Magnesium Alloys for Biomedical Applications [306697]

[anonimizat], [anonimizat], [anonimizat]-particularly in orthopedic applications. [anonimizat], biocompatibility, good mechanical properties and safe biodegradation. [anonimizat]. Current research approached a [anonimizat], adhesion, the most elusive aspect being the biocorrosion of Mg alloys. [anonimizat]. [anonimizat]. [anonimizat], performed either using conversion or deposition methods which were developed for biomedical applications. In this paper was analyzed the influence of different elements in magnesium alloys on the corrosion properties. Their directions as biodegradable materials are also briefly described in this chapter.

Keywords: [anonimizat], [anonimizat] ( approximately 20-28 grams of magnesium) [1]. Skeletal system contain over 50% [anonimizat], muscle and bodily fluids. [anonimizat], in supporting a [anonimizat]. Magnesium is an element from Group 2 (Group IIA in older labelling schemes) and the elements which are from Group 2 are called alkaline earth metals. In figure 1 it is presented the symbol of magnesium and also atomic number and relative atomic mass. Magnesium is a substantial intercellular cation which is involved in more than 300 biological reactions of cell. Also, magnesium is a [anonimizat] a stabilizer for the structures of DNA and RNA [2]. Magnesium is the lightest metal being approximately 34% lighter by volume than aluminum and 50% lighter than titanium which exposes a [anonimizat], excellent vibration and shock absorption, a high damping capacity and electromagnetic shield performance [3-6]. However, the particularity of magnesium which differentiates from other materials is its ability to biodegrade in vivo. An implant needs to last as longs as its intended function in the body without needing surgical retriever. Thus biodegradability represent a [anonimizat] [7]. Stainless steel[8,9] , Co–Cr-based alloys[10-13] , titanium-based alloys[13-19] , zirconium-based alloys[20,21] [anonimizat][22,23] were some of the first metallic materials used as biomedical implants. All these kinds of biomedical metals cannot degrade in vivo and will permanently exist in human body after implantation.

Fig. 1 Magnesium in the Periodic Table of Elements

Although magnesium was first discovered by Joseph Black in 1755, it was first fully isolated by Sir Humphry Davy in 1808, which created Mg metal electrolytically. In 1878, for the first time, magnesium was mentioned as a biodegradable implant material in a report of physician Edward C. Huse about absorbable ligatures for the closure of bleeding vessels. So, after magnesium alloy has been found biodegradable in the human body in the 1930s, it was used in the early 20th century for fast absorbable wound closure and fixing fractured bones and in several investigations as biodegradable material for connectors for vessel anastomosis and wires for aneurysm treatment, applications which has been discussed in a review by Witte in 2010. [24]

However, Heublein et al.  [25,26] were the first which investigated the idea of using magnesium alloys for cardiovascular stents . They chose AE21 alloy which has lower degradation rate in comparation with other magnesium alloys for an initial coronary animal study. The disadvantage of AE21 stent was that its degradation occurred faster that the expected rate.

So due to its advantageous properties in the human body, Mg and its alloys have been studied during the last two centuries as implant materials for numerous medical applications such as cardiovascular, musculoskeletal and general surgery. Because nowadays magnesium production technology is highly developed, there are expectations for the applicability of magnesium and its alloys in various fields. Magnesium and its alloys have many advantage over traditional metallic materials, ceramics and biodegradable polymers, in load-bearing applications is higher tensile strength and a low Young’s modulus (41-45 GPa) that more closely matches bone and it is also remarkable for its compressive yield strength and fracture toughness compared to hydroxyapatite. Moreover, they offer many other benefits over many current implant materials, including low density, high damping capacity, ease of machinability, biocompatibility, bioresorbable properties and osteogenicity. To date magnesium and its alloys have especially been studied in the development of cardiovascular stents, bone fixation materials, and porous scaffolds for bone repair. [27-29] However, the use of magnesium as an implant material has been limited due to its rapid corrosion in the presence of the body fluid; magnesium is one of the most electrochemically active metals. Under an accelerated degradation, toxicity and hydrogen evolution become problems, so rapid evolution of gas causing the appearance of a balloon effect in subcutaneous tissues, it inhibits wound healing and finally, leading to the implant failure . Rapid corrosion is an intrinsic response of magnesium alloys to chloride containing solutions, including the human body fluid or blood plasma, and beside the fact that corrosion is too rapid even for a biodegradable material; it is not homogeneous, due to a strong tendency for localized corrosion exhibited by Mg alloys.

