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Corrosion in Biomedical Applications
Corrosion in Biomedical Applications S Virtanen, University of Erlangen-Nürnberg, Erlangen, Germany © 2018 Elsevier Inc. All rights reserved.
Introduction Human Body as a Corrosive Environment Possible Corrosion Modes in Biological Environments Corrosion Behavior of Selected Passive Alloys Used in Biomedical Applications Corrosion of Biodegradable Metals Concluding Remarks Further Reading
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Introduction Corrosion behavior of metallic materials used in biomedical applications (e.g., implants) is of high importance for biocompatibility, and in general for successful and safe applications. Strong degradation of a permanent implant would lead to failure by loss of the device integrity; for high-corrosion resistant metals and alloys that are typically used in such applications, strong active dissolution is, however, not expected. Nevertheless, specific corrosion modes are possible for passive metals and alloys. In addition, even small amount of corrosion, as takes place upon passive dissolution, can be detrimental, if the corrosion products are toxic or lead to adverse biological reactions. As the biological response to metal release depends on speciation and concentration of the released species, detailed information on the corrosion mode and rate is required. For biodegradable, temporary implants that are intended to corrode in the host body, it is even more self-evident that the nature of the corrosion products must be considered in view of their biological interactions. Corrosion behavior of metals and alloys is similar in the human body as in other aqueous solutions with a corresponding electrolyte composition. Hence, passive metals and alloys, such as Ti and its alloys, typically remain in the passive state also in the biological environment. Similarly, actively corroding materials, such as Mg and its alloys, would be expected to easily degrade in the body fluids. However, the specific chemistry of body fluids, as well as the presence of biomolecules and cells, can have very strong material-dependent influences on the corrosion behavior and metal release. Moreover, implant surfaces can be exposed to additional loads such as wear or mechanical stresses due to the function and design of the device. Therefore, not only purely chemical degradation but in addition mechano-chemical corrosion processes need to be considered. It should be emphasized that corrosion behavior of materials is always a system behavior, hence depending on the combination of the material (composition and microstructure), the environment (e.g., chemistry, temperature, hydrodynamic conditions), and the construction.
Human Body as a Corrosive Environment For metallic implants used in the body for healing, the first contact with the human body upon surgery is with blood. Therefore, simulated body fluids that are used to study the corrosion behavior in the laboratory often mimic the inorganic composition of blood. For this, the corrosive environment can be described as a 0.9% NaCl solution, with small amounts of other inorganic species, such as Ca2 þ, PO4 3 , and HCO3 , buffered to pH 7.4, at a temperature of c. 37 C. Chloride ions are aggressive species and can lead to breakdown of passivity of many metals and alloys, whereas phosphates and carbonates can lead to formation of (partially) protective corrosion product layers. Precipitation of Ca-phosphates from the body fluids can take place on metal and alloy surfaces, and this can further influence the electrochemical reactions. The body temperature of 37 C can accelerate corrosion reactions as compared with the room temperature. Even though body fluids are buffered by the CO2/HCO3 system to pH 7.4, local variations of the pH value can occur; upon inflammatory conditions and directly following surgery, lower pH values are expected. Also, if the implant design leads to formation of crevices, acidification inside the crevice electrolyte due to metal cation hydrolysis reactions can take place. Depending on the targeted biomedical application or on the type of contact of metals and alloys with the human body, different types of exposure scenarios exist; therefore the study of the corrosion behavior should consider the specific chemistry and biology for each case. In dermal contact, in lung or stomach environment, or in dental applications, the material surface is exposed to different chemical, biological, and mechanical interactions. Moreover, mass transfer conditions that can significantly influence corrosion reactions depend on the site of implantation. The redox potential of the biological environment depends on one hand side on the oxygen content (that also is locationdependent) and on the other side on the presence of a variety of other types of redox species; some of which can stem from biological reactions. Especially noteworthy are reactive oxygen species, such as hydroxide radicals and hydrogen peroxidedthese can have effects both on the biological system and on metal surface reactions.
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In addition to the variety of inorganic components of body fluids, presence of organic componentsdbiomolecules, proteins, cells, or bacteriadcan further influence corrosion reactions. All these species show a variety of interactions with the metal and alloy surfaces (and with each other), making the corrosion scenario highly complex. Moreover, they lead to a strongly time-dependent, dynamic system behavior, in that processes such as cell adhesion on surfaces lead to a number of possible follow-up reactions of the living system. Influence of proteins on corrosion has been frequently investigated, however the findings are not quite straightforward in view of the protein influence on corrosion rates. On one hand side this can be due to different experimental approaches, for instance in some studies single protein addition to simulated body fluids is used, while in other studies serum is used as a source of proteins. Moreover, it should be considered that proteins have multiple effects on corrosion and metal release. Proteins can adsorb on metal surfaces and such protein adsorption layers can induce corrosion protection by blocking the active sites of the surface. Protein adsorption depends on the nature of metal surface and of the specific protein, as protein adsorption is influenced by surface energy and surface charges. On the other hand, proteins can bind metal cations, and this could enhance metal dissolution reactions. This effect is also metal/protein-specific. In a tribocorrosion scenario (see what follows), biofilms forming on the surface are influenced by proteins, and may act as lubricants reducing friction. Cells can influence the biomaterial surface reactions by a variety of operative mechanisms. Macrophage cells are central to wound healing and inflammatory tissue reactions, and therefore in the early stage after surgery they accumulate around the foreign materials. Activated macrophages generate reactive oxygen species, and for passive metals and alloys these species can change the thickness and/or composition of passive films. For Ti-based materials it should be noted that H2O2 possibly generated by cells leads to enhanced solubility of the TiO2 passive film. For actively corroding, biodegradable metals (for instance, Mg alloys), reactive oxygen species generated by activated macrophages have been demonstrated to significantly increase the corrosion rate. Other cells, such as fibroblasts and osteoblasts, can adhere to the material surface by contacting cell-adhesive protein adsorption layers. The cells generate an extracellular matrix, containing polysaccharides, water, fibrous proteins, and cell-adhesive proteins. These cell adhesion layers and the extracellular matrix can on one hand side influence diffusion reactions to and from the metal surface, thereby influencing kinetics of electrochemical reactions. The electrolyte chemistry on the metal surface below cell layers can change, due to the diffusion hindrance as well as by cell metabolism products. For instance, pH decrease has been reported to take place near fibroblasts cultured on stainless steel, and this could locally accelerate metal dissolution (“cell-induced crevice corrosion”). On the other side, the cell adhesion layers can inhibit corrosion, by simple blocking of the active surface. As a direct probing of surface reactions in presence of cells on the surface is challenging, there are still many open questions on the mechanisms of cell-influence on metal corrosion (also considering that different metallic materials will show different response to the biological environment). In the real application (in vivo corrosion scenario), the situation is highly dynamic. Following protein adsorption and cell adhesion on the biomaterial surface, complex follow-up reactions take place, and these reactions depend on the material surface as well as on the location of implantation. The metal surface corrosion reactions will adapt to the changing environment, but there clearly is a lack of detailed understanding on the spatial and temporal variations of the implant surface/biology interactions. For instance, blood flow on the implant surface can be expected to influence any diffusion-controlled electrochemical reaction. However, the hydrodynamic conditions in the vicinity of an implant surface would strongly depend on the location of implantation and can change with time (e.g., due to tissue coverage of the implant with time as compared with a bare metal surface exposed to blood flow). Not only the anatomical location of an implant can have a significant influence on the corrosion behavior, but moreover the surgical procedure.
