drugs and proteins including insulin, bovine growth factors, angiogenesis inhibitors (e.g., heparin and cortisone), enzymes (e.g., alkaline phosphatase and β-galactosidase), and anesthetics have been incorporated into polyanhydride matrices with their in vitro and in vivo release characteristics evaluated (Katti et al., 2002; Kumar et al., 2002). The first polyanhydride-based drug delivery system to enter clinical use was for the delivery of chemotherapeutic agents. An example of this application is the delivery of BCNU (bis-chloroethylnitrosourea) to the brain for the treatment of glioblastoma multiforme, a highly fatal brain cancer (Brem et al., 1993). For this application, BCNU-loaded implants made of the polyanhydride derived from bis-p-carboxyphenoxypropane and sebacic acid received FDA regulatory approval in the fall of 1996 and are currently being marketed under the name Gliadel™. Septacin™, a drug delivery vehicle for gentamicin sulfate has been developed for the treatment of osteomyelitis (Li et al., 2002).
Poly(ortho ester)s (POE) are a family of synthetic, degradable polymers that have been under development for a number of years (Heller and Gurny, 1999). Currently, there are four major types of poly(ortho ester)s: first generation (POE I); second generation (POE II); third generation (POE III); and fourth generation (POE IV) (Heller et al., 2002). (See also the chapter on Drug Delivery, Chapter II.5.16.)
POE I were first prepared by Choi and Heller via the transesterification of 2,2′–dimethoxyfuran with a diol (Heller and Gurny, 1999). Unfortunately, the hydrolysis of POE I is an acid-sensitive process that itself results in the release of acidic degradation products. This leads to autocatalysis of the polymer, limiting its utility for clinical applications. The second generation of poly(ortho ester)s (POE II) were based on an acid catalyzed addition reaction of diols with diketeneacetals (Heller, 1990). Unlike POE I, the degradation products of this polymer are neutral in pH, therefore preventing the autocatalysis of the polymer. The mechanical properties of the POE II can be controlled to a large extent by the choice of the diols used in the synthesis. Materials ranging from stiff materials to those that are quite soft at room temperature can be achieved. Since the ortho ester linkage in POE I and II is far more stable in base than in acid, Heller and his co-workers controlled the rate of polymer degradation by incorporating acidic or basic excipients into the polymer matrix. The third generation of poly(ortho ester)s (POE III) were very soft and were even viscous liquids at room temperature (Merkli et al., 1993). POE III can be used in the formulation of drug delivery systems that are injected rather than implanted into the body. Release profiles of the chemotherapeutic agent 5-fluorouracil demonstrated zero order kinetics and were investigated for ocular implants (Heller, 2005). Unfortunately, difficulties in synthesis and reproducibility limited the practical applications of this material, and currently it is no longer under development. The final generation of poly(ortho ester)s (POE IV) was synthesized by Heller and his co-workers to address the need for excipients in POE I and II (Ng et al., 1997). This polymer achieves consistent degradation profiles via the incorporation of glycolic or lactic acid to the polymer backbone. Upon degradation, the acid segments are released, therefore catalyzing the hydrolysis of the ortho ester linkages.
POE IV is naturally “surface eroding”, while POE I and II can erode by “surface erosion” if appropriate excipients are incorporated into the polymeric matrix. One concern about the “surface erodability” of poly(ortho ester)s is that the incorporation of highly water-soluble drugs into the polymeric matrix can result in swelling of the polymer matrix. The increased amount of water imbibed into the matrix can then cause the polymeric device to exhibit “bulk erosion” instead of “surface erosion” (see the section The Process of Bioerosion, pp. 189–190, for a more detailed explanation of these erosion mechanisms) (Okada and Toguchi, 1995). Since surface eroding, slab-like devices tend to release drugs embedded within the polymer at a constant rate, poly(ortho ester)s appear to be particularly suited for controlled release drug delivery applications. For example, poly(ortho ester)s have been studied for the controlled delivery of cyclobenzaprine and steroids, and a significant number of publications describe the use of poly(ortho ester)s for various drug delivery applications (Heller and Gurny, 1999) (see Chapter II.5.16).
Poly(amino acid)s and “pseudo”-poly(amino acid)s: since proteins are composed of amino acids, researchers explored the possible use of poly(amino acid)s in biomedical applications (Anderson et al., 1985). Poly(amino acid)s were regarded as promising candidates since the amino acid side chains offer sites for the attachment of drugs, cross-linking agents, or pendent groups that can be used to modify the physico-mechanical properties of the polymer. Additionally, they initially held potential as degradable polymer materials because they possessed long range order, defined conformations in solution, and their amino acid-derived building blocks lacked toxicity (Kohn, 1993). Poly(amino acid)s have been investigated for use as suture materials (Williams, 1982), as artificial skin substitutes (Spira et al., 1969), and as drug delivery systems (Matsumura, 2008). Various drugs have been attached to the side chains of poly(amino acid)s, usually via a spacer unit that distances the drug from the backbone. Poly(amino acid)–drug combinations investigated for chemotherapy include poly(L-lysine) with methotrexate and pepstatin (Campbell et al., 1980), poly(glutamic acid) with adriamycin and cisplatin (Van Heeswijk et al., 1985), and PEG-polyaspartate with paclitaxel (Matsumura, 2008).
