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Here is an essential text for cardiologists, heart surgeons, intensive care specialists and anyone interested in pacing. It is a comprehensive guide to contemporary devices used in the resynchronization of patients' heartbeats. The treatment of congestive heart failure by implanted biventricular pacemakers, or cardiac resynchronization, has revolutionized the practice of implanting pacemakers and defibrillators.

More cardiac resynchronization therapy devices than conventional pacemakers are now being implanted and the numbers are growing worldwide. The treatment of congestive heart failure by implanted biventricular pacemakers, or cardiac resynchronization, has revolutionized the prac Authors: Daniel Tarsy, Jerrold L.

Vitek, Philip A. Starr, Michael S. In the coming years the World Health Organization predicts that depression will rank just behind heart conditions in terms of the global disease burden. Yet, according to a provocative new book, mental health systems often reinforce the depressive disorders they aim to treat. In Depression and Globalization, Carl Walker analyzes the human cost of recent political and economic events as main contributors to the rise of depression, particularly in the U. Starting in the s, income and educational disparities, financial and job insecurity—by-products of multinational business—have grown in parallel with increasing feelings of hopelessness and isolation.

These sociopo Authors: Gennady, E. Volume 1. Zaikov, DSc, and A. Zaikov, DSc, A. Haghi, PhD, reviewers and Advisory Board members. Biofilm and Materials Science. Authors: Hideyuki, Kanematsu; Dana, M. Summary: This book explains the formation of biofilm on materials surfaces in an industrial setting. The authors describe new developments in understanding of biofilm formation, detection, and control from the viewpoint of materials science and engineering. The book details the range of issues caused by biofilm formation and the variety of affected industries. Authors: Haghi, A. Metallic biomaterials have been used since the early s to replace damaged or diseased hard tissues.

And as early as , a magnesium alloy was used by Lambotte, to secure a bone fracture in the lower leg[20, 21]. Metallic implant materials currently used include stainless steel, cobalt-chrome alloys and titanium and its alloys. At present there are two major problems associated with using the metallic implants. The first involves the mismatch 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 titanium 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 mismatch between bone and different metallic implant materials results in a clinical phenomenon known as stress shielding.

The stress-shielding phenomenon occurs when the implant 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 from mechanical wear and corrosion of the implant and results in the release of toxic metallic ions such as chromium, cobalt and nickel into the body.

This is in total contrast to the corrosion products of magnesium Mg which 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 from 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].


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And recently, Robinson et al. 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. Interestingly, the density of Mg is slightly less than natural bone which ranges from 1.

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 biocompatibility, makes it a promising material for the development of biodegradable orthopaedic implants[33, 35]. Polymeric 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.

Polymers are materials with large molecules composed of small repeating structural units called monomers. The monomers are usually attached by covalent chemical bonds, with cross-linking taking place along the length of the molecule. It is the amount of cross-linking that gives the polymer its physiochemical properties.

Natural polymers such as polysaccharides[], chitosan[], hyaluronic based derivatives[] and protein based materials such as fibrin gel[51, 52] and collagen[], have all produced favourable outcomes in a number of tissue engineering applications. Similar studies using synthetic biopolymers composed of simple high purity constituent monomers, fabricated under controllable formation conditions have produced a variety of tissue scaffolds and implants with tuneable and predictable physio-mechanical properties.

These biopolymers also have low toxicity reactions with the body and their degradation rate can be easily controlled. 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 application. Some 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, biopolymers 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 biopolymers.

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 for biomedical applications are: 1 they are physically strong; 2 they are both chemically and thermally stable; 3 they exhibit good wear resistance, and 4 they are durable in the body environment[82].

In addition, they are readily available, can be shaped to suit the application, they are biocompatible, hemocompatible, 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, some bioceramics have found application in hip joints, coatings on implants, maxillofacial reconstruction, bone tissue engineering and drug delivery devices[81, ]. A composite material consists of two or more distinct parts or phases[85].

The major advantage of using a composite biomaterial stems from the fact a single-phase material may not have all the required properties for a particular application[86]. However, by combining one or more phases with differing physical and chemical properties it is possible to create a composite material with superior properties to those of the individual components. A good example of a natural composite is bone, which is a composed of Type 1 collagen and HAP. A typical manmade example of a biomedical composite is a bioactive coating of HAP or a bioactive glass deposited on to the surface of a titanium implant to promote bone attachment[87].