Pure magnesium (unalloyed) is the main cause which produce an accelerated corrosion due to presence of impurities in it structure such as Fe, Ni, Cu. So by pure purification of smelting, it can greatly reduce content of harmful impurities of heavy metals in magnesium alloys, which can effectively improve the corrosion resistance of magnesium alloys. Alloying is a way of improving the corrosion resistance of magnesium. Rare earth elements make magnesium alloy microstructure improved, and strengthen the grain boundaries and phase boundaries, thus slow down the corrosion of magnesium alloys in NaCl solution. [30]

Types of alloys

Magnesium alloys can be used in various applications, but they easily can be classified into two groups: sand casting alloys and diecasting alloys. Alloys also can be classified as general purpose, high-ductility and high-temperature with higher levels of iron, nickel and copper.

The Mg based biodegradable substrates can be divided into four major groups:

(a) Pure Mg,

(b) Al containing alloys (AZ91, AZ31, LAE422, AM60 etc.)

(c) Rare earth elements (AE21, WE43 etc.), and

(d) Al free alloys (WE43, MgCa0.8, MgZn6 etc.).

These alloying elements improve the mechanical and physical properties of Mg alloys for orthopedic applications by:

(a) Optimizing grain size,

(b) Improve corrosion resistance,

(c) Providing mechanical strength by the formation of intermetallic states,

(d) Eases the manufacture process of Mg alloys.

The composition of some Mg based alloys is given in Table 1.

Table 1. Elemental compositions (%wt) of selected Mg alloys

In this section, the effects of some important alloying elements on the mechanical properties and corrosion behaviors of Mg alloys are also discussed.

Aluminum is the principal alloying element for many magnesium alloys as it can improve the mechanical strength, corrosion properties and stability of magnesium castings. The most widely used general purpose sand casting alloy is AZ91. In the alloy nomenclature, the letters A and Z denote the major alloy.

The binary Mg-Al system was the basis for early magnesium casting alloy development. The maximum solid solubility of aluminium is 12.7% at 437 oC decreasing to about 2% at room temperature. By adding a quantity of Aluminum (1-5%), grain size will decrease considerably, while nitrogen will not influence the grain size. After partial dissolution of Al in Mg solid solutions, this precipitates as Mg17Al12 secondary phases along the grain boundaries. The as-cast Mg-Al alloys show α-Mg matrix and β- phases mainly consisting of Mg17Al12 phases. In the presence of electrolytes, these phases show different electrode potentials. The Mg17Al12 phase exhibits a passive behaviour, acting as a cathode with respect to the α-phase of Mg matrix, thereby accelerating the corrosion of the alloy. However, the inert nature of Mg17Al12 phase can itself act as a corrosion barrier, thereby reducing the corrosion in AZ91D alloys. The suggestion of Song et al. was that by increasing the volume fraction and through its distribution along the grain boundary, it could act as a barrier against corrosion surrounding the α-Mg matrix thereby reducing the corrosion rates. Clark has studied the precipitation hardening mechanism of these alloys, and found that the aging resulted in the transformation of supersaturated solid solution directly to a coarsely dispersed, equilibrium of β precipitate without the appearance of the Guinier–Preston (GP) zones or intermediate precipitates. AZ91C is the most widely used alloy among these zirconium-free alloys. Clark demonstrated that, the discontinuous precipitation, which was most pronounced in the early stages of aging, was made up of alternating lamellae of Mg17Al12 and equilibrium composition matrix growing behind a migrating grain boundary. Calcium and Zn are two essential elements in human body that also provide mechanical strengthening in Mg-based alloys.

Calcium (Ca) acts as a grain refining agent in Mg alloys, stabilizing the grain size at levels up to 0.5% Ca content and decreases slightly with further addition. Moreover, Ca is a major component in human bone and can accelerate the bone growth. It was thought that the presence of Ca ion benefits the bone healing. However, for chloride ions, low corrosion rates were obtained due to the formation of a partially protective Mg(OH)2 layer ( the structure is presented in figure 2).