Despite their apparent potential as biomaterials, poly(amino acid)s have actually found few practical applications due to several synthetic disadvantages: the expense of production; insolubility in common organic solvents; and thermal degradation upon melting limiting processibility. Poly(amino acid)s have a pronounced tendency to swell in aqueous media, therefore making the prediction of drug release rates difficult. Furthermore, the antigenicity of polymers containing three or more amino acids limits their use in biomedical applications (Kohn, 1993). Due to these difficulties, only a few poly(amino acid)s have been investigated for clinical uses. While some poly(amino acid) derivatives of poly(glutamic acid) carrying various pendent chains at the γ-carboxylic acid group have been investigated as implant materials (Lescure et al., 1989), the majority of their clinical success has been in the drug delivery of chemotherapeutics. As of 2008, PEG-poly(glutamic acid) nanocarrier formulations are in Phase 1 clinical trials for the delivery of paclitaxel and camptothecin, while Phase 2 trials are underway for cisplatin delivery against stomach cancer (Matsumura, 2008). So far, no implantable devices made of a poly(amino acid) have been approved for clinical use in the USA.
In an attempt to circumvent the problems associated with conventional poly(amino acid)s, backbone-modified “pseudo”-poly(amino acid)s were introduced in 1984 (Kohn and Langer, 1984, 1987). The first “pseudo”-poly(amino acid)s investigated were a polyester from N-protected trans-4-hydroxy-L-proline, and a polyiminocarbonate derived from tyrosine dipeptide. The tyrosine-derived “pseudo”-poly(amino acid)s are processable by solvent or heat methods and exhibit a high degree of biocompatibility.
The reason for the improved physico-mechanical properties of “pseudo”-poly(amino acid)s relative to conventional poly(amino acid)s can be traced to the reduction in the number of interchain hydrogen bonds. In conventional poly(amino acid)s, individual amino acids are polymerized via repeated amide bonds, leading to strong interchain hydrogen bonding. In natural peptides, hydrogen bonding is one of the interactions leading to the spontaneous formation of secondary structures such as α-helices or β-pleated sheets. Strong hydrogen bonding also results in high processing temperatures and low solubility in organic solvents, which tends to lead to intractable polymers with limited applications. In “pseudo”-poly(amino acid)s, half of the amide bonds are replaced by other linkages (such as, for example, carbonate, ester, or iminocarbonate bonds) which have a much lower tendency to form interchain hydrogen bonds, leading to better processibility and, generally, a loss of crystallinity.
For example, tyrosine-derived polycarbonates (Pulapura and Kohn, 1992; Bourke and Kohn, 2003) are high strength materials that may be useful in the formulation of degradable orthopedic implants. One of the tyrosine-derived pseudo-poly(amino acid)s, poly(DTE carbonate) exhibits a high degree of bone conductivity (e.g., bone tissue will grow directly along the polymeric implant) (Choueka et al., 1996; James et al., 1999; Asikainen et al., 2006). As of 2012, tyrosine derived “pseudo”-poly(amino acid)s have been used in two FDA approved devices (a hernia repair device and an antimicrobial pacemaker pouch from TyRx Pharma, Inc.) and a third is in clinical trials (a resorbable cardiovascular stent from REVA Medical, Inc.).
Polycyanoacrylates are used as bioadhesives and have been investigated for drug delivery. These adhesives are known for their high rates of curing, high adhesive strength, and good adhesion to both tissue and inorganic materials. Cyanoacrylates undergo spontaneous polymerization at room temperature in the presence of water. Their toxicity and erosion rate after polymerization differs with the length of their alkyl chain (Vezin and Florence, 1980). Methyl cyanoacrylates are more commonly used as general-purpose glues and are commercially available as “Crazy Glue.” Methyl cyanoacrylate was used during the Vietnam War as an emergency tissue adhesive, but is no longer used today. Butyl cyanoacrylate is a slowly degrading polymer approved for use in Canada and Europe as a dental adhesive. n-Butyl-cyanoacrylate is used in veterinary glues, while 2-octyl cyanoacrylate (Dermabond™) is used in liquid bandages. Polycyanoacrylates have several limiting properties: first, the monomers (cyanoacrylates) are very reactive compounds that often have significant toxicity; second, upon degradation polycyanoacrylates release formaldehyde resulting in intense inflammation in the surrounding tissue. In spite of these inherent limitations, polycyanoacrylates have been investigated as potential drug delivery matrices, have been suggested for use in ocular drug delivery (Deshpande et al., 1998) and the delivery of nucleic acids (Fattal and Couvreur, 2004), and are in late stage clinical trials for cancer therapy (Vauthier et al., 2003).
