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Biomedical magnesium alloys: material properties, surface modification and potential as biodegradable orthopedic implants

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  • Save American Journal of Biomedical En gineer in g 2012, 2(6): 218-240 DOI: 10.5923/j.ajbe.20120206.02 Biomedical Magnesium Alloys: A Review of Material Properties, Surface Modifications and Potential as a Biodegradable Orthopaedic Implant Gérrard Eddy Jai Poinern*, Sridevi Brundavanam, Derek Fawcett M urdoch Applied Nanotechnology Research Group, Department of Physics, Energy Studies and Nanotechnology, School of Engineering and Energy, M urdoch University, M urdoch, Western Australia, 6150, Australia Abstract Magnesium and magnesium based alloys are lightweight metallic materials that are extremely biocompatib le and have similar mechanical properties to natural bone. These materials have the potential to function as an osteoconductive and biodegradable substitute in load bearing applicat ions in the field of hard t issue engineering. However, the effects of corrosion and degradation in the physiological environ ment of the body has prevented their wide spread applicat ion to date. The aim o f this review is to examine the properties, chemical stability, degradation in situ and methods of improving the corrosion resistance of magnesium and its alloys for potential application in the orthopaedic field. To be an effective imp lant, the surface and sub-surface properties of the material needs to be carefully selected so that the degradation kinetics of the implant can be efficiently controlled. Several surface modification techniques are presented and their effectiveness in reducing the corrosion rate and methods of controlling the degradation period are discussed. Ideally, balancing the gradual loss of materia l and mechanical strength during degradation, with the increasing strength and stability of the newly forming bone tissue is the ultimate goal. If this goal can be achieved, then orthopaedic implants manufactured fro m magnesium based alloys have the potential to deliver successful clinica l outcomes without the need for revision surgery. Keywords Magnesium, Bio logical Co rrosion, Bioco mpatibility, Alloys, Surface Modification 1. Introduction The skeletal system of the human body is a complex three-dimensional structure that is important for two main reasons. The first aris es from the need to structurally support the many body organs and other related tissues. The second is the attachment of the numerous muscle groups that are needed for body movement and loco motion. The skeleton is constructed of two types of tissue, the first is a hard t issue called bone and the second is a softer tissue composed of cartilag inous materials. The adult hu man skeleton consists of 206 bones[1]; some provide protection to the internal organs, while others perform specialized functions such as transmitting sound vibrations in the inner ear. The bone matrix a lso provides a natural reservoir for ce lls and mineral ions that play an important role in maintain ing the biochemical balance within the body. For examp le, calciu m is an important element involved in muscular action and nerve conduction and its level in the body is closely monitored and regulated by a process called homeostasis[2]. *Corresponding author: (Gérrard Eddy Jai Poinern) Published online at Copyright © 2012 Scientific & Academic Publishing. All Rights Reserved Bone is a natural two phase organic-inorganic ceramic composite consisting of collagen fibrils with an embedded inorganic nano-crystalline co mponent. The primary organic phase of the bone matrix is Type I collagen, which is secreted by osteoblast cells to form self-assembled fibrils[3, 4]. The fibrils are bundled together and orientate themselves parallel to the load-bearing axis of the bone. The fibrils are typically 300 n m long, develop a 67 n m periodic pattern in which a 40 nm gap or hole is formed between the ends of the fibrils and the remain ing 27 n m overlaps the bundle behind[5]. This pattern creates discrete and discontinuous sites for the deposition of plate-like nanometre sized hydroxyapatite (HAP) crystals, which forms the second phase of the bone matrix. HAP is a mineral predominantly co mposed of calciu m phosphate which has the general chemical formu la of[Ca10 (OH) 2(PO4)6]. It is the main inorganic co mponent of bone and teeth, accounting for up to 65% by weight of cortical bone and in the case of teeth it accounts for 97 % by weight of dental enamel in mammalian hard tissue[6]. The discontinuous discrete sites limit the growth of the HAP crystals and force the crystals to grow with a specific crystalline orientation which is parallel to the load-bearing axis of the bone and collagen fibrils. The crystal p lates typically have a length of 50 n m, a width of around 25 n m and on average a thickness of 3 n m[7-10]. The HAP also has 219 American Journal of Biomedical Engineer ing 2012, 2(6): 218-240 trace amounts of potassium, manganese, sodium, ch loride, hydrogen phosphate, citrate and carbonate[11]. The final component of the bone matrix consists of the non-collagen organic proteins such as the phosphor-protein group which are believed to regulate the formation of the inorganic crystal phase by influencing the size, orientation and the depositional environ ment within the spaces between the collagen fibrils. The phosphor-protein group is also believed to be the source of calciu m and phosphate ions used in the formation of the minera l phase[12]. The organic phase gives bone its flexib ility, while the inorganic phase provides bone with its structural rig idity[13, 14]. The incorporation of organic and inorganic phases in the matrix g ives bone its unique mechanical properties such as toughness, strength, and stiffness. It is the combination of these properties that give bone and the skeletal system in general, its remarkable ability to withstand the various mechanical and structural loads encountered during normal and intense physical activity[15]. Ho wever, not all bone tissue in the body has the same properties and this is characterized by the presence of two types of bone. The first type consists of a hard outer layer of co mpact (cortical) t issue, while the second type forms the less dense and spongy (trabecular) tissue which fills the interior of the bone. This spongy interior contains marrow and the many blood vessels that supply nutrients and remove waste products from the bone tissues. Both the cortical bone and the trabecular bone are co mposed of the same organic and inorganic phases