Possible Corrosion Modes in Biological Environments Corrosion of metals and alloys is determined by the thermodynamic and kinetic factors of the electrochemical reactions driving corrosion. For permanent implants, high-corrosion resistant materials that are passive under a large range of environmental conditions are employed. These materials include Ti and its alloys, CoCrMo alloys, and specific grades of stainless steels. In the next section, the corrosion behavior of these three material groups in a biological environment is shortly summarized. Another section is dedicated to considerations of the corrosion behavior of biodegradable metals (mainly Mg- and Fe-based alloys). First, general types of corrosion that can take place in the biomedical applications are introduced. In the case of stable passivity with no localized corrosion occurring, passive dissolution rates are orders of magnitude lower than dissolution in the active state. However, even such small passive dissolution rates may not be negligible in all cases, especially as concerns the composition of the passive alloys used; these alloys contain for instance Al, Co, Cr, and Ni; the release of these elements raises questions on the biocompatibility. Also, the protectiveness of passive films and hence the rate of passive dissolution depend on the alloy and the environment. In the biomedical applications, proteins and cell metabolism products can be involved in passive dissolution and metal release. Moreover, local acidification due to for instance inflammation or metal cation hydrolysis in sites of hindered mass transfer can increase passive dissolution rates. As the biological environment is a Cl-containing solution, it is straightforward to consider pitting corrosion as a possible degradation mode. Pitting corrosion leads to formation of cavities, surrounded by an intact passivated surface. For many materials, pit propagation is an autocatalytic process, related to formation of aggressive chemistry (acidification, accumulation of aggressive chloride, oxygen depletion) in the pit local environment. Moreover, an unfavorable area ratio of small anodes (pits) and a large cathode (the passive surface) accelerates local dissolution rates. Pitting corrosion leads to an increased metal release as compared with
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uniform passive dissolution, but moreover it can be dangerous for implants under mechanical load, as pits could act as stressconcentration sites and could possibly lead to local loss of the mechanical integrity of the device. Crevice corrosion occurs in sites on the metal surface where mass transfer is limited, leading to formation of locally aggressive electrolyte in narrow crevices or under deposits; similarly to what occurs in pits. Materials that are susceptible to pitting corrosion are also prone to crevice corrosion, but far less aggressive conditions are required to trigger crevice corrosion than to initiate pitting. In biomedical applications, crevices can occur in modular implants or at interfaces between implant surface and bone cement (or body tissues). Galvanic corrosion that is driven by a potential difference of dissimilar materials in contact and in a joint electrolyte is a possible corrosion risk in modular implants. The corrosion rate of the less noble metal is accelerated by the contact with a more noble metal, and the most critical case is given by a small area of a less noble metal forming a local anode in contact with a large area of a more noble material. For passive metals and alloys, the resistance given by the passive film, however, decreases the risk of galvanic corrosion. If any type of localized breakdown of passivity is possible, galvanic coupling with a more noble metal can significantly accelerate the propagation of localized corrosion of the less noble metal. Tribocorrosiondthat is, the simultaneous action of mechanical wear and corrosive attackdis a degradation mode for many biomedical implants where micromovements between surfaces take place and the implant surface is under mechanical loading. For biomedical implants fretting corrosion that takes place at the interface of two close surfaces subjected to small oscillatory slip and concurrent corrosive action is a highly relevant failure mode, for instance in hip and knee replacements. Here, information is required not only on the materials (including the tribological or fretting partners) and the environment, but moreover the biomechanical situation needs to be considereddhence the implant design and loading are important. As in most mechano-chemical corrosion modes, synergetic interactions between the mechanical wear and chemical corrosion can lead to drastically increased damage as compared with the sole action of wear or corrosion. For passive materials, fretting wear can easily mechanically destroy thin passive films. Therefore, during fretting corrosion continuous activation/repassivation cycles take place, and the repassivation ability of the materials is of critical importance for resulting damage and for metal release.