Composites, such as a 2-phase HAP-polymer mixture have also been developed to create a biomaterial with similar properties 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 properties 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 implants has been severely limited due to the electrolytic aqueous environment of the chloride rich body fluid pH ranges between 7. Furthermore, there are two serious consequences of the rapid corrosion rate of Mg implants. The first is the rapid evolution of subcutaneous hydrogen gas bubbles which are produced at a rate too high for 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 from exposure to the body fluid environment means that the implant is unable to provide the necessary support for the healing bone tissue. Generally, the implant 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 application of biodegradable Mg based alloys for orthopaedic implants.


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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 following sections. Table 1. Biological Corrosion of Magnesium 2. Corrosion Mechanism When unprotected chemically pure magnesium is exposed to humid atmospheric air it develops a thick dull gray amorphous layer composed of magnesium hydroxide[Mg OH 2 ].

The oxidation rate of this protective oxide layer is typically around 0. In magnesium alloys, controlling the alloying chemistry and the overall microstructure of the alloy can significantly reduce the corrosion rate. Table 2. The body fluids are composed of water, dissolved oxygen, proteins and electrolytic ions such as chloride and hydroxide.

In this environment, magnesium with a negative electrochemical potential of 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 rapid, evolving hydrogen gas and consuming the metal substrate surface. But in reality the electrochemical reaction results in the migration of ions from the metal surface into solution, which forms species that result in the formation of an oxide layer that adheres to the metal surface.

The Mg OH 2 layer formed on the metal surface is slightly soluble and reacts with chorine ions to form highly soluble magnesium chloride 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 expressed in the following equations.

The primary anodic reaction is expressed by the partial reaction presented in equation 1 , at the same time the reduction of protons is expressed by the partial reaction occurring at the cathode 2. The rapid formation of hydrogen gas resulting from the rich chorine environment produces subcutaneous gas bubbles, which generally appear within the first week after surgery and then disappear after 2 to 3 weeks[92].

During the initial gas formation a subcutaneous needle can be used to draw off the gas. In , Song postulated that a hydrogen evolution rate of 0. 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: 3 Mg OH 2 layer: 4 The general reaction of the corrosion process is presented in equation 5. 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 implant; and 5 the influence of the surrounding tissues[97, 98 and 99].

Types of Biological Corrosion An important property of the oxide 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 which will result in further corrosion.

The clinical repercussion of the corrosion process is the loss of mechanical strength and the ultimate failure of the implant. Typical forms of Mg corrosion encountered within the body environment are discussed in the following sections. Galvanic Corrosion Galvanic corrosion takes place between two dissimilar metals, each with a different electrochemical 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 becomes 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 serum or interstitial fluid would preferentially attack the Mg plate; see Figure 1[]. Therefore, it would be good design practice to use 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[]. Therefore, selection of Ti alloy fixation screws to secure the Mg plate ensures the lowest possible corrosion rate. Galvanic corrosion can also result from the presence of inter-metallic alloying elements or impurities present in the Mg matrix, see Figure 2. Figure 1. Galvanic corrosion between dissimilar metals Figure 2. Galvanic corrosion resulting from inter-metallic elements 2.

Granular Corrosion In many metal alloys, inter-granular corrosion can occur from the presence of impurities and inclusions which are deposited in the grain boundary regions during solidification. Following solidification, numerous 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[]. Pitting Corrosion Pitting corrosion of Mg results from the rapid corrosion of small-localized areas which damage the protective surface oxide layer; see Figure 3. This form of corrosion is more serious than other forms of corrosion since the surface pits are difficult to see due to the presence of corrosion products.

The pits are small, highly corrosive and continue to grow downwards, perforating the metal matrix[]. After initial nucleation at the surface, the presence of impurities in the Mg alloy microstructure often assists in further corrosion due to the galvanic differences in the materials[, ]. In addition, the mouth of the pit is small and prevents any dilution of the pit contents, which adds to the accelerating autocatalytic growth of the pit. During this process, electrons flowing from 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 biomedical implant, 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[]. The formation of stress corrosion cracking and metal fatigue cracks in the pits can lead to failure of the implant during normal loading conditions.