discussed above, but they differ in the amount of each phase present. The two bone types also differ in their respective porosities and in their structural arrange ment. The a mount of cortical and trabecular tissue found in bone is dependent on the external load being applied and the frequency of the load[16]. Despite its remarkable mechanical and structural properties bone can fracture fro m three main causes: 1) a fracture caused by sudden injury; 2) Fatigue or stress fractures resulting fro m repeated cyclic loads; and 3) Pathological fractures resulting fro m bone infections and tumours[17]. The surgical imp lantation of artificial biomaterials of specific size and shape is an effective solution in restoring the load bearing capacity and functionality of damaged bone tissue. The design and selection of bio materials is highly dependent on the specific med ical applicat ion. Therefore, it is imperative that new biomaterials being developed for load bearing orthopaedic implant applications should have excellent bioco mpatibility, comparable strength to natural bone, and produce no cytotoxicity effects[18, 19]. Metallic b io materials have been used since the early 1900s to replace damaged or diseased hard tissues. And as early as 1907, a magnesium alloy was used by Lambotte, to secure a bone fracture in the lo wer leg[20, 21]. Metallic imp lants are generally used in load bearing applications where their high mechanical strength and fracture toughness make’s them superior to ceramics, poly meric materials and poly mer / ceramic co mposites. Metallic implant materials currently used include stainless steel, cobalt-chro me alloys and titanium and its alloys. At present there are two major problems associated with using the metallic imp lants. The first involves the mis match between the mechanical properties of the metallic alloy and the surrounding natural bone tissue. The elastic modulus of both stainless steel and cobalt-chrome alloys is around ten times greater than that of bone, while a titaniu m alloy such as Ti-6Al-4V is around five times greater[22]. Bone tissue is constantly undergoing remodelling and modification in response to imposed stresses produced by normal everyday activities. The mechanical mis match between bone and different metallic implant materials results in a clin ical phenomenon known as stress shielding. The stress-shielding phenomenon occurs when the imp lant carries the bulk of the load and the surrounding bone tissue experiences a reduced loading stress. The reduced loading stress experience by the surrounding bone tissue ultimately leads to bone resorption[23, 24]. The second problem stems fro m mechanical wear and corrosion of the implant and results in the release of toxic metallic ions such as chromiu m, cobalt and nickel into the body. These harmful metallic ions solicit an inflammatory response from the body’s immune system and the surrounding tissues which reduces the biocompatibility of the imp lant[25, 26, and 27]. This is in total contrast to the corrosion products of magnesiu m (Mg) wh ich can be considered physiologically beneficial, with the adult body storing around 30 g of Mg in both muscle and bone tissue[28]. The importance of Mg to the body stems fro m the fact it is bivalent ion which is used to form apatite in the bone matrix and is also used in a number of metabolic processes within the body[29]. And recently, Robinson et al. reported the novel antibacterial properties of Mg metal against Escherichia coli, Pseudomonas aeruginosa and Staphylococcus aureus[30]. Mg is a lightweight, silvery-white metal that is relatively weak in its pure state and is generally used as an alloy in engineering applications. The density of Mg and its alloys are around 1.74 g/cm3 at 20oC, which is 1.6 and 4.5 times less dense than alumin iu m and steel, respectively[31]. Interestingly, the density of Mg is slightly less than natural bone which ranges fro m 1.8 to 2.1 g/cm3, while the elastic modulus of pure Mg is 45 GPa and human bone varies between 40 and 57 GPa[32, 33 and 34]. Because of this close similarity in the respective elastic moduli, using Mg in hard tissue engineering applications would greatly reduce the possibility of stress shielding and prevent bone resorption. Thus, Mg with its similar mechanical properties to natural bone, combined with its bioco mpatibility, makes it a promising material for the development of biodegradable orthopaedic imp lants[33, 35]. Poly meric materials have also been used in a number of tissue engineering applications since they have many attractive properties such as being lightweight, ductile in nature, biocompatible and biodegradable. Poly mers are materials with large mo lecules co mposed of small repeating structural units called mono mers. The mono mers are usually attached by covalent chemical bonds, with cross-linking taking place along the length of the molecule. It is the Gérrard Eddy Jai Poinern et al.: Biomedical M agnesium Alloys: A Review of M aterial Properties, 220 Surface M odifications and Potential as a Biodegradable Orthopaedic Implant amount of cross-lin king that gives the polymer its physiochemical propert ies. Many polymeric materials have been investigated since the body’s natural processes can easily handle the by-products resulting fro m their degradation, with the by-products being easily excreted in the urine. Natural poly mers such as polysaccharides[36-40], chitosan[41-46], hyaluronic based derivatives[47-50] and protein based materials such as fibrin gel[51, 52] and collagen[53-56], have all produced favourable outcomes in a number of tissue engineering applications. Similar studies using synthetic biopolymers composed of simp le high purity constituent monomers, fabricated under controllable formation conditions have produced a variety of tissue scaffolds and imp lants with tuneable and predictable physio-mechanical properties. These biopolymers also have low to xicity reactions with the body and their degradation rate can be easily controlled. Examp les of synthetic biodegradable polymers include Poly (lact ic acid), PLA[57-62], Po ly (L-lactic acid), PLLA [63-66], Po ly (lactic-co-glycolic acid), PLGA[67-70], Poly-capro lactone PCL[71-74] and Poly (g lycolic acid ), PGA [75-78]. These biopolymers are generally poly-α-hydro x esters that de-esterifies in the body as the polymer degrades to simp le metabolites[79]. Currently available b iodegradable sutures in clinical use are made fro m PLA and PGA. These synthetic biopolymers can also be made into different shapes and structures, such as pellets, rods, disks, films, and fibres as required for the specific