Fig. 2 Structure of Mg-Ca biodegradable alloy

Li et al. studied the Mg-Ca binary alloys with increasing concentration of Ca from 0.5% to 20%. The high content of Mg2Ca secondary phase distribution towards grain boundaries result from increasing the quantity of Ca. The Mg2Ca secondary phase is brittle reducing the ductility of Mg-Ca alloys with increase in Ca concentration. This also influences the corrosion properties of Mg alloys. A high volume fraction of the Mg2Ca secondary phase reduces the corrosion resistance of Mg-Ca alloys due to the formation of micro-galvanic cells. Therefore, it can be observed that the excessive Ca concentration accelerates the corrosion of Mg-Ca alloys, and the optimum Ca concentration should be ≤1%. The addition of 1% calcium improves the creep strength of Mg-Al alloys but makes them susceptible to hot cracking.

Manganese (Mn) is frequently used as a secondary element in Mg alloys. It has been noticed that when the content of Mg increases in Mg-Al-Mn alloy the grain size decreases but at levels above 0.4% this effect ceases. It has also been reported that Mn addition can improve the tensile strength and fatigue life of extruded AZ31, AZ61 and AZ21 alloys. Song et al. suggested that Mn improve the corrosion resistance in Al containing Mg alloys by transforming the Iron (Fe) and other impurities into harmless intermetallic compounds. However, the excessive addition of Mn reduces the corrosion resistance of Mg-Al alloys due to the formation of a large amount of Mn-containing Mg-Al intermetallic phases which can be prone to galvanic effects.

Zinc (Zn) has similar characteristics to those of manganese, meaning it has the ability to transform impurities such as Iron (Fe), Copper (Cu) and Nickel (Ni) into harmless intermetallic compounds, thereby reducing their corrosion enhancement effect. It was observed that with the addition of zinc occurs the grain refinement and formation of secondary phases, thus influencing the mechanical and corrosion properties of Mg alloys. Yin et al. showed that the addition of 3% Zn in Mg-Zn-Mn alloys forms Mg-Zn secondary phases which precipitate out of the Mg-matrix, thus improving the strength through a dispersion strengthening mechanism.

Song et al. also studied the effect of Zn alloying on corrosion behavior of Mg alloys. It was found that micro-galvanic effects dominate the corrosion behaviour of Mg-Zn alloys, thereby restricting Zn to levels b5%. Above 5% addition results in the formation of high volume fraction of Mg-Zn secondary phases which act as cathodes, accelerating the corrosion of α-Mg matrix around the Mg-Zn phases in Mg alloys.

Lithium (Li) can alter the lattice structure from hexagonal close packed (h.c.p.) to body-centered-cubic (b.c.c.) crystal structure in Mg alloys. Therefore, it can be employed to enhance the ductility and formability of Mg alloys. Li is more reactive than Mg and has a pronounced effect on the corrosion behaviour of Mg alloys. Li content below 9% in pure Mg is beneficent to corrosion resistance, whereas excess addition of Li is detrimental to the corrosion resistance.

The group of rare-earth elements (REEs) contains seventeen elements, including fifteen lanthanides, scandium (Sc) and yttrium (Y). They are normally added to Mg alloys as master alloys or hardeners and can improve the strength and corrosion resistance by both solid solution and precipitation hardening. The REEs can be classified into two groups:

High solubility group: (yttrium (Y), gadolinium (Gd), terbium (Tb), dysprosium(Dy), holmium (Ho), erbium(Er), thulium(Tm), ytterbium (Yb), and lutetium(Lu))

Limited solubility group (neodymium (Nd), lanthanum (La), cerium (Ce), praseodymium (Pr), samarium (Sm), europium (Eu)).

For example, Y has high solid-solution solubility in Mg and is often introduced along with other rare earth elements to improve creep and corrosion resistance. Moreover, most of REEs form intermetallic phases with Mg and Al which have a pronounced effect on the strength and corrosion of Mg alloys. Several REE-doped Mg alloys such as WE43, Mg-5Gd, LAE442 and Mg-4Y have been investigated. In WE43, main precipitated phases are

Mg12YNd and Mg14YNd. For Mg-Gd alloy, Mg5Gd intermetallic phases can precipitate both in grains and at grain boundaries. In Mg- Al-REE systems, the REE tend to form intermetallic phases with Al as Al12REE and Al11REE3 improving strength and corrosion resistance [31]. Furthermore, the REE elements with limited solubility tend to form inter-metallic phases early during the solidification process; for example, Ce aggregate at solid-liquid interfaces during solidification in Mg-Al-Ce alloys. During solidification, Al-Ce secondary phases form and segregate along the grain boundaries, effectively blocking the sliding of boundaries during deformation. Al-Ce particles also influence the corrosion rate in Mg-Al-Ce alloys. Higher content of Ce in alloys leads to Al11Ce3 particles forming a network surrounding the Mg matrix acting as a micro-galvanic cathode and thus delaying the corrosion of the Mg alloys due to a very small potential difference. Additionally, Al-Ce phase exhibits passivation in a wide range of pH, which further retards the corrosion of Mg alloys.