Polyphosphazenes are polymers whose backbone consists of nitrogen–phosphorous bonds. These polymers are at the interface between inorganic and organic polymers, and have unusual material properties. Polyphosphazenes have found industrial applications, mainly because of their high thermal stability. They have also been used in investigations for the formulation of controlled drug delivery systems (Allcock, 1990). Polyphosphazenes are interesting biomaterials in many respects. Most polyphosphazenes are not degradable, but the incorporation of certain chemical groups into the polymer structure may render these materials degradable (Allcock, 1999). They have been claimed to be biocompatible, and their chemical structure provides a readily accessible “pendent chain” to which various drugs, peptides or other biological compounds can be attached and later released via hydrolysis. Polyphosphazenes have been examined for use in skeletal tissue regeneration and have been shown to be osteocompatible (Qiu and Zhu, 2000; Nair et al., 2006). Another novel use of polyphosphazenes is in vaccine design, where these materials were used as immunological adjuvants (Andrianov et al., 1998).
Poly(propylene fumarate)s (PPF) and Poly(trimethylene carbonate) (PTMC) are polyester based materials, like PGA and PLA. PPF is most commonly synthesized using a two-step process (Shung et al., 2002) that results in the formation of either a viscous liquid or a yellow solid. The formation of the liquid or the solid depends on the molecular weight of the synthesized polymer, with high molecular weights yielding the solid, and low molecular weights yielding the liquid. PPF, unlike PGA and PLA, is an unsaturated polyester. Since a double bond is present within the backbone, various chemical modifications, including cross-linking, can be used to alter the physico-mechanical properties of the material. Cross-linking of the liquid polymer allows the material to be processed and molded into any number of shapes both pre-implantation and in situ. The ability to harden the materials in situ via thermal or photo initiated processes coupled with their osteoconductivity in vitro (Yaszemski et al., 1995) and in vivo (Lewandrowski et al., 2000; Fisher et al., 2002) has prompted exploration of these materials as cements and scaffolding materials for bone. PFF is degraded via hydrolytic cleavage of its ester bonds yielding propylene glycol and fumarate upon complete degradation. Both propylene glycol and fumarate are non-toxic in vivo. PPF, like PGA and PLA, releases acidic products upon degradation. However, PPF seems to lack susceptibility to autocatalyzed degradation as cross-linking density governs degradation rates to a far greater extent than local pH (Timmer et al., 2003).
PTMC is an elastomeric polyester with high flexibility but poor mechanical properties. It is synthesized via the ring opening polymerization of trimethyl carbonate and has a degradation rate that is 20 times slower than poly (ε-caprolactone) in phosphate buffered saline in vitro (Zhu et al., 1991). Interestingly, the in vivo degradation time of PTMC is significantly faster than its in vitro degradation, thus strongly suggesting a role for enzymatic degradation (Zhu et al., 1991; Zhang et al., 2006). This makes PTMC a polymer that truly undergoes “biodegradation.” Copolymers of trimethylene carbonate and glycolide have been used as sutures and orthopedic fixation devices under the brand names Maxon™ and Acufex™ respectively. Additionally, a terpolymer of trimethylene carbonate, glycolide, and dioxane has been used to create a degradable suture under the brand name of BioSyn™ (Suzuki and Ikada, 2005).
Polyurethanes represent a class of elastomeric materials that have been used in the fabrication of a number of medical devices. While considered to be resistant to hydrolysis, polyurethane pacemaker leads were found to degrade through oxidative, enzymatic, and cellular mechanisms (Mcmillin, 1991). Recent work has focused on the incorporation of degradable linkages to impart predictable degradation kinetics. Linkages based on chemicals such as lysine diisocyanate have allowed the creation of degradable polyurethanes that are non-toxic. Polyurethanes are of particular interest in implantables because they have robust mechanical properties and good biocompatibility (Gunatillake and Adhikari, 2003). Non-degradable polyurethanes are currently used in a number of clinical devices, but have gained particular interest as vascular grafts. This stems from the observation that protein adsorption, and therefore clotting, is less pronounced on polyurethanes as compared to other materials. Degradable polyurethanes have also been investigated as artificial skin (Bruin et al., 1990), orthopedic injectables, and tissue engineering scaffolds (Degrapol™) (Saad et al., 1997; Zhang et al., 2000).