applicat ion. So me of these applications include biodegradable sutures, bone and dental cements, bone grafting materials, plates, screws, pins, fixation devices and low load bearing applications in orthopaedics[80, 81]. However, even with their many attractive properties, biopoly mers have low mechanical strength when compared to ceramics and metals, which has resulted in them being used in soft tissue reconstruction and low-load bearing applications. The major advantage that Mg and its alloys have over biopolymers is its superior mechanical strength, which is typically double that of biopolyme rs . Ceramics are non-metallic, inorganic materials that are used in hard tissue engineering applications where they are collectively termed bioceramics. The important properties of bioceramics that make them highly desirable fo r b io medical applications are: 1) they are physically strong; 2) they are both chemically and thermally stable; 3) they exh ibit good wear resistance, and 4) they are durable in the body environment[82]. In addit ion, they are readily available, can be shaped to suit the application, they are bioco mpatible, hemoco mpatible, nontoxic, non-immunogenic and can be easily sterilised[83]. But unlike Mg and its alloys, bioceramics such as HAP, tend to be brittle, have low fracture toughness and are not as resilient. However, so me bioceramics have found application in hip jo ints, coatings on implants, maxillofacial reconstruction, bone tissue engineering and drug delivery devices[81, 84-86]. A composite material consists of two or more distinct parts or phases[85]. The major advantage of using a composite biomaterial stems fro m the fact a single-phase material may not have all the required properties for a particular applicat ion[86]. However, by co mbin ing one or more phases with diffe ring physical and chemica l properties it is possible to create a composite material with superior properties to those of the indiv idual co mponents. A good example of a natural co mposite is bone, which is a composed of Type 1 collagen and HAP. A typical man made examp le of a bio medical co mposite is a b ioactive coating of HA P or a bioactive glass deposited on to the surface of a titanium implant to pro mote bone attachment[87]. Co mposites, such as a 2-phase HAP-poly mer mixtu re have also been developed to create a bio material with similar p roperties to natural bone for hard tissue engineering applications[88]. Unfortunately, as mentioned above, biopolymers biodegrade with time and as a result, the load bearing capacity and fracture toughness of the implant will decline with time. When comparing the propert ies of Mg and its alloys with metals, polymers, ceramics and composites it can be shown that Mg and its alloys have many properties that are comparable, if not superior, see Table 1. However, despite its many advantages, Mg has the disadvantage of having a high corrosion rate in the body. And as a result, medical application of Mg based imp lants has been severely limited due to the electrolytic aqueous environment of the chloride rich body fluid (pH ranges between 7.4 and 7.6). Furthermore, there are two serious consequences of the rapid corrosion rate of Mg imp lants. The first is the rap id evolution of subcutaneous hydrogen gas bubbles which are produced at a rate too high fo r the surrounding tissues to handle[89, 90]. These bubbles usually appear within the first week after surgery and can be easily treated by drawing off the gas using a subcutaneous needle[91]. The second consequence of the high corrosion rate is the loss of mechanical integrity of the Mg implant being used in the load bearing application. The rapid decrease in mechanical properties resulting fro m exposure to the body fluid environ ment means that the implant is unable to provide the necessary support for the healing bone tissue. Generally, the imp lant would be expected to maintain its mechanical integrity between 12 to 18 weeks while the healing process takes place and then slowly degrade while natural bone tissues replace the implant[92]. This article reviews the biological performance, mechanical properties and potential applicat ion of biodegradable Mg based alloys for orthopaedic imp lants. The major disadvantage of using Mg in many engineering applications is its low corrosion resistance, especially in electrolytic, aqueous environments where it rapidly degrades. To slow the degradation rate in situ, factors influencing the corrosion rate such as alloying elements, surface modification and surface treatments are examined and discussed in the follo wing sections. 221 American Journal of Biomedical Engineer ing 2012, 2(6): 218-240 Table 1. Some mechanical propert ies of selected materials Tissue /Material Den sity (g cm-3) Com pressi ve Strength (MPa) Te nsile Strength (MPa) El a sti c Modulus (GPa) Natural Materials Arterial wall Collagen Collagen(Rat tail tendon) Cancellous bone Cortical bone 1.0 – 1.4 1.8 – 2.0 - 1.5 – 9.3 160 Trans. 240 Long. 0.50 -1.72 60 - 1.5 – 38 35 Trans. 283 Long. 0.001 1.0 3.75 – 11.5 0.01 – 1.57 5 - 23 Magnesium Alloys Pure magnesium AZ31 (Extruded) AZ91D (Die cast) Other metal alloys Cobalt-Chrome Alloys Stainless Steel Titanium Alloys Ceram ics Synthetic- Hydroxyapatite Alumina Ceramics (Al2O3 80% - 99%) Polymers Po lymethy lmet hacry late (P MMA) Polyethylene- terephthalate (P ET) 1.74 1.78 1.81 7.8 7.9 4.4 3.05 – 3.15 3.30 – 3.99 20 - 115 83 - 97 160 - 100 - 900 2000 – 4000 1.12 – 1.20 1.31 – 1.38 45 – 107 65 – 90 90 - 190 241 - 260 230 450 - 960 480 – 620 550 – 985 40 – 200 - 38 – 80 42 – 80 45 45 45 195 - 230 193 – 200 100 – 125 70 – 120 260 – 410 1.8 – 3.3 2.2 – 3.5 Note: Table compiled from references[122, 126, 213, 218, 219, 220 and 221] 2. Biological Corrosion of Magnesium 2.1. Corrosion Mechanism When unprotected chemically pure magnesiu m is exposed to humid at mospheric air it develops a thick dull gray amorphous layer co mposed of magnesium hydro xide[Mg (OH)2]. The o xidation rate of this protective o xide layer is typically around 0.01 mm/yr, while the o xidation rate in salt water is around 0.30 mm/yr[93]. In magnesium alloys, controlling the alloying chemistry and the overall microstructure of the alloy can significantly reduce the corrosion rate. Table 2. Corrosion rates for some magnesium alloys immersion in various media Ma teri al Pure Mg (99.95%) AZ31 AZ91 LAE442 In vitro corrosion rate (m -2.h -1) Hanks Solu tion Simulated Body Flui d 0.011 0.038 0.0065 - 0.0028 - - - In vivo corrosion rate (m -2.y r-1 ) 1.17 1.38 0.39 WE43 - 0.085 1.56 Note: Table compiled from references[92, 96, 126, 213, 214 and 215] For orthopaedic