Figure 3. Pitting corrosion site at the surface of a magnesium component Figure 4. Crevice corrosion occurring between magnesium components in a body fluid environment 2. 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 allow the flow of the body fluids through the gap and prevent any stagnant flow, see Figure 4. The subsequent corrosion cell then starts to attack the metal components of the implant[].

Fretting Corrosion Fretting corrosion is the result of damage produced by metal components 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 implant[].

During daily activity, the micro-motions remove the passive surface layer of the metallic components in direct contact, exposing fresh metal underneath. Then both the fresh metal surfaces and the metallic surface debris undergo oxidation.

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The surface debris has a further detrimental effect 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 or metal alloys used in the implant and solution chemistry in the region around the fretting zone[, ]. During the corrosion process metallic ions are produced which can form a wide range of organic-metallic complexes and some metallic implants can release toxic metallic ions such as chromium, cobalt and nickel.

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 initiating fatigue cracks in the fretting zone. Once formed the crack can propagate into the bulk of the metal matrix and can lead to the failure of the implant. Erosion Corrosion Erosion corrosion occurs from the wearing away of the metal 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 implant, transferring energy into the region of the collision and plastically deforming the surface. During the deformation process the surface becomes work harden to the point where the next impact exceeds the strain required for surface fracturing, pitting or chip formation. With the passage of time, the numerous impacts result in material loss from the metal surface[]. For example, a femoral head of a Cobalt-Chromium implant will have numerous scratches after 17 years of implantation in a patient[].

All bio-metals used in implants inevitably corrode at some finite rate when immersed in the complex electrolytic environment 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 implant. However, 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 implant.

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 activity of the metal. This stress initiated corrosion mechanism effectively increases the corrosion rate, usually by two to three times above the normal uniform rate. This usually results in the formation of small 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 initiated cracks grow rapidly and extend between the grains throughout the metal matrix[, ]. The progress of SCC is also influenced by the strain rate resulting from the implant loading cycles and the presence of hydrogen gas produced by the corrosion process[, ]. Current research suggests that chloride ions produce pitting in the protective surface layer, which ultimately leads to a break down in the layer exposing the underlining Mg matrix to the electrolytic fluids 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[]. Fracture and failure of the implant will occur when the SCC is below the normal operating stress of the implant. Corrosion Fatigue Corrosion fatigue is the result of a material being exposed to the combined effects of a cycling load and a corrosive environment[].

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 from corrosion, they become crack nucleation sites which can significantly speed up crack growth rates. Mg in particular is susceptible to corrosion fatigue due to the presence of chloride ions in the body fluids.

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Corrosion within the crack promotes crack propagation and in combination with cyclic loading, the crack growth 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 lower failure stresses and considerably shorter failure times.

Magnesium and its Alloys For biomedical applications, the composition of the material being considered is a crucial factor since many of the elements that make up commercially available materials for industrial applications are extremely toxic to the human body. Therefore, in addition to meeting the mechanical properties needed for a particular biomedical application, the material must also be biocompatible. Ideally, a biodegradable biomedical device should be composed 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, calcium and zinc, see Table 3. Furthermore, the material should have a controllable dissolution rate or slow corrosion rate that permits the biomedical device or implant 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 biomedical implant are no longer required and the implant 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 implant it is necessary to control their corrosion behaviour in the body fluid environment[]. The use of alloying elements such Al, Ca, Li, Mn, Y, Zn, Zr and RE in Mg alloys can significantly improve the physical and mechanical properties of the alloy by: 1 refining the grain structure; 2 improving the corrosion resistance; 3 form inter-metallic phases that can enhance the strength; and 4 assist in the manufacture and shaping of Mg alloys.

Impurities commonly 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 from 2 to 4 ppm by weight, while Cu is ppm , Fe ppm and Ni ppm []. Since both Be and Ni are carcinogenic, their use in biomedical applications should be avoided as alloying elements.

While elements such as Ca, Mn and Zn are essential trace elements for human life and RE elements exhibiting 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[].

Mn is added to many commercial alloys to improve corrosion resistance and reduce the harmful effects of impurities[]. 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 calcium phosphate layer formed over the surface of the oxide layer[92, ]. Recent studies by Witte et al. Rods of 15 mm diameter and 20 mm long were inserted into the femur of guinea pigs and the rods degradation profile monitored. The implants were harvested at 6 and 18 weeks, with complete implant degradation occurring at 18 weeks.