Applications of Synthetic, Degradable Polymers as Biomaterials
Classification of Degradable Medical Implants
Some typical short-term applications of biodegradable polymers are listed in Table I.2.6.1. From a practical perspective, it is convenient to distinguish between five main types of degradable implants: the temporary support device; the temporary barrier; the drug delivery device; the tissue-engineering scaffold; and the multi-functional implant.
A temporary support device is used in those circumstances in which the natural tissue bed has been weakened by disease, injury or surgery, and requires some artificial support. A healing wound, a broken bone or a damaged blood vessel are examples of such situations. Sutures, bone fixation devices (e.g., bone nails, screws or plates), and vascular grafts would be examples of the corresponding support devices. In all of these instances, the degradable implant would provide temporary, mechanical support until the natural tissue heals and regains its strength. In order for a temporary support device to work properly, a gradual stress transfer should occur: as the natural tissue heals, the degradable implant should gradually weaken. The need to adjust the degradation rate of the temporary support device to the healing of the surrounding tissue represents one of the major challenges in the design of such devices.
A successful example of a temporary support device, the first synthetic, degradable sutures were made of PGA and became available under the trade name “Dexon™” in 1970. This represented the first routine use of a degradable polymer in a major clinical application (Frazza and Schmitt, 1971). Later, copolymers of lactic acid and glycolic acid were developed. The widely used “Vicryl™” suture, for example, is a 90:10 copolymer of lactic and glycolic acid, introduced into the market in 1974. Sutures made of polydioxanone (PDS) became available in the USA in 1981. Sutures of ε-caprolactone and trimethylene carbonate with glycolide have been developed under the names “MonoCryl™” and “Maxon™,” respectively. “Biosyn™” is a terpolymer suture composed of trimethylene carbonate, glycolide, and dioxane. In spite of extensive research efforts in many laboratories, no other degradable polymers are currently used to any significant extent in the formulation of degradable sutures.
A temporary barrier has its major medical use in the prevention of postsurgical adhesions. During surgical closure, barrier devices are interposed between tissues that must remain separate following surgery. Adhesions are formed between two tissue sections by clotting of blood in the extravascular tissue space, followed by inflammation and fibrosis. If this natural healing process occurs between surfaces that were not meant to bond together, the resulting adhesion can cause pain, functional impairment, and problems during subsequent surgery. Surgical adhesions are a significant cause of morbidity, and represent one of the most significant complications of a wide range of surgical procedures such as abdominal, cardiac, spinal, and tendon surgery. A temporary barrier could take the form of a thin polymeric film or a mesh-like device that would be placed between adhesion-prone tissues at the time of surgery. To be useful, such a temporary barrier would have to prevent the formation of scar tissue connecting adjacent tissue sections, followed by the slow resorption of the barrier material. This sort of barrier has also been investigated for sealing of breaches of the lung tissue that cause air leakage.
Another important example of a temporary barrier is in the field of skin reconstruction. Several products have become available that are generally referred to as “artificial skin.” The first such product consists of an artificial, degradable collagen/glycosaminoglycan matrix that is placed on top of the skin lesion to stimulate the regrowth of a functional dermis. Another product consists of a degradable collagen matrix with pre-seeded human fibroblasts. Synthetic degradable polymers have also been used to create “artificial skin.” Dermagraft™ is an artificial skin product that uses a PLGA-based mesh scaffold to provide structural support. Again, the goal is to stimulate the regrowth of a functional dermis. These products are used in the treatment of burns and other deep skin lesions, and represent an important application for temporary barrier type devices.
Implantable drug delivery devices and injectable polymer-drug depot delivery systems are capable of delivering drugs to specific locations over an extended period of time. They are, by necessity, temporary devices, for the drug reservoir will eventually be depleted or the need for delivery of a specific drug will be eliminated once the disease is adequately treated. The development of implantable drug delivery systems is probably the most widely investigated application of degradable polymers (Nair and Laurencin, 2007). One can expect that the future acceptance of implantable and injectable drug delivery devices by physicians and patients alike will depend on the availability of degradable systems that do not have to be surgically explanted.
Since PLA and PGA have an extensive safety profile based on their use as sutures, these polymers have been very widely investigated in the formulation of implantable controlled release devices. Several implantable, controlled release formulations based on copolymers of lactic and glycolic acid have already become available. However, a wide range of other degradable polymers has also been investigated. Particularly noteworthy is the use of a new type of polyanhydride in the formulation of an intracranial, implantable device for the administration of BCNU (a chemotherapeutic agent) to patients suffering from glioblastoma multiforme, a usually lethal form of brain cancer (Gliadel™). A clinical trial by Westphal et al. in 2003 demonstrated a 29% reduction in the risk of death in patients treated with Gliadel™ versus those who were not (Westphal et al., 2003). The topic of drug delivery systems is discussed in more detail in Chapter II.5.16.