Properties and limitations

Magnesium (Mg) is a lightweight, silvery-white metal which exhibits a high strength to weight ratio, good thermal and electrical conductivity, 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 aluminium and steel, respectively. [35,36]

Interestingly, density and modulus of elasticity of magnesium have values close to those of natural bone. Because of this similarity, magnesium can be easily used in tissue applications because it reduces 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. [37,38]

Magnesium has a hexagonal crystal structure with a =0.320 nm, c = 0.520 nm, c/a = 1.624. The basal plane is close packed and the axial ratio is only slightly greater than the theoretical value for incompressible spheres. The atomic diameter is 0.320 nm, so there is a favorable size factor with a diverse range of the solute elements aluminum (Al), zinc (Zn), cerium (Ce), yttrium (Y), silver (Ag), zirconium (Zr), and thorium (Th). [39]

Magnesium and its alloys have many outstanding properties relative to other engineering materials such as: low density, great damping capability [40, 41], excellent fluidity for casting, good electric shielding effect, high strength, non-magnetic, satisfactory heat conductivity, low heat capacity, negative electrochemical potential and non-toxicity. [42] Generally, magnesium is used as an alloying addition in aluminum alloys [43] and it is indispensable in nodular cast iron. It is also used in aerospace and general transport industries, in electronics and at equipment for material handling. In these engineering applications, magnesium is rarely used in its unalloyed form. The magnesium alloys find principal use in non-structural application although the structural uses are expected to grow in importance in the future. Recently, die cast magnesium alloys have found applications in computer disk drive and magnetic card readers at supermarket checkouts. [43] Also, because of the ability of magnesium based materials to degrade it has led to a multitude of medical applications. Recently, studies have focused on the biological environment-biomaterial interaction and the cellular mechanisms to highlight how the materials are biologically influenced by the dissolution of corrosion byproducts from the bulk of the material. [44-46]

Advanced studies are also exploring how polyaprolactone (PCL) and polylactic acid (PLA) polymer coatings influence the corrosion behavior and drug eluting kinetics for biodegradable stent applications . Orthopedic applications of magnesium shows excellent interfacial resistance at implantation. Magnesium materials have also been used for

different types of fixation devices for orthopedic surgery, such as screws, plates, and fasteners

In recent studies it was also shown that implanting a magnesium device shows no significant changes in blood composition, without causing damage to excretory organs like the liver or the kidneys . [47-50]

Unlike other lightweight alloys, magnesium alloys has many advantages which gives them a wide field of application. For example, magnesium preserves the light weight of a design without sacrificing strength and rigidity. This benefit is important when portability is a key element of the product design, such as with chainsaws, cellular phones or circular saws. What makes it special is its ability to absorb energy. Thus by increasing the capacity of absorption of vibrations ensure quiet operation of equipment.

Structural changes can affect the size of some metals after prolonged exposure to high temperatures, but the magnesium alloys is not the case. Magnesium shrinkage rates are more consistent and predictable in comparison to other nonferrous metals, hence they have minimal residual casting stress.

Magnesium alloys possess also a low galling tendency and can be used as a bearing surface in conjunction with a shaft hardness above 400 Brinell measurement.