All Mg based alloy implants were found to be beneficial and promoted new in situ bone tissue formation, while the polymer control rods produced a less significant effect. The LAE alloy had the greatest resistance to corrosion, while the other alloys all had similar, but lower values of corrosion resistance and degraded at similar rates[92].

While Mg is potentially an ideal biocompatible implant 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 example an AZ91 alloy rods were implanted into the femur of a number of rabbit models and the subsequent analysis revealed that after 3 months the implant had degraded and been replaced by new bone tissue[, ].

At the end of this degradation process most of the alloying elements such as Al would have been released 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[]. In humans, Al is a neurotoxicant and its long term accumulation in brain tissues has been linked to neurological disorders such as Alzheimers disease, dementia and senile dementia[]. In addition, the administration of RE elements such as cerium, praseodymium and yttrium has resulted in severe hepatotoxicity in rats[].

Furthermore, using heavy metal elements as alloying components are also potentially toxic to the human body due to their ability to form stable complexes and disrupt the normal molecular functions of DNA, enzymes and proteins[].

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Therefore, there is a definite requirement to carefully select alloying elements that are non-toxic to the human body, see Table 4. Non-toxic alloying elements such as Ca[] and Zr[] 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 viable implant material[33]. Surface Modifications and Treatment Processes for Biomedical Mg Alloys The high degradation rate of Mg and Mg alloy implants in the human physiological environment would result in the reduction of mechanical integrity of the implant before the bone tissues had sufficient time to heal[26].

There are two methods of reducing the degradation rate; the first, which was discussed in Section 3, involved alloying Mg with biocompatible elements that can resist the corrosion process. The second method is discussed in this section and involves the surface modification of the implant, through a treatment 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, reparative and remodelling. The reparative phase usually takes 3 to 4 months, during which time integration of the implant with the new and regenerated tissues takes place.

The final remodelling phase, which is the longest phase, can take from several months to years to complete[]. For 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 implant is compromised. The minimum period for this to take place is at least 12 weeks[26].

Unfortunately, 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 properties 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 modifications and treatments can have a significant role 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 properties of Mg bio-alloys. A brief overview of some of these surface modification processes are presented in the following the four sections. Mechanical Modifications to Induce Surface and Subsurface Properties The surface structure of an implant is very important, since it is the initial response of the surrounding tissues to the surface of the implant material that determines whether or not there is effective tissue-biomaterial integration. Studies of conventional types of permanent implant 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[]. In the case of Ti alloys, 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 proliferation than roughen surfaces[].

For Mg alloys, the influence of different mechanical processing operations during fabrication has the potential to greatly influence surface and subsurface properties[, ]. 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[], while chip removal from the surface during machining can directly influence the surface topography[]. 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[]. A recent study by Denka et al. The presence of residual compressive stresses after rolling also has the advantage of reducing micro-crack formation from 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 for biomedical applications[, ]. The importance of surface and sub-surface treatments on Mg alloy implants was recently investigated by Von Der Hoh et al. In their study three surface machining treatments were applied to an Mg-Ca 0.

The alloy was used to make three different geometric sample types. The smooth cylinders were machined with no further surface treatment, so they retained the micro-surface topography produced by the cutting tool. After 6 months of in vivo implantation in adult New Zealand white rabbits, the smooth cylinders revealed good integration with the surrounding tissues and also had the least structural loss.

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The sand blasted cylinders had the greatest material loss with the initial cylindrical shape completely consumed, while the threaded cylinders ranged between these two extremes. The results indicated that the smoother micro-topographic surface features of the cylinders were suitable for resorbable Mg alloys, while 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. Physical and Chemical Modifications From an engineering point of view, the most effective way to prevent corrosion is to coat the metal component with a protective barrier that effectively isolates the metal from the surrounding environment.

To be effective against corrosion, the protective coating must be uniform, well adhered and free from any imperfections such as pits, scratches and cracks. Therefore, surface cleaning and a suitable pre-treatment of the metal surface is a crucial factor in achieving an effective surface coating. 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 of problems to overcome.