Corrosion Behavior of Selected Passive Alloys Used in Biomedical Applications For permanent implants, high-corrosion resistant alloys such as Ti-based materials (e.g., cp-Ti, Ti–6Al–4V, Ti–6Al–7Nb, Ni–Ti shape memory alloys, and others), Co–Cr–Mo alloys, and specific grades of stainless steels (Fe–Cr–Ni alloys) are used. All these materials in general show a very good passive behavior in a variety of environments. Considering specific localized corrosion modes of passive metals and alloys, significant differences in the resistance of the different alloy groups exist. Moreover, the exact chemical composition and microstructure of the specific alloy plays a significant role for the corrosion behavior. For instance, Co–Cr–Mo alloys exist in low-carbon and high-carbon variation, and they can be cast, forged, wrought, or produced by powder metallurgical routes; the processing and the carbon content lead to different carbide content and distribution in the alloydand this in its turn influences the corrosion behavior. Even in cases of no localized corrosion attack or tribocorrosion, passive dissolution leads to metal release and this can strongly vary for different passive materials under different conditions. Typical for high-corrosion resistant alloys is that they show a stable passivity under a large pH range. However, many of the biomedical alloys contain alloying elements that are not inherently stable under the conditions of biological solutions or in slightly acidified conditions (for instance, Fe and Co require alkaline environment for passivation, Al shows stable passivity only in the near-neutral pH range). Therefore, selective release of these alloying elements is possible during passive dissolution. Many of the alloying elements in implant materials raise concerns on the biocompatibility; namely Al, V, Co, Ni, and Cr. For Cr it is important to note that the influence of metal release on the biological reactions will strongly depend on the oxidation state, Cr3 þ or Cr6 þ. The hexavalent Cr species are carcinogenic in nature and therefore of large concern. Oxidation of the Cr2O3 passive film into soluble hexavalent Cr species leads to transpassive dissolution of stainless steels and Co– Cr–Mo alloys, and therefore to strongly increased metal release. The speciation of released Cr will depend on the pre-vailing redox conditions (and other environmental factors). For hexavalent Cr release relatively high oxidative conditions that may not reflect the situation encountered in typical biomedical applications are required. Ti alloys represent the most corrosion resistant alloys used in the biomedical applications, due to the very high stability of the TiO2 passive film under many chemical and electrochemical conditions. Al-containing Ti alloys, however, not only bear the risk of Al release, but moreover the stability of the passive film is somewhat compromised as compared with pure Ti. Strongly enhanced selective Al release has been observed from these alloys under simulated inflammation conditions (slightly acidified solution and in presence of H2O2). In dental applications, the influence of fluoride on passivity of Ti alloys needs to be considered. Due to the concerns on Al release from the widely used Ti–6Al–4V alloy, new Al-free Ti alloys with promising corrosion (and mechanical) properties have been developed. The susceptibility of different passive alloys to chloride-induced pitting corrosion strongly varies, surgical stainless steels showing the lowest resistance against pitting corrosion among the alloys discussed here. Fig. 1 shows a comparison of the electrochemical behavior as determined by potentiodynamic polarization curves in Hank’s simulated physiological solution for stainless steel grade AISI 316L, Co–28Cr–6Mo, and two Ti alloys. All alloys are spontaneously passive in this electrolyte, but show a different behavior upon anodic polarization. The breakdown of passivity (increased anodic current densities) of the stainless steel is due to pitting corrosion. The pitting potentials typically determined in the laboratory for surgical-grade stainless steels are in a range that can be relevant considering the typical redox conditions that can be encountered in vivo. It should be mentioned
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Fig. 1 Potentiodynamic polarization curves for stainless steel, Co–28Cr–6Mo, Ti–6Al–4V, and Ti–6Al–7Nb alloys measured in Hank’s simulated physiological solution. Scan rate 1 mV/s. Ecorr: corrosion potential, Eb: breakdown potential. Reprinted from Milosev, I. Metallic Materials in Biomedical Applications: Laboratory and Clinical Studies. Pure Appl. Chem. 2011, 83, 309–324, with permission from IUPAC.
that the pitting corrosion resistance of stainless steels strongly depends on the exact alloy composition; higher Cr and Mo contents increase the resistance, whereas a higher impurity content (mainly sulfur) strongly decreases the resistance to pitting. For the Co–28Cr–6Mo alloy, the increase of the anodic current is not due to pitting corrosion but breakdown of passivity is due to oxidation of the Cr2O3-rich passive film into hexavalent Cr species leading to transpassive dissolution. Both Ti alloys show a significantly higher resistance to passivity breakdown in chloride-containing solutions. For the Ti–Al alloys, pitting potentials are well above 1 V and for pure Ti even z10 V. Therefore, pitting corrosion will not take place for Ti alloys under the conditions in the human body. Stainless steels are highly susceptible to crevice corrosion as compared with other implant alloys. The better crevice corrosion resistance of Co–Cr–Mo and Ti alloys is mainly related to the high stability of their passive films even in aggressive crevice electrolytes. Hence, in view of general passive behavior, chloride-induced pitting corrosion, and possible transpassive dissolution, the Ti alloys perform better than stainless steels and Co–Cr–Mo alloys. However, the Ti–Al–V alloys are more susceptible to fretting corrosion. Repassivation behavior determines degradation and metal ion release during cyclic activation/repassivation events, such as taken place under fretting conditions. There are only few studies directly comparing the repassivation behavior of different implant alloys in biological solutions. In a study of repassivation time after abrasion of the surface, stainless steel showed the longest, Co– 28Cr–6Mo alloy medium, and Ti–6Al–4V alloy shortest repassivation time. Moreover, metal release during cyclic activation/repassivation of alloy surface is mostly not stoichiometric; for instance for Co–28Cr–6Mo alloys a strong selective release of Co has been observed. As pure Co does not passivate in simulated body fluids, a repassivation reaction leads to formation of Cr2O3 (the new passive film) and soluble Co cations; hence to selective Co dissolution. Concerning tribocorrosion and fretting corrosion, the experimental approach used for the investigations is of utmost importance, as the biomechanical situation and therefore the degradation behavior depends on the design of experiment (and in the real case, of design of the implant). Due to this complexity, a comparison of findings from different type of bio-tribocorrosion experiments is not straightforward.