The term tissue engineering scaffold will be used in this chapter to describe a degradable implant that is designed to act as an artificial extracellular matrix by providing space for cells to grow into and reorganize into functional tissue (Hutmacher, 2001).
It has become increasingly obvious that man-made implantable prostheses do not function as well as the native tissue, or maintain the functionality of native tissue over long periods of time. Therefore, tissue engineering has emerged as an interdisciplinary field that utilizes degradable polymers, among other substrates and biologics, to develop treatments that enhance the body’s potential to heal itself without the need for permanently implanted, artificial prosthetic devices. In the ideal case, a tissue engineering scaffold is implanted to restore lost tissue function, maintain tissue function or enhance existing tissue function (Langer and Vacanti, 1993). These scaffolds can take the form of a felt-like material obtained from knitted or woven fibers or from fiber meshes. Alternatively, various processing techniques can be used to obtain foams or sponges. For all tissue engineering scaffolds, pore volume and pore interconnectivity are key properties, as cells need to be able to migrate and grow throughout the entire scaffold. Thus, industrial foaming techniques used, for example, in the fabrication of furniture cushions, are not applicable to the fabrication of tissue engineering scaffolds, as these industrial foams are designed to contain “closed pores,” while tissue engineering scaffolds require an “open pore” structure to attain the requisite interconnectivity. Tissue engineering scaffolds may be pre-seeded with cells in vitro prior to implantation. Alternatively, tissue-engineering scaffolds may consist of a cell-free structure that is either invaded and “colonized” by cells after implantation, or a surgeon may seed the scaffold with autologous cells intraoperatively (e.g., combining bone marrow aspirate with demineralized bone matrix to treat bone defects). In either case, the tissue engineering scaffold must allow the formation of functional tissue in vivo, followed by the safe resorption of the scaffold material.
There has been some debate in the literature as to the exact definition of the related term “guided tissue regeneration” (GTR). Guided tissue regeneration is a term traditionally used in dentistry. This term sometimes implies that the scaffold encourages the growth of specific types of tissue. For example, in the treatment of periodontal disease, periodontists use the term “guided tissue regeneration” when using implants that favor new bone growth in the periodontal pocket over soft tissue ingrowth (scar formation) (Asikainen et al., 2006).
One of the major challenges in the design of tissue engineering scaffolds is the need to adjust the rate of scaffold degradation to the rate of tissue healing. Depending upon the application of the scaffold, the polymer may need to maintain integrity in the order of days to months. Scaffolds intended for the reconstruction of bone illustrate this point: in most applications, the scaffold must maintain some mechanical strength to support the bone structure while new bone is formed. Premature degradation of the scaffold material can be as detrimental to the healing process as a scaffold that remains intact for excessive periods of time. The future use of tissue engineering scaffolds has the potential to revolutionize the way aging, trauma, and disease related loss of tissue function could be treated.
Multifunctional devices, as the name implies, combine several of the above-mentioned functions within a single device. Over the last few years, there has been a trend toward increasingly sophisticated applications for degradable biomaterials. Usually, these applications envision the combination of several functions within the same device and require the design of custom-made materials with a narrow range of predetermined material properties. For example, the availability of biodegradable bone nails and bone screws made of ultra high strength poly(lactic acid) opens the possibility of combining the “mechanical support” function of the device with a “site-specific drug delivery” function: a biodegradable bone nail that holds the fractured bone in place and can simultaneously stimulate the growth of new bone tissue at the fracture site by slowly releasing bone growth factors (e.g., bone morphogenetic protein or transforming growth factor β) throughout its degradation process.
Likewise, biodegradable stents for implantation into coronary arteries are currently being investigated (Ramcharitar and Serruys, 2008). The stents are designed to mechanically prevent the collapse and restenosis (reblocking) of arteries that have been opened by balloon angioplasty. Ultimately, the stents could deliver an anti-inflammatory or anti-thrombogenic agent directly to the site of vascular injury. Again, it would potentially be possible to combine a mechanical support function with site-specific drug delivery.
Various functional combinations involve the tissue engineering scaffold. Perhaps the most important multi-functional device for future applications is a tissue engineering scaffold that also serves as a drug delivery system for cytokines, growth hormones or other agents that directly affect cells and tissue in the vicinity of the implanted scaffold.
The Process of Bioerosion
One of the most important prerequisites for the successful use of a degradable polymer for any medical application is a thorough understanding of the way the device will degrade/erode and ultimately resorb from the implant site. Within the context of this chapter, we are limiting our discussion to the case of a solid, polymeric implant. The transformation of such an implant into water-soluble material(s) is best described by the term “bioerosion.” The bioerosion process of a solid, polymeric implant is associated with macroscopic changes in the appearance of the device, changes in its physico-mechanical properties and in physical processes such as swelling, deformation or structural disintegration, weight loss and the eventual depletion of drug or loss-of-function. It is important to note that the bioerosion of a solid device is not necessarily due to the chemical cleavage of the polymer backbone, cross-links or side chains. Rather, simple solubilization of the intact polymer chains, for instance, due to changes in pH, may also lead to the erosion of a solid device.