For example, in most PVD processes the substrate temperature range is usually between and o C, but in the case of Mg the substrate temperature must be kept below o C for material stability reasons. The lower substrate temperature of Mg also influences the adhesive and corrosion resistance properties of the coating[, ].

The subsequent corrosion studies have revealed that the binary-alloyed surface coating were capable of increasing the corrosion resistance of the various Mg alloys[]. The chemical vapour deposition process has also been used to produce a variety of coating processes that can create a protective coating or modify the existing Mg alloy surface. During the deposition of a solid material from the vapour phase onto a usually heated substrate a chemical reaction over the surface takes place.

This results in changes to the sub-surface of the substrate, which chemically modifies the surface properties. For example, the deposition of diamond like carbon DLC films on metallic implants can improve the surface properties of the implant, thus making it biocompatible with the surrounding body tissues[].

Ion Implantation and Plating Ion implantation consists of bombarding the surface of a substrate with ionized particles. The ionized particles penetrate the surface and become embedded in the sub-surface of the substrate. The ionized particles soon neutralize in the interstitial positions within the grain structure forming a solid solution. During this process physiochemical changes take place in the sub-surface of the substrate, while the bulk properties of the substrate remain unchanged.

To date there have been relatively few studies carried out that have used ion implantation to enhance the surface properties of Mg alloys. A recent study by Liu et al. The study revealed that a compact surface oxide layer was formed, which was predominantly composed of TiO 2 with a smaller amount of MgO. In a similar study by Fang et al.

The results revealed that after ion implantation, the Zn had improved the surface hardness and elastic modulus of the alloy. The surface oxide layer formed during corrosion testing in simulated body fluid enhanced the Mg-Ca alloys corrosion resistance. However, Wan et al. Subsequent analysis of the results suggests that Zn was an unsuitable metal for ion implantation with Mg-Ca alloys for biomedical applications.

Ion plating is a technique that deposits noble metal ions onto a less noble metal substrate to form a dense and well-adhered layer. The plating layers improve surface properties such as topography, roughness, surface chemistry and wear resistance. Zhang et al. The results not only revealed a substantial improvement in corrosion resistance, but also found that an interfusion layer had formed between the Ti coating and the Mg substrate.

Thermal Spray Coatings During this coating process, materials such as metals, metal alloys, ceramics, polymers and composites are feed powder or wire form into a gun. The material is then heated to a molten or semi molten state within a gas stream. The resulting micrometer size droplets are accelerated in the gas stream, which is directed towards the surface of the substrate[]. This technique was successfully used by Zhang et al. There was significant diffusion of Al and Mg around the interface of the coating, which enhanced both the corrosion resistance and anti-wear properties of the coating.

Ceramic coatings such hydroxyapatite HAP , TiO 2 , Al 2 O 3 , and ZrO 2 have also been successfully applied to Ti alloys to improve their corrosion resistance, wear resistance and biocompatibility[]. However, in a study by Zeng et al. The study also revealed that galvanic corrosion occurred between the surface of the Mg alloy and the coating layer, which effectively reduced any protective properties offered by the coating.

This highlights the weakness of thermally sprayed ceramic coatings which have rough surfaces, high porosity and poor adhesion properties[]. Laser Surface Melting, Alloying and Cladding The high-density energy of a laser beam can be effectively used to modify the surface region of Mg alloys. The surface region of the alloy can be melted to create a meta-stable solid solution. This is then followed by rapidly cooling the substrate, which results in the refinement of the surface microstructure.

This technique can also be used to improve the surface properties of an Mg alloy substrate by melting a metallic coating and the underlining sub-surface. During the rapid melting process both the coating and sub-surface mix before re-solidifying during subsequent cooling to form a new surface alloy which coats the bulk of the substrate. Furthermore, if the appropriate alloying metals are incorporated into this surface modification technique it is possible to significantly improve surface properties such as corrosion resistance[, ].

For example, improved surface properties of a Mg alloy AZ91 have been achieved with the dispersion of hard metallic particles such as TiC and SiC in the molten pool generated by laser melting[, ]. Also, laser cladding of an Al-Si alloy onto a number of Mg alloys such as AS41, AZ91 and WE54 have also been attempted, but unfortunately, the surface properties were not significantly improved[, ].