Corrosion of Biodegradable Metals For temporary implantsdthat is for biomedical devices that are only needed in the human body for the time of healingdcorrosion in body fluids is required. Such applications could be cardiovascular stents or bone-fracture healing devices. To prevent harmful long-term reactions due to a permanent presence of foreign material in the body, when it is no longer required, such devices could be fabricated from biodegradable metals that gradually dissolve in the body. Therefore, low-corrosion resistant metals and alloys are here in the focus of interest. At the same time, the metals should be nontoxic, hence mainly Fe and Mg base materials have been considered for such applications. The desired corrosion or biodegradation rate depends on the targeted application. In general, corrosion of Mg base alloys is significantly faster than corrosion of Fe base materials. For both metals, the high chloride concentration in body fluids assists active
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dissolution. However, the corrosion behavior is strongly influenced by the specific chemistry of biological fluids and significantly differs from corrosion in simple saline solutions. For instance, the presence of phosphates and carbonates in body fluids changes the composition of corrosion products, and partially protective corrosion product layers that strongly slow down dissolution may form on the surface. For both Mg and Fe, phosphates are known to be inhibitive species. Moreover, as the corrosion behavior of both Mg and Fe is strongly pH-dependent, buffering of the body solutions is noteworthy. This is especially important for Mg, as in nonbuffered solutions the pH value in the surroundings of dissolving Mg strongly increases; hence in nonbuffered solutions, corrosion leads to an alkaline shift of the pH value toward the region of Mg passivity. In buffered body fluids, the alkalization is less strong, and therefore does not help in passivating the Mg surface. It should be mentioned that true passivation would not be desired in view of biodegradable implants, as the implant should show a time-controlled dissolution rate optimized for the specific application, and in best case the materials should be completely bioabsorbable. However, formation of insoluble corrosion products may lead to conversion of the metal into compounds instead of complete dissolutiondthe effect of such corrosion products on the biocompatibility must be considered. Also, depending on the size of an implant, full metal degradation can take place or notdthis, as initial corrosion rate is typically fast, but the corrosion rate strongly slows down with time by the formation of protective corrosion product layers as described earlier. As concerns the biocompatibility, one needs to consider not only the effects of dissolved metal cations but also other products of the corrosion reactions. Due to the low standard potential of Mg, the cathodic reaction is reduction of water and hence formation of hydrogen gas takes place. One can easily understand that too strong hydrogen evolution could lead to accumulation of gas and be detrimental for the biological performance. Another concern is pH increase in the vicinity of corroding Mg surfaces, due to poor hydrolysis of Mg cations. Alkalization compromises biological equilibria and can lead to cell death. Both H2 gas accumulation and pH increase depend on the dissolution rate and on the mass transfer conditions around the implant. Another drawback is that for most Mg alloys typically very nonuniform corrosion takes place; this is often related to micro-galvanic coupling of the different phases of the alloy, as most alloying elements and impurities are more noble than Mg and hence act as local cathodes on the surface. The localized nature of the attack on one hand side makes a life-time prediction of an implant more difficult, but even more serious is the possibility of local loss of mechanical integrity of the device; especially as many biomedical implants are also under mechanical loading. In order to control the biodegradation rate, research is searching novel, optimized alloys as well as developing biocompatible coatings for biodegradable metals and alloys. Alloy development includes not only new chemistries, but also modification of the microstructure (for instance, ultra-fine grained materials have been studied as they may combine beneficial mechanical properties and good corrosion behavior). Surface modification and coatings have been very widely explored, and the different approaches include for instance anodization, Ca-phosphate coatings, and a variety of biodegradable polymer coatings. In addition to considering the life-time of biodegradable implants, the corrosion behavior as a function of time should be optimized to the healing process. This is schematically illustrated for biodegradable stents in Fig. 2: in an ideal situation the initial corrosion rate is very slow, enabling to maintain the mechanical integrity during early stages of healing (in addition to the mechanical integrity considerations, the healing and remodeling will not be interfered by strong corrosion that would lead to the negative side effects of Mg corrosion, namely H2 gas evolution and pH increase). After the remodeling process has been completed, the degradation should occur with a sufficient rate, but without causing accumulation of degradation products at the site of implantation. In view of coatings for the control of degradation rate, such a time-dependence of degradation could be possible by multilayered coatings.
Fig. 2 Schematic illustration of an ideal compromise between mechanical integrity and degradation of a biodegradable stent. Reprinted from Hermawan, H.; Dube, D. M.; Mantovani, D. Developments in Metallic Biodegradable Stents. Acta Biomater. 2010, 6, 1693–1697, with permission from Elsevier.
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Concluding Remarks The complex and dynamic nature of biocorrosion (i.e., corrosion of metals in the human body) is difficult to simulate in the laboratory. Therefore, research in the field ranges from highly simplified (but well controlled) laboratory experiments to retrieval studies. Electrochemical characterization of the corrosion behavior of metals and alloys has been carried out in simulated body fluids mimicking the inorganic composition of blood, but also in the presence of various organic components of biological fluids, such as amino acids, proteins, and cells. In spite of a clear demonstration that the organic species play a role in corrosion and metal release, detailed mechanistic understanding of the interactions between the components of the biological environment and the corroding metal surfaces needs to be further developed. For devices under mechanical loads, not only the chemical and biological environment determines the corrosion behavior, but moreover biomechanics of the system must be considered (i.e., mechanical loads and design/geometry of the device determine the tribological action). Correlation between in vitro corrosion studies with the in vivo behavior is important. For this, expertise in materials science, biology, and medicine needs to be combined.
See also: Corrosion of Titanium and Titanium Alloys; Functional Self-Healing Coatings: A New Trend in Corrosion Protection by Organic Coatings; Kinetics of Anodic Oxidation of Aluminum and Titanium: Formation of Porous Alumina and Titanium Oxide Nanotube Layers.
Further Reading Elias, N., Ed. Degradation of Implants Materials, Springer: New York, 2012. Gilbert, J. L. In Electrochemical Behavior of Metals in the Biological Milieu; Ducheyne, P., Ed.; Comprehensive Biomaterials, Vol. 1; Elsevier: Amsterdam, 2011; pp 21–48. Milosev, I. Metallic Materials in Biomedical Applications: Laboratory and Clinical Studies. Pure Appl. Chem. 2011, 83, 309–324. Landolt, D., Michler, S., Eds. Tribocorrosion of Passive Metals and Coatings, Woodhead Publishing Limited: Oxford, 2011. Yan, Y., Ed. Bio-Tribocorrosion in Biomaterials and Medical Implants, Woodhead Publishing Limited: Oxford, 2013. Yan, Y.; Dowson, D.; Neville, A. In-Situ Electrochemical Study of Interaction of Tribology and Corrosion in Artifical Hip Prosthesis Simulators. J. Mech. Behav. Biomed. Mater. 2013, 18, 191–199. Special Issue: Biomaterials Corrosion (2003), in Corrosion reviews, Vol. 21, Nos. 2–3. Hiromoto, S.; Hanawa, T. Corrosion of Implant Metals in Presence of Cells. Corros. Rev. 2016, 24, 323–352. Hedberg, Y.; Odnevall-Wallinder, I. Metal Release from Stainless Steel in Biological Environments: A Review. Biointerphases 2016, 11, 1–17. Staiger, M. P.; Pietak, A. M.; Huadmai, J.; Dias, G. Magnesium and its Alloys as Orthopedic Biomaterials: A Review. Biomaterials 2006, 27, 1728–1734. Zheng, Y. F.; Gu, X. N.; Witter, F. Biodegradable Metals. Mater. Sci. Eng. R 2014, 77, 1–34. Walker, J.; Shadanbaz, S.; Woodfield, T. B. F.; Staiger, M. P.; Dias, G. J. Magnesium Biomaterials for Orthopaedic Application: A Review From a Biological Perspective. J. Biomed. Mater. Res. B 2014, 201B, 1316–1331. Hermawan, H.; Dube, D. M.; Mantovani, D. Developments in Metallic Biodegradable Stents. Acta Biomater. 2010, 6, 1693–1697.