applications pure magnesium finds the human body a highly aggressive corrosive environment, see Table 2. The body flu ids are co mposed of water, dissolved oxygen, proteins and electro lytic ions such as chloride and hydroxide. In this environment, magnesium with a negative electrochemical potential o f -2.37 V, is very susceptible to corrosion and results in free ions migrat ing fro m the metal surface into the surrounding fluid environ ment. These ions can form chemical species, such as metal oxides, hydroxides, chlorides and other compounds. In thermodynamic terms, with the assumption that there is no barrier to oxidation of the metal surface, the reaction would be very rap id, evolv ing hydrogen gas and consuming the metal substrate surface. But in reality the electrochemical reaction results in the migration of ions fro m the metal surface into solution, which forms species that result in the formation of an o xide layer that adheres to the metal surface. The Mg (OH)2 layer fo rmed on the metal surface is slightly soluble and reacts with chorine ions to form highly soluble magnesiu m chlo ride and hydrogen gas[94, 95]. When the oxide layer fully covers and seals the metal surface, it forms a kinetic barrier or passive layer that physically limits or prevents further migration of ionic species across the metal oxide solution interface. The corrosion of Mg in an aqueous physiological environment can be exp ressed in the follo wing equations. The primary anodic reaction is expressed by the partial reaction presented in equation (1), at the same time the reduction of protons is exp ressed by the partial reaction occurring at the cathode (2). Gérrard Eddy Jai Poinern et al.: Biomedical M agnesium Alloys: A Review of M aterial Properties, 222 Surface M odifications and Potential as a Biodegradable Orthopaedic Implant Anodic reaction: Mg → Mg2- + 2e- (1) Cathodic reaction: 2H2O + 2e- → 2OH- + H2 (2) Another undesirable consequence of the corrosion process in Mg and its alloys is the formation of hydrogen gas. The rapid formation o f hydrogen gas resulting fro m the rich chorine environ ment produces subcutaneous gas bubbles, which generally appear within the first week after surgery and then disappear after 2 to 3 weeks[92]. During the init ial gas formation a subcutaneous needle can be used to draw off the gas. In 2007, Song postulated that a hydrogen evolution rate of 0.01 ml/cm2/day can be tolerated by the hu man body and does not constitute a serious threat[96]. If the Mg corrosion rate can be regulated so that the hydrogen evolution rate is below this value, then the implant will not create a gas threat. The reactions of solid Mg and the Mg (OH)2 layer with chorine ions in the aqueous environment are presented in equations (3) and (4). Solid Mg: Mg (s) + 2Cl-(aq) → Mg Cl2 + 2e- (3) Mg (OH)2 layer: Mg (OH)2 (s) + 2CL-(aq) → MgCl2 + 2OH- (4) The general reaction of the corrosion process is presented in equation (5). Mg (s) + 2H2O (l) → Mg (OH)2 (s) + H2 (g) (5) Corrosion in the aqueous environment of the body is not as straight forward as corrosion in the industrial environ ment. This is due to the corrosion rate being influenced by a variety of other factors such as: 1) the pH of body fluids; 2) variations in the pH value; 3) concentration of ions; 4) the presence of proteins and protein adsorption on the orthopaedic imp lant; and 5) the influence of the surrounding tissues[97, 98 and 99]. metals with similar electrochemical properties when designing implant devices. For example, the fixation screws used to attach an Mg plate during a bone reconstruction procedure should be made of a titanium (Ti) alloy, since Ti is the closest metal to Mg in the electrochemical series. Mg is the most reactive metal in the electrochemical series and will always be the anode in any corrosion reaction[101]. Therefore, selection of Ti a lloy fixation screws to secure the Mg plate ensures the lowest possible corrosion rate. Galvanic corrosion can also result fro m the presence of inter-metallic alloying elements or impurit ies present in the Mg matrix, see Figure 2. Figure 1. Galvanic corrosion between dissimilar metals 2.2. Types of Biol ogical Corrosion An important property of the o xide layer is its ability to remain fixed to the metal surface during a variety of mechanical loading situations. If the oxide layer ruptures during mechanical loading it will expose the pure Mg substrate to body fluids wh ich will result in further corrosion. The clinical repercussion of the corrosion process is the loss of mechanical strength and the ultimate failure of the imp lant. Typical forms of Mg corrosion encountered within the body environment are d iscussed in the following sections. 2.2.1. Galvanic Corrosion Galvanic co rrosion takes place between two dissimilar meta ls, each with a diffe rent electrochemica l potential, when they are in contact in the presence of an electrolyte which provides a pathway for the transfer of electrons. The less noble metal beco mes anodic, corrodes and produces a build up of corrosion by-products around the contact site. For example, if gold screws are used to attach an Mg plate to bone during reconstructive procedure, the resulting electrolytic effect of the body fluids (seru m or interstitial flu id) would preferentially attack the Mg plate; see Figure 1[100]. Therefore, it would be good design practice to use Figure 2. Galvanic corrosion resulting from inter-metallic elements 2.2.2. Granular Corrosion In many metal alloys, inter-granular corrosion can occur fro m the presence of impurities and inclusions which are deposited in the gra in boundary regions during solidification. Following solidification, nu merous galvanic reactions takes place between the metal matrix and the various impurities and inclusions. The ensuing corrosion rate at the various grain boundary regions exceeds that of the grains and results in an accelerated corrosion rate of the metal matrix. However, in the case of Mg alloys, inter-granular corrosion does not occur since the grains tend to be anodic, while their boundaries are cathodic in nature compared to the interior of the grains. The resulting grain boundary corrosion undercuts nearby grains which subsequently fall out of the matrix[102]. 