Wet Chemical Processes 4. Electrochemical Deposition of Metallic Coatings The corrosion resistance of Mg and its alloys can be increased by an electroplating technique. In this technique a metal salt is reduced in solution to its metallic form, the electrons for reduction are supplied from an external source and the resulting metallic ions are deposited on to the surface of the substrate. However, most metals are more electrochemically noble than Mg, which can cause serious problems if there are any imperfections in the deposited layer.

Such imperfections will expose the underlining substrate and result in the formation of small localized areas of corrosion. The corrosion sites form highly corrosive pits that tunnel down into the Mg substrate and seriously weaken the substrate[]. From an industrial point of view, electroplating is a highly effective technique for coating Mg and its alloys with metallic coatings such as nickel, chrome and aluminium coatings[].

These coatings have good mechanical properties and provide effective corrosion protection. Unfortunately, these metals are also harmful to human tissues, which make them highly unsuitable for biomedical applications. Chemical Conversion Coatings Chemical conversion coatings are formed by chemically treating the surface of Mg and Mg alloys to produce a thin outer coating of metal oxides, phosphates or other compounds that are chemically bonded to the surface[]. The conversion coating acts as protective barrier that isolates the substrate from the surrounding environment and prevents the corrosion.

Many of the processes used to produce conversion coatings involve the use of toxic materials that are detrimental to human health. An alternative treatment to chromate conversion coatings are: 1 phosphate; 2 phosphate-permanganate; and 3 stannate coatings. All three of these conversion treatments have comparable corrosion resistant properties to those of chromate treatments.

Xu et al. During the treatment process a biocompatible brushite layer[CaHPO 4. Subsequent immersion in SBF revealed that the brushite layer transformed into a coating of HAP, with the excess phosphate ions being released into the surrounding environment. The released phosphate ions were also found to neutralize the alkalization effect produced by the corrosion process. The treatment process did not prevent corrosion, but it did significantly slow down the degradation rate[].

And a phosphate-permanganate process developed by Han et al. While a stannate treatment developed by Gonzalez-Nunez et al. The coating was adherent, continuous, and crystalline which produced a passivating effect on the substrate surface[]. Unfortunately, no degradation rate data was reported, indicating that more studies are needed to indeed gauge the effectiveness of this process for in vivo applications. Magnesium fluoride MgF 2 conversion coatings on Mg alloys have produced mixed results in providing corrosion protection.

Degradation studies carried out by Zeng et al. While in a similar study Hassel et al. And an in vivo study carried out by Witte et al. In a similar study by Gao et al. The two RE elements under investigation were Ce and Y, with each element being used individually to form a surface treatment solution. The study revealed that the second conversion layer had improved corrosion resistance compared to the first.

However, both coating provided limited corrosion resistance due to their thin thickness and soft structure, which was incapable of withstanding minor mechanical damage. Furthermore, both the toxicology and metabolic pathways within the human body of RE elements such as Ce and Y are still unclear and need to be fully investigated before they can be used in biomedical applications. Calcium Phosphate Surface Coatings A more biocompatible form of conversion coating can be derived from a variety of calcium phosphate compounds.

There are three major advantages in using HAP in hard tissue engineering applications: 1 it has good biocompatibility and bioactivity properties with respect to bone cells and other body tissues; 2 it has a slow biodegradability in situ ; and 3 it offers good osteoconductivity and osteoinductivity capabilities[, ]. These properties are very important because bone tissue constantly undergoes remodelling, a process in which bone tissue is simultaneously replaced and removed by the bone cells, osteoblasts and osteoclasts respectively. It is these advantages that make HAP and TCP tri-calcium phosphate compounds attractive for coating metallic orthopaedic implants.

In this application, both HAP and TCP coatings promote bone formation which enhances bonding between the implant and the surrounding tissues. It is also due to these positive biological responses within the human body that has made calcium phosphate coating an attractive option for potentially reducing the biodegradation rate of Mg orthopaedic implants. Several techniques have been used to deposit calcium phosphate coatings onto Mg substrates, these range from anodization[], bio-mimetic coatings[], electro-deposition[,],hydrothermal[] and wet chemical methods[, ].

Unfortunately, the layer formed was porous and did not prevent corrosion in a simulated body fluid. However, the degradation rate was significantly reduced and provided the Mg alloy substrate with reasonable protection against the corrosive effects of the simulated body fluid. The study also found that the brushite coating was able to improve the surface biocompatibility of Mg alloy substrate, since the brushite coating transformed into a HAP phase with time.