Introduction Human Body as a Corrosive Environment Possible Corrosion Modes in Biological Environments Corrosion Behavior of Selected Passive Alloys Used in Biomedical Applications Corrosion of Biodegradable Metals Concluding Remarks Further Reading
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Introduction Corrosion behavior of metallic materials used in biomedical applications (e.g., implants) is of high importance for biocompatibility, and in general for successful and safe applications. Strong degradation of a permanent implant would lead to failure by loss of the device integrity; for high-corrosion resistant metals and alloys that are typically used in such applications, strong active dissolution is, however, not expected. Nevertheless, specific corrosion modes are possible for passive metals and alloys. In addition, even small amount of corrosion, as takes place upon passive dissolution, can be detrimental, if the corrosion products are toxic or lead to adverse biological reactions. As the biological response to metal release depends on speciation and concentration of the released species, detailed information on the corrosion mode and rate is required. For biodegradable, temporary implants that are intended to corrode in the host body, it is even more self-evident that the nature of the corrosion products must be considered in view of their biological interactions. Corrosion behavior of metals and alloys is similar in the human body as in other aqueous solutions with a corresponding electrolyte composition. Hence, passive metals and alloys, such as Ti and its alloys, typically remain in the passive state also in the biological environment. Similarly, actively corroding materials, such as Mg and its alloys, would be expected to easily degrade in the body fluids. However, the specific chemistry of body fluids, as well as the presence of biomolecules and cells, can have very strong material-dependent influences on the corrosion behavior and metal release. Moreover, implant surfaces can be exposed to additional loads such as wear or mechanical stresses due to the function and design of the device. Therefore, not only purely chemical degradation but in addition mechano-chemical corrosion processes need to be considered. It should be emphasized that corrosion behavior of materials is always a system behavior, hence depending on the combination of the material (composition and microstructure), the environment (e.g., chemistry, temperature, hydrodynamic conditions), and the construction.
Human Body as a Corrosive Environment For metallic implants used in the body for healing, the first contact with the human body upon surgery is with blood. Therefore, simulated body fluids that are used to study the corrosion behavior in the laboratory often mimic the inorganic composition of blood. For this, the corrosive environment can be described as a 0.9% NaCl solution, with small amounts of other inorganic species, such as Ca2 þ, PO4 3 , and HCO3 , buffered to pH 7.4, at a temperature of c. 37 C. Chloride ions are aggressive species and can lead to breakdown of passivity of many metals and alloys, whereas phosphates and carbonates can lead to formation of (partially) protective corrosion product layers. Precipitation of Ca-phosphates from the body fluids can take place on metal and alloy surfaces, and this can further influence the electrochemical reactions. The body temperature of 37 C can accelerate corrosion reactions as compared with the room temperature. Even though body fluids are buffered by the CO2/HCO3 system to pH 7.4, local variations of the pH value can occur; upon inflammatory conditions and directly following surgery, lower pH values are expected. Also, if the implant design leads to formation of crevices, acidification inside the crevice electrolyte due to metal cation hydrolysis reactions can take place. Depending on the targeted biomedical application or on the type of contact of metals and alloys with the human body, different types of exposure scenarios exist; therefore the study of the corrosion behavior should consider the specific chemistry and biology for each case. In dermal contact, in lung or stomach environment, or in dental applications, the material surface is exposed to different chemical, biological, and mechanical interactions. Moreover, mass transfer conditions that can significantly influence corrosion reactions depend on the site of implantation. The redox potential of the biological environment depends on one hand side on the oxygen content (that also is locationdependent) and on the other side on the presence of a variety of other types of redox species; some of which can stem from biological reactions. Especially noteworthy are reactive oxygen species, such as hydroxide radicals and hydrogen peroxidedthese can have effects both on the biological system and on metal surface reactions.
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In addition to the variety of inorganic components of body fluids, presence of organic componentsdbiomolecules, proteins, cells, or bacteriadcan further influence corrosion reactions. All these species show a variety of interactions with the metal and alloy surfaces (and with each other), making the corrosion scenario highly complex. Moreover, they lead to a strongly time-dependent, dynamic system behavior, in that processes such as cell adhesion on surfaces lead to a number of possible follow-up reactions of the living system. Influence of proteins on corrosion has been frequently investigated, however the findings are not quite straightforward in view of the protein influence on corrosion rates. On one hand side this can be due to different experimental approaches, for instance in some studies single protein addition to simulated body fluids is used, while in other studies serum is used as a source of proteins. Moreover, it should be considered that proteins have multiple effects on corrosion and metal release. Proteins can adsorb on metal surfaces and such protein adsorption layers can induce corrosion protection by blocking the active sites of the surface. Protein adsorption depends on the nature of metal surface and of the specific protein, as protein adsorption is influenced by surface energy and surface charges. On the other hand, proteins can bind metal cations, and this could enhance metal dissolution reactions. This effect is also metal/protein-specific. In a tribocorrosion scenario (see what follows), biofilms forming on the surface are influenced by proteins, and may act as lubricants reducing friction. Cells can influence the biomaterial surface reactions by a variety of operative mechanisms. Macrophage cells are central to wound healing and inflammatory tissue reactions, and therefore in the early stage after surgery they accumulate around the foreign materials. Activated macrophages generate reactive oxygen species, and for passive metals and alloys these species can change the thickness and/or composition of passive films. For Ti-based materials it should be noted that H2O2 possibly generated by cells leads to enhanced solubility of the TiO2 passive film. For actively corroding, biodegradable metals (for instance, Mg alloys), reactive oxygen species generated by activated macrophages have been demonstrated to significantly increase the corrosion rate. Other cells, such as fibroblasts and osteoblasts, can adhere to the material surface by contacting cell-adhesive protein adsorption layers. The cells generate an extracellular matrix, containing polysaccharides, water, fibrous proteins, and cell-adhesive proteins. These cell adhesion layers and the extracellular matrix can on one hand side influence diffusion reactions to and from the metal surface, thereby influencing kinetics of electrochemical reactions. The electrolyte chemistry on the metal surface below cell layers can change, due to the diffusion hindrance as well as by cell metabolism products. For instance, pH decrease has been reported to take place near fibroblasts cultured on stainless steel, and this could locally accelerate metal dissolution (“cell-induced crevice corrosion”). On the other side, the cell adhesion layers can inhibit corrosion, by simple blocking of the active surface. As a direct probing of surface reactions in presence of cells on the surface is challenging, there are still many open questions on the mechanisms of cell-influence on metal corrosion (also considering that different metallic materials will show different response to the biological environment). In the real application (in vivo corrosion scenario), the situation is highly dynamic. Following protein adsorption and cell adhesion on the biomaterial surface, complex follow-up reactions take place, and these reactions depend on the material surface as well as on the location of implantation. The metal surface corrosion reactions will adapt to the changing environment, but there clearly is a lack of detailed understanding on the spatial and temporal variations of the implant surface/biology interactions. For instance, blood flow on the implant surface can be expected to influence any diffusion-controlled electrochemical reaction. However, the hydrodynamic conditions in the vicinity of an implant surface would strongly depend on the location of implantation and can change with time (e.g., due to tissue coverage of the implant with time as compared with a bare metal surface exposed to blood flow). Not only the anatomical location of an implant can have a significant influence on the corrosion behavior, but moreover the surgical procedure.