Two distinct modes of bioerosion have been described in the literature: bulk erosion and surface erosion. In bulk erosion, the rate of water penetration into the solid device exceeds the rate at which the polymer is transformed into water-soluble material(s). Consequently, an erosion process that occurs throughout the entire volume of the solid device follows the uptake of water. Due to the rapid penetration of water into the matrix of the polymer, most of the currently available polymers will give rise to bulk eroding devices. In a typical bulk erosion process, cracks and crevices will form throughout the device, which may crumble into pieces. A good illustration for a typical bulk erosion process is the disintegration of an aspirin tablet that has been placed into water. Depending on the specific application, the uncontrollable tendency of bulk eroding devices to crumble into little pieces can be a disadvantage.
Alternatively, in surface erosion, the bioerosion process is limited to the surface of the device. Therefore, the device will become thinner with time, while maintaining its structural integrity throughout much of the erosion process. In surface erosion, the polymer must impede the rapid transport of water into the interior of the device. In addition, the rate at which the polymer is transformed into water-soluble material(s) has to be fast relative to the rate of water penetration into the device. Under these conditions, scanning electron microscopic evaluation of surface eroding devices has sometimes shown a sharp border between the eroding surface layer and the intact polymer in the core of the device (Mathiowitz et al., 1990).
The determination of whether a material is surface eroding or bulk eroding is dependent on several factors. Since the transport of water into the interior of a device is necessary for erosion, polymer hydrophobicity plays an important role. However, polymer chemistry is not the sole mediator of whether a material undergoes surface or bulk erosion. Von Burkersroda et al. have proposed that the physical architecture of the device may determine if it is surface or bulk eroding (Von Burkersroda et al., 2002). In this publication, a critical thickness is calculated for a selection of polymeric materials. Above this critical thickness, a material is surface eroding, while below this critical thickness it is bulk eroding. This critical thickness is so large for some polymer classes (in the order of meters) that with regard to implantable devices they cannot practically be surface eroding. Therefore, surface eroding devices have so far been obtained only from a small number of hydrophobic polymers containing hydrolytically highly reactive linkages in the backbone. Currently, polyanhydrides, and poly(ortho ester)s are the best-known examples of polymers that can be fabricated into surface eroding devices.
A possible exception to this general rule is enzymatic surface erosion. The inability of enzymes to penetrate into the interior of a solid, polymeric device may result in an enzyme-mediated surface erosion mechanism. In the case of enzymatic surface erosion, the term biodegradation may be correctly applied, for enzymatic degradation of the polymer backbone dominates the rate of hydrolysis under physiological conditions. It is possible for a polymeric device to be classified as bulk eroding and biodegrading if the polymer imbibes sufficient water to allow enzymes to penetrate the swollen network of polymer chains.
Mechanisms of Chemical Degradation
Although bioerosion can be caused by the solubilization of an intact polymer, chemical degradation of the polymer is usually the underlying cause for the bioerosion of a solid, polymeric device. Several distinct types of chemical degradation mechanisms have been identified (Figure I.2.6.2) (Rosen et al., 1988). Chemical reactions can lead to cleavage of cross-links between water-soluble polymer chains (Mechanism I), to cleavage of polymer side chains resulting in the formation of polar or charged groups (Mechanism II), or to cleavage of the polymer backbone (Mechanism III). Obviously, combinations of these mechanisms are possible: for instance, a cross-linked polymer may first be partially solubilized by the cleavage of cross-links (Mechanism I), followed by the cleavage of the backbone itself (Mechanism III). It should be noted that water is key to all of these degradation schemes. Even enzymatic degradation occurs in an aqueous environment.
FIGURE I.2.6.2 Mechanisms of chemical degradation, adapted from Gombotz and Pettit (1995). Mechanism I involves the cleavage of degradable cross-links between water-soluble polymer chains. Mechanism II involves the cleavage or chemical transformation of hydrophobic polymer side chains, resulting in the formation of charged or polar groups. The presence of charged or polar groups leads then to the solubilization of the intact polymer chain. Mechanism III involves the cleavage of unstable linkages in the polymer backbone, followed by solubilization of the low molecular weight fragments.
Since the chemical cleavage reactions described above can be mediated by water or by biological agents, such as enzymes and microorganisms, it is possible to distinguish between hydrolytic degradation and biodegradation. It has often been stated that the availability of water is virtually constant in all soft tissues and varies little from person to person. On the other hand, the levels of enzymatic activity may vary widely not only from person to person, but also between different tissue sites in the same person. Thus polymers that undergo hydrolytic cleavage tend to have more predictable in vivo erosion rates than polymers whose degradation is mediated predominantly by enzymes. The latter polymers tend to be generally less useful as degradable medical implants.