223 American Journal of Biomedical Engineer ing 2012, 2(6): 218-240 2.2.3. Pitt ing Corrosion Pitting corrosion of Mg res ults from the rapid corrosion of small-localized areas which damage the protective surface oxide layer; see Figure 3. This form of corrosion is mo re serious than other forms of corrosion since the surface pits are difficu lt to see due to the presence of corrosion products. The pits are small, highly corrosive and continue to grow downwards, perforating the metal matrix[103]. After init ial nucleation at the surface, the presence of impurit ies in the Mg alloy microstructure often assists in further corrosion due to the galvanic differences in the materials[104, 105]. The environment within the pit is very aggressive, with chlorides species from the body flu ids and Mg+ ions from anodic dissolution greatly aggravating the situation. In addition, the mouth of the pit is s mall and prevents any dilution of the pit contents, which adds to the accelerating autocatalytic growth of the pit. During this process, electrons flowing fro m the pit make the surface surrounding the pit entrance become cathode-protected and the protective oxide layer is further weakened. Once pitting starts, an Mg component can be totally penetrated within a relatively short period of time and in the case of a biomed ical imp lant, its load bearing capacity would be greatly reduced to the point of failure. Another problem associated with pitting arises from localised increase in stress produced by the pit, which has the potential to form cracks[106]. The fo rmation of stress corrosion cracking and metal fatigue cracks in the pits can lead to failure of the imp lant during normal loading conditions. Figure 3. Pitting corrosion site at the surface of a magnesium component 2.2.4. Crevice Corrosion Crev ice corrosion is local contact corrosion that occurs between metal and metal/non metal co mponents. For example, if a magnesium plate is to be fixed in location by a set of screws with a small gap between the screw head and plate. The gap must have sufficient width to allo w the flo w of the body fluids through the gap and prevent any stagnant flow, see Figure 4. The stagnant flo w results in the build up of Mg+ ions, with an Mg+ ion concentration gradient soon set up between the entrance and the dead end of gap. The subsequent corrosion cell then starts to attack the metal components of the implant[107]. 2.2.5. Fretting Corrosion Fretting corrosion is the result of damage p roduced by metal co mponents in direct physical contact with each other in the presence of small vibratory surface motions. The micro -motions are produced by normal every day activities experienced by the human body which result in mechanical wear and metallic debris between the surfaces of metal components making up the biomedical imp lant[108]. During daily activ ity, the micro-motions remove the passive surface layer o f the metallic co mponents in direct contact, exposing fresh metal underneath. Then both the fresh metal surfaces and the metallic surface debris undergo o xidation. The surface debris has a further detrimental e ffect by acting as an abrasive agent during subsequent micro-motions. The corrosion rate is dependent on the applied load, the resulting fretting motion, the microstructure of the metal o r metal alloys used in the implant and solution chemistry in the region around the fretting zone[109, 110]. During the corrosion process metallic ions are produced which can form a wide range of organic-metallic co mp lexes and some metallic imp lants can release toxic metallic ions such as chromiu m, cobalt and nickel. These harmfu l metallic ions significantly reduce the biocompatibility of the imp lant and solicit a major inflammatory response from the body’s immune system[25, 26, and 27]. In the case of magnesium, metallic ions released during fretting, can be considered physiologically beneficial since these ions can be consumed or absorbed by the surrounding tissues, or be dissolved and readily excreted through the kidneys. Fretting corrosion is common in load bearing surfaces and is also capable initiat ing fatigue cracks in the fretting zone. Once formed the crack can propagate into the bulk of the metal mat rix and can lead to the failure of the imp lant. Fi gure 4. Crevice corrosion occurring bet ween magnesium component s in a body fluid environment 2.2.6. Erosion Corrosion Erosion corrosion occurs from the wearing away of the meta l surface or passive layer by the impact of wear debris in the body environment surrounding the implant. The metallic debris impacts on the surface of the imp lant, transferring energy into the region of the collision and plastically deforming the surface. During the deformation p rocess the surface becomes work harden to the point where the next impact exceeds the strain required fo r surface fracturing, Gérrard Eddy Jai Poinern et al.: Biomedical M agnesium Alloys: A Review of M aterial Properties, 224 Surface M odifications and Potential as a Biodegradable Orthopaedic Implant pitting or chip format ion. With the passage of time, the numerous impacts result in material loss fro m the metal surface[111]. For examp le, a femo ral head of a Cobalt-Chro miu m imp lant will have numerous scratches after 17 years of imp lantation in a patient[112]. A ll bio-metals used in implants inevitably corrode at some fin ite rate when immersed in the complex electrolytic environ ment of the body; even Ti alloys with the lowest corrosion rate produce corrosion debris. The debris can significantly influence the wear behaviour and erosion resistant properties of the imp lant. Ho wever, the effects of erosion may not be noticed until there is a significant loss of metal which ultimately leads to the clinical failure of the imp lant. 2.2.7. Stress Corrosion When an electrochemical potential is formed between stressed and unstressed regions of a metal implant under load, there is an increase in the chemical act ivity of the metal. This stress initiated corrosion mechanism effect ively increases the corrosion rate, usually by two to three times above the normal uniform rate. Th is usually results in the formation of s mall cracks that concentrate stress within the loaded implant, a mechanism know as stress corrosion cracking (SCC). Mg SCC can occur in any load stressed implant immersed in the dilute chloride environment of the body fluids. SCC init iated cracks grow rapidly and extend between the grains throughout the metal matrix[113, 114]. The progress of SCC is also influenced by the strain rate resulting fro m the implant loading cycles and the presence of hydrogen gas produced by the corrosion process[115, 116]. Current research suggests that chloride ions produce pitting in the protective surface layer, wh ich ultimately leads to a break down in the layer exposing the underlining Mg matrix to the electrolytic flu ids of the body environment. The resulting hydrogen diffuses into the stressed zone of the metal matrix ahead of the crack tip and allows the SCC crack to advance through the zone[117-119]. Fracture and failu re of the imp lant will occur when the SCC is belo w the normal operating stress of the implant. 