Possible Corrosion Modes in Biological Environments Corrosion of metals and alloys is determined by the thermodynamic and kinetic factors of the electrochemical reactions driving corrosion. For permanent implants, high-corrosion resistant materials that are passive under a large range of environmental conditions are employed. These materials include Ti and its alloys, CoCrMo alloys, and specific grades of stainless steels. In the next section, the corrosion behavior of these three material groups in a biological environment is shortly summarized. Another section is dedicated to considerations of the corrosion behavior of biodegradable metals (mainly Mg- and Fe-based alloys). First, general types of corrosion that can take place in the biomedical applications are introduced. In the case of stable passivity with no localized corrosion occurring, passive dissolution rates are orders of magnitude lower than dissolution in the active state. However, even such small passive dissolution rates may not be negligible in all cases, especially as concerns the composition of the passive alloys used; these alloys contain for instance Al, Co, Cr, and Ni; the release of these elements raises questions on the biocompatibility. Also, the protectiveness of passive films and hence the rate of passive dissolution depend on the alloy and the environment. In the biomedical applications, proteins and cell metabolism products can be involved in passive dissolution and metal release. Moreover, local acidification due to for instance inflammation or metal cation hydrolysis in sites of hindered mass transfer can increase passive dissolution rates. As the biological environment is a Cl-containing solution, it is straightforward to consider pitting corrosion as a possible degradation mode. Pitting corrosion leads to formation of cavities, surrounded by an intact passivated surface. For many materials, pit propagation is an autocatalytic process, related to formation of aggressive chemistry (acidification, accumulation of aggressive chloride, oxygen depletion) in the pit local environment. Moreover, an unfavorable area ratio of small anodes (pits) and a large cathode (the passive surface) accelerates local dissolution rates. Pitting corrosion leads to an increased metal release as compared with
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uniform passive dissolution, but moreover it can be dangerous for implants under mechanical load, as pits could act as stressconcentration sites and could possibly lead to local loss of the mechanical integrity of the device. Crevice corrosion occurs in sites on the metal surface where mass transfer is limited, leading to formation of locally aggressive electrolyte in narrow crevices or under deposits; similarly to what occurs in pits. Materials that are susceptible to pitting corrosion are also prone to crevice corrosion, but far less aggressive conditions are required to trigger crevice corrosion than to initiate pitting. In biomedical applications, crevices can occur in modular implants or at interfaces between implant surface and bone cement (or body tissues). Galvanic corrosion that is driven by a potential difference of dissimilar materials in contact and in a joint electrolyte is a possible corrosion risk in modular implants. The corrosion rate of the less noble metal is accelerated by the contact with a more noble metal, and the most critical case is given by a small area of a less noble metal forming a local anode in contact with a large area of a more noble material. For passive metals and alloys, the resistance given by the passive film, however, decreases the risk of galvanic corrosion. If any type of localized breakdown of passivity is possible, galvanic coupling with a more noble metal can significantly accelerate the propagation of localized corrosion of the less noble metal. Tribocorrosiondthat is, the simultaneous action of mechanical wear and corrosive attackdis a degradation mode for many biomedical implants where micromovements between surfaces take place and the implant surface is under mechanical loading. For biomedical implants fretting corrosion that takes place at the interface of two close surfaces subjected to small oscillatory slip and concurrent corrosive action is a highly relevant failure mode, for instance in hip and knee replacements. Here, information is required not only on the materials (including the tribological or fretting partners) and the environment, but moreover the biomechanical situation needs to be considereddhence the implant design and loading are important. As in most mechano-chemical corrosion modes, synergetic interactions between the mechanical wear and chemical corrosion can lead to drastically increased damage as compared with the sole action of wear or corrosion. For passive materials, fretting wear can easily mechanically destroy thin passive films. Therefore, during fretting corrosion continuous activation/repassivation cycles take place, and the repassivation ability of the materials is of critical importance for resulting damage and for metal release.