Factors that Influence the Rate of Bioerosion
Although the solubilization of intact polymer, as well as several distinct mechanisms of chemical degradation have been recognized as possible causes for the observed bioerosion of a solid, polymeric implant, virtually all currently available implant materials erode due to the hydrolytic cleavage of the polymer backbone (Mechanism III in Figure I.2.6.2). We therefore limit the following discussion to solid devices that bioerode in this manner.
The main factors that determine the overall rate of the erosion process are the chemical stability of the hydrolytically-susceptible groups in the polymer backbone, the hydrophilic–hydrophobic character of the repeat units, the morphology of the polymer, the initial molecular weight and molecular weight distribution of the polymer, the fabrication process used to prepare the device, the presence of catalysts, additives or plasticizers, and the geometry (specifically the surface area to volume ratio) of the implanted device.
The susceptibility of the polymer backbone to hydrolytic cleavage is probably the most fundamental parameter. Anhydride bonds are the most susceptible to hydrolysis, followed by carbonate, ester, urethane, ortho ester, and amide bonds (Figure I.2.6.3) (Gombotz and Pettit, 1995). Thus, polyanhydrides will tend to degrade faster than polyesters, which in turn will have a higher tendency to bioerode than polyamides. Based on the known susceptibility of the polymer backbone structure to hydrolysis, it is possible to make predictions about the bioerosion of a given polymer.
FIGURE I.2.6.3 Relative hydrolytic susceptibility of degradable bonds at physiologic pH, adapted from Gombotz and Pettit (1995). In general, polymers with high rates of main chain bond hydrolysis will degrade more quickly than polymers with low rates of main chain bond hydrolysis, but many other factors, such as pH and autocatalysis at low pHs, copolymer composition and structure, percent crystallinity, and related equilibrium water content, and geometry of the degradable polymer, will influence the polymer/device degradation rate. For example, at acidic pHs, the order in this table would change, because poly(ortho ester)s will have a higher rate of hydrolysis, perhaps somewhere between polyanhydrides and polycarbonates. In a related example, with thicker samples of polyesters such as PLGA, the acidic byproducts (lactic acid and glycolic acid) may accumulate within the bulk, lowering the internal bulk pH and leading to more rapid degradation within the bulk of the sample than at the surface.
However, the actual erosion rate of a solid polymer cannot be predicted on the basis of the polymer backbone structure alone. The observed erosion rate is strongly dependent on the ability of water molecules to penetrate into the polymeric matrix. The hydrophilic versus hydrophobic character of the polymer, which is a function of the structure of the monomeric starting materials, can therefore have an overwhelming influence on the observed bioerosion rate. For instance, the erosion rate of polyanhydrides can be slowed by about three orders of magnitude when the less hydrophobic sebacic acid is replaced by the more hydrophobic bis(carboxy phenoxy)propane as the monomeric starting material (Tamada and Langer, 1993). Likewise, devices made of PGA erode faster than identical devices made of the more hydrophobic PLA, although the ester bonds have about the same chemical reactivity toward water in both polymers.
The observed bioerosion rate is further influenced by the morphology of the polymer. Polymers can be classified as either semicrystalline or amorphous. At body temperature (37°C) amorphous polymers with Tg (glass transition temperature) above 37°C will be in a glassy state, and polymers with a Tg below 37°C will be in a rubbery state. In this discussion it is therefore necessary to consider three distinct morphological states: semicrystalline; amorphous-glassy; and amorphous-rubbery.
In the crystalline state, the polymer chains are densely packed and organized into crystalline domains that resist the penetration of water. Consequently, backbone hydrolysis tends to occur in the amorphous regions of a semicrystalline polymer and at the surface of the crystalline regions. This phenomenon is of particular importance to the erosion of devices made of poly(L-lactic acid) and poly(glycolic acid), which tend to have high degrees of crystallinity, around 50%.
Another good illustration of the influence of the polymer morphology on the rate of bioerosion is provided by a comparison of poly(L-lactic acid) and poly(D,L-lactic acid); although these two polymers have chemically identical backbone structures and an identical degree of hydrophobicity, devices made of poly(L-lactic acid) tend to degrade much more slowly than identical devices made of poly(D,L-lactic acid). The slower rate of bioerosion of poly(L-lactic acid) is due to the fact that this stereoregular polymer is semicrystalline, while the racemic poly(D,L-lactic acid) is an amorphous polymer.
Likewise, a polymer in its glassy state is less permeable to water than the same polymer when it is in its rubbery state. This observation could be of importance in cases where an amorphous polymer has a glass transition temperature that is not far above body temperature (37°C). In this situation, water absorption into the polymer could lower its Tg below 37°C, resulting in abrupt changes in the bioerosion rate.