2.2.8. Corrosion Fatigue Corrosion fatigue is the result of a material be ing e xposed to the combined effects of a cycling load and a corrosive environment[120]. In general, metal fatigue is the damage caused by the repeated loading and unloading of a metal component. The cyclic stress initiates the formation of microscopic cracks on the metal surface and also damages the protective passive layer. If there are any surface imperfections such as pores or pitting fro m corrosion, they become crack nucleat ion sites which can significantly speed up crack g rowth rates. In the body’s environment the cracks become localized electrochemical cells that promote further corrosion. Mg in particular is susceptible to corrosion fatigue due to the presence of chloride ions in the body fluids. Corrosion within the crack pro motes crack propagation and in co mbination with cyclic loading, the crack gro wth rate significantly increases. Eventually the loading stress exceeds the SCC threshold and the crack grows to a critical size resulting in the fracture of the metallic implant. The body environment can significantly reduce the fatigue life of Mg alloys, producing lo wer failure stresses and considerably shorter failure times. 3. Magnesium and its Alloys For bio medical applicat ions, the composition of the material being considered is a crucial factor since many of the elements that make up co mmercially available materials for industrial applications are extremely to xic to the human body. Therefore, in addition to meeting the mechanical properties needed for a particular bio medica l applicat ion, the material must also be bioco mpatible. Ideally, a biodegradable bio medical device should be co mposed of materials or alloys that are non toxic or carcinogenic. It would also be very advantageous if the material was composed of elements and minerals already present and compatible within the body such as magnesium, calciu m and zinc, see Tab le 3. Fu rthermore, the material should have a controllable d issolution rate or slow corrosion rate that permits the biomed ical device or imp lant to maintain its mechanical integrity until the surrounding tissues heal and are capable of carrying the load once again. After the healing process has taken place, the load bearing properties of the biomed ical implant are no longer required and the imp lant material should then be able to slowly dissolve away. Furthermore, the resultant by-products of the degradation process should be non-toxic; capable of being consumed or absorbed by the surrounding tissues, or being dissolved and readily excreted through the kidneys. Thus, for Mg and its alloys to be used as an effective biodegradable imp lant it is necessary to control their corrosion behaviour in the body flu id environ ment[121]. 3.1. The Influence of Alloyi ng Elements on Physical and Mechanical Properties There are three ma jor groups of Mg alloys: the first group consists of pure Mg; the second group consists of aluminiu m (Al) containing alloys such as AZ91, AZ31 and rare earth ele ments (RE) such as AE21; and the fina l group consists of the Al free alloys such as Mg-Ca, W E, MZ and WZ. The use of alloying elements such Al, Ca, Li, Mn, Y, Zn, Zr and RE in Mg alloys can significantly imp rove the physical and mechanical properties of the alloy by: 1) refin ing the grain structure; 2) imp roving the corrosion resistance; 3) form inter-metallic phases that can enhance the strength; and 4) assist in the manufacture and shaping of Mg alloys. Impurit ies common ly found in Mg alloys are Be, Cu, Fe and Ni and the levels of theses impurities are restricted to within specific limits during the production of the alloy, see Table 3. The range of acceptable levels for Be ranges fro m 2 to 4 ppm by weight, while Cu is (100-300 pp m), Fe (30-50 ppm) and Ni (20-50 pp m)[122]. Since both Be and Ni are carcinogenic, their use in b io medical applications should be 225 American Journal of Biomedical Engineer ing 2012, 2(6): 218-240 avoided as alloying elements. While elements such as Ca, Mn and Zn are essential trace elements for hu man life and RE elements exhib iting anti-carcinogenic properties should be the first choice for incorporation into an alloy. Studies by Song have suggested that very small quantities of RE elements and other alloying metals such as Zn and Manganese (Mn) could be tolerated in the human body and could also increase corrosion resistance[123]. Mn is added to many co mmercial alloys to improve corrosion resistance and reduce the harmful effects of impurit ies[124]. Mg alloys containing rare earth elements have also been found to increase the resistance to the flow of Mg2+ ions out of the Mg matrix via the Mg oxide layer[125]. During the degradation process the RE elements remained localised in the corrosion layer, which also contained high levels of both calcium and phosphorous. Also during this period a thin amorphous calciu m phosphate layer formed over the surface of the o xide layer[92, 126]. Recent studies by Witte et al. have investigated the degradation behaviour of Mg based alloy rods and polymer based control rods[poly (lactic acid)] in animal models. Rods of 15 mm d iameter and 20 mm long were inserted into the femu r of guinea p igs and the rods degradation profile monitored. The percentage compositions by weight of the Mg alloys investigated consisted of two alu min iu m-zinc alloys co mposed of 3% Al and1% Zn {AZ31} and 9% Al and 1% Zn {AZ91} with the balance of the alloys composed of pure Mg. In addition two RE alloys were studied, the first consisted of 4% yttriu m and a 3% rare earth mixtu re composed of neodymiu m, ceriu m and dysprosium {W E43} and the second composed of 4% lithiu m, 4%, alu miniu m and a 2% rare earth mixture o f ceriu m, lanthanum, neodymiu m and praseodymiu m {LA E442}[92, 127]. The imp lants were harvested at 6 and 18 weeks, with co mplete imp lant degradation occurring at 18 weeks. During this time radiographs were regularly taken, while a micro -tomography-based technique using X-ray synchrotron radiation was used to characterize the imp lant’s degradation process. All Mg based alloy imp lants were found to be beneficial and pro moted new in situ bone tissue format ion, while the polymer control rods produced a less significant effect. The LA E442 alloy had the greatest resistance to corrosion, wh ile the other alloys all had similar, but lower values of corrosion resistance and degraded at similar rates [92] . While