Corrosion Behavior of Selected Passive Alloys Used in Biomedical Applications For permanent implants, high-corrosion resistant alloys such as Ti-based materials (e.g., cp-Ti, Ti–6Al–4V, Ti–6Al–7Nb, Ni–Ti shape memory alloys, and others), Co–Cr–Mo alloys, and specific grades of stainless steels (Fe–Cr–Ni alloys) are used. All these materials in general show a very good passive behavior in a variety of environments. Considering specific localized corrosion modes of passive metals and alloys, significant differences in the resistance of the different alloy groups exist. Moreover, the exact chemical composition and microstructure of the specific alloy plays a significant role for the corrosion behavior. For instance, Co–Cr–Mo alloys exist in low-carbon and high-carbon variation, and they can be cast, forged, wrought, or produced by powder metallurgical routes; the processing and the carbon content lead to different carbide content and distribution in the alloydand this in its turn influences the corrosion behavior. Even in cases of no localized corrosion attack or tribocorrosion, passive dissolution leads to metal release and this can strongly vary for different passive materials under different conditions. Typical for high-corrosion resistant alloys is that they show a stable passivity under a large pH range. However, many of the biomedical alloys contain alloying elements that are not inherently stable under the conditions of biological solutions or in slightly acidified conditions (for instance, Fe and Co require alkaline environment for passivation, Al shows stable passivity only in the near-neutral pH range). Therefore, selective release of these alloying elements is possible during passive dissolution. Many of the alloying elements in implant materials raise concerns on the biocompatibility; namely Al, V, Co, Ni, and Cr. For Cr it is important to note that the influence of metal release on the biological reactions will strongly depend on the oxidation state, Cr3 þ or Cr6 þ. The hexavalent Cr species are carcinogenic in nature and therefore of large concern. Oxidation of the Cr2O3 passive film into soluble hexavalent Cr species leads to transpassive dissolution of stainless steels and Co– Cr–Mo alloys, and therefore to strongly increased metal release. The speciation of released Cr will depend on the pre-vailing redox conditions (and other environmental factors). For hexavalent Cr release relatively high oxidative conditions that may not reflect the situation encountered in typical biomedical applications are required. Ti alloys represent the most corrosion resistant alloys used in the biomedical applications, due to the very high stability of the TiO2 passive film under many chemical and electrochemical conditions. Al-containing Ti alloys, however, not only bear the risk of Al release, but moreover the stability of the passive film is somewhat compromised as compared with pure Ti. Strongly enhanced selective Al release has been observed from these alloys under simulated inflammation conditions (slightly acidified solution and in presence of H2O2). In dental applications, the influence of fluoride on passivity of Ti alloys needs to be considered. Due to the concerns on Al release from the widely used Ti–6Al–4V alloy, new Al-free Ti alloys with promising corrosion (and mechanical) properties have been developed. The susceptibility of different passive alloys to chloride-induced pitting corrosion strongly varies, surgical stainless steels showing the lowest resistance against pitting corrosion among the alloys discussed here. Fig. 1 shows a comparison of the electrochemical behavior as determined by potentiodynamic polarization curves in Hank’s simulated physiological solution for stainless steel grade AISI 316L, Co–28Cr–6Mo, and two Ti alloys. All alloys are spontaneously passive in this electrolyte, but show a different behavior upon anodic polarization. The breakdown of passivity (increased anodic current densities) of the stainless steel is due to pitting corrosion. The pitting potentials typically determined in the laboratory for surgical-grade stainless steels are in a range that can be relevant considering the typical redox conditions that can be encountered in vivo. It should be mentioned
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Fig. 1 Potentiodynamic polarization curves for stainless steel, Co–28Cr–6Mo, Ti–6Al–4V, and Ti–6Al–7Nb alloys measured in Hank’s simulated physiological solution. Scan rate 1 mV/s. Ecorr: corrosion potential, Eb: breakdown potential. Reprinted from Milosev, I. Metallic Materials in Biomedical Applications: Laboratory and Clinical Studies. Pure Appl. Chem. 2011, 83, 309–324, with permission from IUPAC.
that the pitting corrosion resistance of stainless steels strongly depends on the exact alloy composition; higher Cr and Mo contents increase the resistance, whereas a higher impurity content (mainly sulfur) strongly decreases the resistance to pitting. For the Co–28Cr–6Mo alloy, the increase of the anodic current is not due to pitting corrosion but breakdown of passivity is due to oxidation of the Cr2O3-rich passive film into hexavalent Cr species leading to transpassive dissolution. Both Ti alloys show a significantly higher resistance to passivity breakdown in chloride-containing solutions. For the Ti–Al alloys, pitting potentials are well above 1 V and for pure Ti even z10 V. Therefore, pitting corrosion will not take place for Ti alloys under the conditions in the human body. Stainless steels are highly susceptible to crevice corrosion as compared with other implant alloys. The better crevice corrosion resistance of Co–Cr–Mo and Ti alloys is mainly related to the high stability of their passive films even in aggressive crevice electrolytes. Hence, in view of general passive behavior, chloride-induced pitting corrosion, and possible transpassive dissolution, the Ti alloys perform better than stainless steels and Co–Cr–Mo alloys. However, the Ti–Al–V alloys are more susceptible to fretting corrosion. Repassivation behavior determines degradation and metal ion release during cyclic activation/repassivation events, such as taken place under fretting conditions. There are only few studies directly comparing the repassivation behavior of different implant alloys in biological solutions. In a study of repassivation time after abrasion of the surface, stainless steel showed the longest, Co– 28Cr–6Mo alloy medium, and Ti–6Al–4V alloy shortest repassivation time. Moreover, metal release during cyclic activation/repassivation of alloy surface is mostly not stoichiometric; for instance for Co–28Cr–6Mo alloys a strong selective release of Co has been observed. As pure Co does not passivate in simulated body fluids, a repassivation reaction leads to formation of Cr2O3 (the new passive film) and soluble Co cations; hence to selective Co dissolution. Concerning tribocorrosion and fretting corrosion, the experimental approach used for the investigations is of utmost importance, as the biomechanical situation and therefore the degradation behavior depends on the design of experiment (and in the real case, of design of the implant). Due to this complexity, a comparison of findings from different type of bio-tribocorrosion experiments is not straightforward.