Also during this transformation, acidic phosphate ions were released into solution, which tended to have a neutralizing effect on the alkalization process[]. Furthermore, the surface treatment enhanced the bioactivity of the Mg-Mn-Zn alloy and promoted bone formation[]. A similar calcium-phosphate coating was produced by Wang et al. The DCPD layer was effective in providing protection for the Mg substrate during the first 21 days of immersion in a simulated body fluid.

Recently, Yanovska et al.


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The involved dipping a Mg substrate into an aqueous solution containing Ca NO 3 2. Deposition of both DCPD and HAP phases under the influence of magnetic fields lead to crystal orientation during the formation of the phases. The technique also produced coatings with enhanced corrosion resistance, which in turn reduced the degradation rate[]. In an alternative method, Song et al. The study found the FHA coating had long-term stability and remained intact even after 1 month of immersion in a simulated body fluid and provided effective corrosion resistance to the Mg alloy[].

Alkali Heat Treatments Heat treatment can be a beneficial way of improving the microstructure and enhance the surface properties of Mg and Mg alloys. The corrosion behaviour and cytotoxicity of alkali heat-treated pure Mg samples immersed in simulated body fluid SBF were investigated by Li et al. SBF solutions with and without chloride ions were used to study the influence of chloride ions on the corrosion behaviour of treated Mg samples. All the treated samples showed a significant improvement in corrosion resistance in both SBF Cl - and SBF solutions compared to the untreated Mg samples.

In addition, after 14 days of immersion in the SBF fluids, a calcium phosphate compound with a molar ratio of 1. While the subsequent cytotoxicity testing revealed no signs of morphological changes in the cells and no inhibitory effect of the surface treatments on cell growth could be detected. In a similar study by Liu et al.

In addition, the study found that samples microstructure significantly influenced the overall corrosion morphology. For example, the surface of the untreated samples displayed deep and uniform corrosion, while the surface of the treated samples had only shallow pitting[]. Anodization The anodization of magnesium is an electro-chemical process that changes the surface chemistry of the metal, via oxidation, to produce a stable anodic oxide layer.

The structure of this layer is characterized by a thin barrier layer at the metal-oxide interface, followed by a less dense porous oxide layer. The porous layer can display a variety of different structures and properties which are dependent on the composition, substrate micro-structure and processing parameters[]. The processing parameters that influence oxide layer formation include: 1 the type, temperature and concentration of electrolyte; 2 current density; and 3 the applied anodization voltage. These parameters can also significantly influence the resulting corrosion behaviour of the substrate[].

The anodization process can also produce an oxide layer consisting of pores, whose size and density is dependent on the selection of the appropriate processing parameters. Industrially, porous oxide layers are usually coloured and then sealed or form part of a pre-treatment process prior to painting or coating. Many of the industrial coating and surface treatments used on anodized Mg alloy components to reduce corrosion are toxic to the human body. And as a result, research efforts have focused on searching for biocompatible surface treatments and process that are non toxic.

For example, Hiromoto et al. The advantage of this technique comes from the Ca and P elements being deposited on the substrate, since both bioactive materials are known to induce osteoinduction and promote new bone tissue growth[]. Micro-arc oxidation MAO is an electrochemical process which uses a high anodic voltage and high current density to create an intense micro-arc plasma near the metal surface to induce oxidation. The oxide layer formed during this process is substantially thicker than conventional anodization, since the sub-surface of the metal substrate is also oxidized[].

Liu, Y. Huang, X. Fathi, T. Zarede, M. Ayad, M. Lieser, J. Leyens, M. III, pp. Popescu, H. Iancau, L. Hancu, A. Wuchina, E. Opila, M. Opeka, W. Fahrenholtz, I. Haruna, A. Abdulrahman, P. Zubairu, L. Isezuo, M. Abdulrahman, D. Ashby, K. Abba, I. Nur, S. Sahari, S. Jagadeeshgouda, P. Reddy, K. Dhal, S. Rao, D. Rao, N. Gladston, N. Sherrif, I. Dinaharan, J. Callister, D.

Jain, A. George, A. Bockarie, H.

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