Mg is potentially an ideal bioco mpatible imp lant material due to its non-toxicity to the human body, the safe long term use of an Mg based alloy needs to be carefully studied. Magnesium based alloys have also been used in vivo; for examp le an AZ91 alloy rods were imp lanted into the femur of a nu mber of rabbit models and the subsequent analysis revealed that after 3 months the implant had degraded and been replaced by new bone tissue[128, 129]. At the end of this degradation process most of the alloying ele ments such as Al would have been re leased into the bodies of the rabbits. The long term health effects on the rabbits are unknown, but in the case of the human body, the release of Al into the body will create undesirable health problems[130]. In hu mans, Al is a neurotoxicant and its long term accu mulation in brain t issues has been linked to neurological disorders such as Alzheimers disease, dementia and senile dementia[131]. In addition, the administration of RE elements such as cerium, praseodymiu m and yttrium has resulted in severe hepatotoxicity in rats[132]. Furthermore, using heavy metal elements as alloying components are also potentially to xic to the human body due to their ability to form stable co mplexes and disrupt the normal mo lecular functions of DNA, en zy mes and proteins[133]. Therefore, there is a definite requirement to carefully select alloying elements that are non-toxic to the hu man body, see Table 4. Non-toxic alloying elements such as Ca[134] and Zr[135] have the potential to significantly improve the corrosion resistance of the Mg alloy and reduce the degradation rate to make the Mg metal alloy a viab le implant material[33]. Table 3. Chemical analysis of alloying element s for a select ion of magnesium alloys Alloy Nominal element component (wt. %) Maximum values of trace element s (wt . %) Al Zn Mn Ca Li Nd Zr Y AZ31 3.5 1.4 0.3 - - - - - Fe (max 0.003), Cu (0.008), Si (1.2), Ni (0.001) and Be (5 – 15 ppm) AZ91 9.5 0.5 0.3 - - - - - Fe (max 0.004), Cu (0.025), Si (0.05), Ni(0.001) and Be (5 – 15 ppm) AM60 6.0 0.2 0.2 - - - Fe (max 0.004), Cu (0.008), Si (0.05), Ni(0.001) and Be (5 – 15 ppm) LAE442 4.0 4.0 2.0 Contains some heavy metal rare earth elements WE43 - - - - - 3.2 0.5 4.0 Note: Table compiled from references[92, 93, 216, 217 and 218] Contains some heavy metal rare earth elements Gérrard Eddy Jai Poinern et al.: Biomedical M agnesium Alloys: A Review of M aterial Properties, 226 Surface M odifications and Potential as a Biodegradable Orthopaedic Implant Table 4. Common alloying elements used in magnesium alloys Alloying Element Aluminium Mech an ical P ro pert ies Enhancement to Mg Matrix Rapidly diffuses through Mg mat rix, and act s as a passivating element and improves corrosion resistance. Improves die cast-ability P atho phy sio lo gy Blood serum level 2.1-4.8 µg/L Calcium Copper Manganese Lit h ium Rare earth Element s Zinc Adding to improve corrosion resistance in Mg-Ca alloys. Can increase strength of Mg casts, however, it also accelerat es corrosion rate when exposed to a NaCl medium. Adding to reduce the harmful effect s of impurities and improve corrosion resistance Improvement in corrosion resistance Improvement in corrosion resistance Improves yield st ress, Mg alloys containing Zn have an Elastic Modulus similar to bone. The presence of Zn can reduce hydrogen gas evolution during bio-corrosion. Blood serum level 0.919-0.993 mg/L. Levels controlled by Homeost at is of skeleton. Abundant mineral that is mainly stored in bones and teeth. Activator/stabilizer of enzymes. Involved in blood clotting Blood serum level 74-131 µmol/L Essentialtrace element Blood serum level <0.8 µg/L Essentialtrace element Influences cellular functions/immune system/blood clotting/bone growth. Influences met abolic cycle of lipids/amino acids and carbohydrates Blood serum level 2-4ng/g Used in drugs to treat psychiatric disorders Many rare earth element s have anticancerogenic properties and are used in the treatment of cancer. Blood serum level 12.4-17.4µmol/L Essentialtrace element Essentialto enzymes and immune system Note: Table compiled from references[122, 134, 213, 222, 223, 224, 225, 226 227 and 228] Toxicology Tends to diffuse out of Mg mat rix Neurotoxic (influences funct ion of the blood brain barrier) Linked to Alzheimer’s disease Accumulat es in amyloid fibres/brain plaques. Accumulat es in bone t issue/decreases osteclast viability Met abolic disorder of calcium levels result s in the format ion of excess calcium in the kidneys (stones). Excessive amounts of Cu have been linked to neuro-degenerative diseases. Can produce cellular cytotoxicity. Excessive amounts of Mn can produce neurological disorder. (manganism) Overdose causes central nervous centre disorders, lung dysfunctions, impaired kidney function. Accumulate in the liver and bone In high concentrations is neurotoxic and can hinder bone development. 4. Surface Modifications and Treatment Processes for Biomedical Mg Alloys The high degradation rate of Mg and Mg alloy implants in the human physiological environ ment would result in the reduction of mechanical integrity of the imp lant before the bone tissues had sufficient time to heal[26]. There are t wo methods of reducing the degradation rate; the first, which was discussed in Section 3, involved alloy ing Mg with biocompatible elements that can resist the corrosion process. The second method is d iscussed in this section and involves the surface modification of the imp lant, through a treat ment process that provides a resistive barrier against the body environment. An important factor that needs to be taken into account before any surface treatment is investigated is the healing or regenerative processes of bone and other associated body tissues. The healing process consists of three phases; inflammatory, reparat ive and remodelling. The in itial inflammatory phase usually lasts between 3 to 7 days and this is the natural response of the body’s immune system to the presence of the biomedical device or imp lant. The reparative phase usually takes 3 to 4 months, during which t ime integration of the imp lant with the new and regenerated tissues takes place. The final remodelling phase, which is the longest phase, can take fro m several months to years to complete[136]. Fo r Mg to be an effective bio-absorbable implant the degradation rate must be slow enough for the healing process to take place and the new tissues have sufficient time to provide their own structural support before the structural integrity of the imp lant is compro mised. The minimu m period for this to take place is at least 12 weeks[26]. Un fortunately, Mg alloys can completely degrade before the end of this timeframe and as a result there is a need to reduce the biodegradation rate. The bulk propert ies of Mg based alloys dictate its mechanical properties, but it is the surface properties that influence the interaction between the metal and the surrounding tissue environment of the body. As a consequence, surface 227 American Journal of Biomedical Engineer ing 2012, 2(6): 218-240 modifications and treatments can have a significant ro le to play in governing the degradation rate of the implant. To date, numerous surface modification techniques have been developed to change the surface characteristics of biomaterials. Many of these methods have been applied to modifying the surface p roperties of Mg bio-alloys. A brief overview of so me of these surface modificat ion processes are presented in the following the four sections. 