Corrosion of Biodegradable Metals For temporary implantsdthat is for biomedical devices that are only needed in the human body for the time of healingdcorrosion in body fluids is required. Such applications could be cardiovascular stents or bone-fracture healing devices. To prevent harmful long-term reactions due to a permanent presence of foreign material in the body, when it is no longer required, such devices could be fabricated from biodegradable metals that gradually dissolve in the body. Therefore, low-corrosion resistant metals and alloys are here in the focus of interest. At the same time, the metals should be nontoxic, hence mainly Fe and Mg base materials have been considered for such applications. The desired corrosion or biodegradation rate depends on the targeted application. In general, corrosion of Mg base alloys is significantly faster than corrosion of Fe base materials. For both metals, the high chloride concentration in body fluids assists active
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dissolution. However, the corrosion behavior is strongly influenced by the specific chemistry of biological fluids and significantly differs from corrosion in simple saline solutions. For instance, the presence of phosphates and carbonates in body fluids changes the composition of corrosion products, and partially protective corrosion product layers that strongly slow down dissolution may form on the surface. For both Mg and Fe, phosphates are known to be inhibitive species. Moreover, as the corrosion behavior of both Mg and Fe is strongly pH-dependent, buffering of the body solutions is noteworthy. This is especially important for Mg, as in nonbuffered solutions the pH value in the surroundings of dissolving Mg strongly increases; hence in nonbuffered solutions, corrosion leads to an alkaline shift of the pH value toward the region of Mg passivity. In buffered body fluids, the alkalization is less strong, and therefore does not help in passivating the Mg surface. It should be mentioned that true passivation would not be desired in view of biodegradable implants, as the implant should show a time-controlled dissolution rate optimized for the specific application, and in best case the materials should be completely bioabsorbable. However, formation of insoluble corrosion products may lead to conversion of the metal into compounds instead of complete dissolutiondthe effect of such corrosion products on the biocompatibility must be considered. Also, depending on the size of an implant, full metal degradation can take place or notdthis, as initial corrosion rate is typically fast, but the corrosion rate strongly slows down with time by the formation of protective corrosion product layers as described earlier. As concerns the biocompatibility, one needs to consider not only the effects of dissolved metal cations but also other products of the corrosion reactions. Due to the low standard potential of Mg, the cathodic reaction is reduction of water and hence formation of hydrogen gas takes place. One can easily understand that too strong hydrogen evolution could lead to accumulation of gas and be detrimental for the biological performance. Another concern is pH increase in the vicinity of corroding Mg surfaces, due to poor hydrolysis of Mg cations. Alkalization compromises biological equilibria and can lead to cell death. Both H2 gas accumulation and pH increase depend on the dissolution rate and on the mass transfer conditions around the implant. Another drawback is that for most Mg alloys typically very nonuniform corrosion takes place; this is often related to micro-galvanic coupling of the different phases of the alloy, as most alloying elements and impurities are more noble than Mg and hence act as local cathodes on the surface. The localized nature of the attack on one hand side makes a life-time prediction of an implant more difficult, but even more serious is the possibility of local loss of mechanical integrity of the device; especially as many biomedical implants are also under mechanical loading. In order to control the biodegradation rate, research is searching novel, optimized alloys as well as developing biocompatible coatings for biodegradable metals and alloys. Alloy development includes not only new chemistries, but also modification of the microstructure (for instance, ultra-fine grained materials have been studied as they may combine beneficial mechanical properties and good corrosion behavior). Surface modification and coatings have been very widely explored, and the different approaches include for instance anodization, Ca-phosphate coatings, and a variety of biodegradable polymer coatings. In addition to considering the life-time of biodegradable implants, the corrosion behavior as a function of time should be optimized to the healing process. This is schematically illustrated for biodegradable stents in Fig. 2: in an ideal situation the initial corrosion rate is very slow, enabling to maintain the mechanical integrity during early stages of healing (in addition to the mechanical integrity considerations, the healing and remodeling will not be interfered by strong corrosion that would lead to the negative side effects of Mg corrosion, namely H2 gas evolution and pH increase). After the remodeling process has been completed, the degradation should occur with a sufficient rate, but without causing accumulation of degradation products at the site of implantation. In view of coatings for the control of degradation rate, such a time-dependence of degradation could be possible by multilayered coatings.
Fig. 2 Schematic illustration of an ideal compromise between mechanical integrity and degradation of a biodegradable stent. Reprinted from Hermawan, H.; Dube, D. M.; Mantovani, D. Developments in Metallic Biodegradable Stents. Acta Biomater. 2010, 6, 1693–1697, with permission from Elsevier.
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Concluding Remarks The complex and dynamic nature of biocorrosion (i.e., corrosion of metals in the human body) is difficult to simulate in the laboratory. Therefore, research in the field ranges from highly simplified (but well controlled) laboratory experiments to retrieval studies. Electrochemical characterization of the corrosion behavior of metals and alloys has been carried out in simulated body fluids mimicking the inorganic composition of blood, but also in the presence of various organic components of biological fluids, such as amino acids, proteins, and cells. In spite of a clear demonstration that the organic species play a role in corrosion and metal release, detailed mechanistic understanding of the interactions between the components of the biological environment and the corroding metal surfaces needs to be further developed. For devices under mechanical loads, not only the chemical and biological environment determines the corrosion behavior, but moreover biomechanics of the system must be considered (i.e., mechanical loads and design/geometry of the device determine the tribological action). Correlation between in vitro corrosion studies with the in vivo behavior is important. For this, expertise in materials science, biology, and medicine needs to be combined.
See also: Corrosion of Titanium and Titanium Alloys; Functional Self-Healing Coatings: A New Trend in Corrosion Protection by Organic Coatings; Kinetics of Anodic Oxidation of Aluminum and Titanium: Formation of Porous Alumina and Titanium Oxide Nanotube Layers.
Further Reading Elias, N., Ed. Degradation of Implants Materials, Springer: New York, 2012. Gilbert, J. L. In Electrochemical Behavior of Metals in the Biological Milieu; Ducheyne, P., Ed.; Comprehensive Biomaterials, Vol. 1; Elsevier: Amsterdam, 2011; pp 21–48. Milosev, I. Metallic Materials in Biomedical Applications: Laboratory and Clinical Studies. Pure Appl. Chem. 2011, 83, 309–324. Landolt, D., Michler, S., Eds. Tribocorrosion of Passive Metals and Coatings, Woodhead Publishing Limited: Oxford, 2011. Yan, Y., Ed. Bio-Tribocorrosion in Biomaterials and Medical Implants, Woodhead Publishing Limited: Oxford, 2013. Yan, Y.; Dowson, D.; Neville, A. In-Situ Electrochemical Study of Interaction of Tribology and Corrosion in Artifical Hip Prosthesis Simulators. J. Mech. Behav. Biomed. Mater. 2013, 18, 191–199. Special Issue: Biomaterials Corrosion (2003), in Corrosion reviews, Vol. 21, Nos. 2–3. Hiromoto, S.; Hanawa, T. Corrosion of Implant Metals in Presence of Cells. Corros. Rev. 2016, 24, 323–352. Hedberg, Y.; Odnevall-Wallinder, I. Metal Release from Stainless Steel in Biological Environments: A Review. Biointerphases 2016, 11, 1–17. Staiger, M. P.; Pietak, A. M.; Huadmai, J.; Dias, G. Magnesium and its Alloys as Orthopedic Biomaterials: A Review. Biomaterials 2006, 27, 1728–1734. Zheng, Y. F.; Gu, X. N.; Witter, F. Biodegradable Metals. Mater. Sci. Eng. R 2014, 77, 1–34. Walker, J.; Shadanbaz, S.; Woodfield, T. B. F.; Staiger, M. P.; Dias, G. J. Magnesium Biomaterials for Orthopaedic Application: A Review From a Biological Perspective. J. Biomed. Mater. Res. B 2014, 201B, 1316–1331. Hermawan, H.; Dube, D. M.; Mantovani, D. Developments in Metallic Biodegradable Stents. Acta Biomater. 2010, 6, 1693–1697.