The manufacturing process may also have a significant effect on the erosion profile. For example, Mathiowitz and co-workers (Mathiowitz et al., 1990) showed that polyanhydride microspheres produced by melt encapsulation were very dense and eroded slowly, whereas when the same polymers were formed into microspheres by solvent evaporation, the microspheres were very porous (and therefore more water permeable) and eroded more rapidly.
The above examples illustrate an important technological principle in the design of bioeroding devices; the bioerosion rate of a given polymer is not an unchangeable property, but depends to a very large degree on readily controllable factors such as the presence of plasticizers or additives, the manufacturing process, the initial molecular weight of the polymer, and the geometry of the device.
Storage Stability, Sterilization, and Packaging
It is important to minimize premature polymer degradation during fabrication and storage. Traces of moisture can seriously degrade even relatively stable polymers, such as poly(bisphenol A carbonate), during injection molding or extrusion. Degradable polymers are particularly sensitive to hydrolytic degradation during high temperature processing. The industrial production of degradable implants therefore often requires the construction of facilities with controlled atmospheres where the moisture content of the polymer and the ambient humidity can be strictly controlled.
After fabrication, γ-irradiation or exposure to ethylene oxide may be used for the sterilization of degradable implants. Both methods have disadvantages, and as a general rule the choice is between the lesser of two evils. γ-Irradiation at a dose of 2 to 3 Mrad can result in significant backbone degradation. Since the aliphatic polyesters PLA, PGA, and PDS are particularly sensitive to radiation damage, these materials are usually sterilized by exposure to ethylene oxide, and not by γ-irradiation. Unfortunately, the use of the highly dangerous ethylene oxide gas represents a serious safety hazard by potentially leaving residual traces in the polymeric device. Polymers sterilized with ethylene oxide must be degassed for extended periods of time.
Additionally, for applications in tissue engineering, biodegradable scaffolds may be pre-seeded with viable cells or may be impregnated with growth factors or other biologics. There is currently no method that could be used to sterilize scaffolds that contain viable cells without damaging the cells. Therefore, such products must be manufactured under sterile conditions, and must be used within a very short time after manufacture. Currently, a small number of products containing preseeded, living cells are in clinical use. These products are expensive, are shipped in special containers, and have a short shelf life.
Likewise, it has been shown that sterilization of scaffolds containing osteoinductive or chondroinductive agents leads to significant losses in bioactivity, depending on the sterilization method used (Athanasiou et al., 1998). The challenge of producing tissue engineering scaffolds that are pre-seeded with viable cells or that contain sensitive biological agents has not yet been fully solved.
After sterilization, degradable implants are usually packaged in air-tight aluminum-backed plastic-foil pouches. In some cases, refrigeration may also be required to prevent backbone degradation during storage.
Evaluating Toxicity of Bioerodible Materials
A polymeric medical device is not completely pure. Device toxicity could originate from many sources, such as those listed below:
Polymer impurities: Initiators, residual monomers and oligomers, byproducts.
Solvents: Contamination from solvents used in synthesis, to clean equipment or in fabrication (e.g., solvent casting processes to produce films, microspheres or fibers).
Additives: Stabilizers, emulsifiers, antioxidants, plasticizers.
Processing byproducts: Byproducts of high-temperature processing and sterilization.
Degradation products: Products of hydrolytic or enzymatic degradation and their subsequent metabolites.
Testing for potential toxicity is governed by ISO 10993 and the FDA G95-1 Memorandum. Cytotoxicity, sensitization, irritation, carcinogenicity, genotoxicity, hemocompatibility, and systemic toxicity are evaluated using cell culture methods and animal testing:
Mammalian cell culture: Indirect (agar overlay) or direct sample contact or exposure to device extract.
Animal testing: Mouse, rat, guinea pig, and dog models are typically used with evaluation by histopathology, gross observation of redness or swelling, lymph node extraction, micronucleous assay, and blood thrombogenicity, lysis, and complement activation.
Glossary of key Terms
Degradation A chemical process by which covalent bonds are cleaved.
Erosion A process that results in physical changes to the size, shape, or mass of a material.
Biodegradation Degradation as a consequence of a biological agent, such as an enzyme, cell, or microorganism.
Bioerosion Conversion of a water-insoluble material into a water-soluble material(s) under physiological conditions, regardless of the mechanism of conversion.
Surface erosion A process of polymer erosion where the rate of conversion to water-soluble material(s) exceeds the rate of water infiltration into the interior of the device. As a consequence, material erodes from the surface of the device, while the core remains structurally consistent.
Bulk erosion A process of polymer erosion where the rate of water penetration into the interior of the device exceeds the rate of conversion into water-soluble material(s). As a consequence, material erosion is equally likely to happen throughout the bulk of the device.
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