4.1. Mechanical Modificati ons to Induce Surface and Subsurface Properties The surface structure of an imp lant is very important, since it is the init ial response of the surrounding tissues to the surface of the implant material that determines whether or not there is effective t issue-biomaterial integration. Studies of conventional types of permanent imp lant materials have shown that surface roughness can influence both cell morphology, cell growth and implant integration. In addition, modification of the surface topography by the physical placement of grooves, columns, pits and other depressions can influence cell orientation and attachment[137-139]. In the case of Ti a lloys, surface modifications such as grooves, surface sand blasting and acid etching has revealed that grooved surface features provide superior cell attachment and promote greater cell pro liferation than roughen surfaces[140]. For Mg alloys, the influence of different mechanical p rocessing operations during fabricat ion has the potential to great ly influence surface and subsurface properties[141, 142]. Mechanical processing techniques involve operations such as rolling, shot peening, and milling. In the case of milling, at low cutting speeds, the surface formed by honed cutting tools tends to produce a rougher surface than those of sharp cutting tools. Also, during milling and similar metal chip removing processes, the exact effect on the underlining sub-surface is not fully understood[143], while chip removal fro m the surface during machining can direct ly influence the surface topography[144]. Besides machining techniques for chip removal, the use of rolling operations can also generate high passive forces acting normal to the surface, which can induce work hardening of the sub-surface. During the rolling operation the sub-surface grain structure is changed by the compressive stresses induced and the resultant micro -topography of the surface is significantly changed[145]. A recent study by Denka et al. has revealed a significant reduction in the corrosion rate (a factor of 100 was achieved in corrosion studies) of an Mg-Ca alloy that was deep-rolled, co mpared to the same alloy that was mach ined[146]. The presence of residual co mpressive stresses after ro lling also has the advantage of reducing micro -crack formation fro m pre-existing crack nucleation points within the substrate. The suppression of crack formation is also an important factor in improving the fatigue life cycle of a material being considered fo r b io medical applications[146, 147]. The importance of surface and sub-surface treatments on Mg alloy implants was recently investigated by Von Der Hoh et al.[148]. In their study three surface machining treatments were applied to an Mg-Ca (0.8 % wt calciu m) alloy. The alloy was used to make three different geometric sample types. The first test sample was a machined 3 mm diameter s mooth cylinder, the second was like the first, except that it was sand blasted for 30 s using particles ranging in size fro m 300 to 400 µm and the final surface topography was a threaded cylinder. The smooth cylinders were machined with no fu rther surface treat ment, so they retained the micro-surface topography produced by the cutting tool. After 6 months of in vivo imp lantation in adult New Zealand white rabbits, the smooth cylinders revealed good integration with the surrounding tissues and also had the least structural loss. The sand blasted cylinders had the greatest material loss with the init ial cylindrical shape completely consumed, while the threaded cylinders ranged between these two extremes. The results indicated that the smoother mic ro-topographic surface features of the cylinders were suitable for resorbable Mg alloys, wh ile the test samples with the rougher surfaces promoted higher degradation rates. The results of this study clearly indicated that differences in surface roughness of the test samples could significantly influence the in vivo degradation rates. The study also highlighted the need for further investigation into the effects of different surface modifications on other biocompatible Mg alloys. 4.2. Physical and Chemical Modifications Fro m an engineering point of view, the most effective way to prevent corrosion is to coat the metal co mponent with a protective barrier that effect ively isolates the metal fro m the surrounding environment. To be effective against corrosion, the protective coating must be uniform, well adhered and free fro m any imperfections such as pits, scratches and cracks. The major problem with Mg, as mentioned earlier, is its chemical reactivity when exposed to air or an aqueous environmental wh ich results in the formation of an oxide/hydroxide layer over the metal surface. The presence of the o xide/hydroxide layer will have a detrimental effect on the ability of the coating to adhere to the metal surface and form a unifo rm p rotective layer. Therefore, surface cleaning and a suitable pre-treat ment o f the metal surface is a crucial factor in achiev ing an effective surface coating. 4.2.1. Physical Vapour Deposition (PVD) & Chemical Vapour Deposition (CVD) The PVD process involves the deposition of thin layers of metal and metal alloys fro m ato ms or mo lecules fro m the vapour phase onto a substrate surface. During the process a metal or metal alloy is heated in vacuum chamber until it evaporates and then the subsequent vapour condenses onto the cooler substrate. This process has been successfully used on a variety of metals, but in the case of Mg there are a number o f problems to overcome. For examp le, in most PVD processes the substrate temperature range is usually between

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