In tables 2 and 3 are briefly presented mechanical and physical characteristics of different magnesium alloys. [51]

Table 2. Typical Mechanical Properties of Magnesium at Room Temperature

Table 3. Typical Physical Properties of Magnesium

However, the main disadvantage of Mg and Mg alloys is their low corrosion resistance; magnesium is one of the most electrochemically active metals. Low corrosion resistance results in the rapid release of degradation products. The degradation under physiological conditions may cause a reduction in mechanical integrity of the implant before the bone or tissue is sufficiently healed. Low corrosion resistance of magnesium also lead to the emission of hydrogen gas which finally will lead to the formation of gas bubbles. The accumulation of these bubbles around the implant will delay healing of the tissue and increase the pH around the implant, this increase causing local alkalization and severely affect the pH dependent physiological processes in the vicinity of the implant. [52]

The corrosion process happens basing on an electrochemical reaction. The overall corrosion reaction of magnesium in aqueous solution can be given as below:

Mg(s)+2 H2O(aq)⇌Mg(OH)2(s)+H2(g) (1)

This overall reaction may be divided into the following partial reactions:

Mg(s)⇌Mg2+(aq)+2 e− (anodic reaction) (2)

2H2O(aq)+2e−⇌H2(g)+2OH−(aq) (cathodic reaction) (3)

Mg2+(aq)+2OH−(aq)⇌Mg(OH)2(S) (production formation) (4)

Although during the corrosion process the magnesium hydroxide layer accumulated on the underlying magnesium substrate can function as a protective layer preventing subsequently corrosion, if the concentration of chloride in the environment rises above 30 mmol/l, the magnesium hydroxide is prone to convert into magnesium chloride, which is highly soluble in water. Therefore, the spontaneously formed magnesium hydroxide layer cannot protect the substrate effectively.

Accentuated corrosion and accelerated degradation are inevitable for magnesium-based implant in vivo where the chloride content is about 150 mmol/l [53], which far exceeds the value that magnesium hydroxide can bare. To successfully integrate bioabsorbable metal implants, the time frame of degradation must be sufficient such that the cells to be able to synthesize and deposit an extracellular matrix for their own support and function before the structural integrity of the implant is compromised. In order to favor solidarity of the cells , surfaces have been treated or coated with a variety of chemistries and polymers. After their adhesion, cells create a substrate on the implant comprised of the proteins necessary for their function and survival. For example, osteoblasts are bone-forming cells which are found at the surface of bone. They can be stimulated to proliferate and and then mineralize it to make new bone. Alternatively, an elastic protein, aptly named “elastin,” is abundant in arteries to expand and recoil when the large volume of blood comes from heart contraction. After the cells are supported, absorption of the implant would leave behind a naturally synthesized protein structure appropriate for those cells at that specific site. For these reasons, a significant, uncontrolled, local change in Mg concentration due to implant degradation can have a deleterious effect on human physiology and must be managed through proper engineering design. Hence for extensive use of Mg and its alloys in biomedical implant applications, a better understanding and control over the degradation rate is required. [54, 55]

So magnesium is a resorbable biomaterial, promising, with deficiencies related to corrosion, emission of hydrogen and biointegration. The coatings seem to be a solution in reducing and controlling the corrosion rate and increasing their initial biocompatibility.

Coatings material and pre-deposition procedures

As can see from the above, Mg, its biodegradable alloys and their corrosion products are well tolerated by the human body. However, in the majority of the cases, the in vivo corrosion kinetics of magnesium and its alloys is much faster than that of the bone healing; so considering this issue, kinetics for implant applications must be slowed down. Numerous researches have shown that surface modifications, such as polishing, oxidation, passivation, coating deposition, ion-implantation, etc. have influence on the properties and functional activity of an implantable biomaterial [56, 57]. From all these techniques, the most effective way to achieve surface modification, and moreover to improve the osseointegration and biocompatibility of metallic implants. There are many ways to improve the corrosion resistance of pure Mg. Briefly, they comprise the following approaches: microstructure tailoring including grain size and texture, alloying, preparation of biocomposites, surface treatment and deposition of protective coatings. [58] The focus of this review concerns the development of biocompatible and biodegradable coatings for Mg and Mg alloys, with the intent of reducing and controlling the corrosion rate and increasing their initial biocompatibility. It should be emphasized that, to improve corrosion resistance, the surface of Mg and its alloys can be coated with either calcium orthophosphates alone (the vast majority of publications cited in this review) or calcium orthophosphate-based biocomposites. As the use of implantable calcium orthophosphates started in 1920 [59, 60], while that of calcium orthophosphate based biocomposites and hybrid biomaterials started only in 1981[61], one can conclude that the latter is just at the initial stages and many more publications are expected in the near future. The coatings are used in order to minimize the initial localized corrosion of magnesium and its alloys. The coatings on magnesium-based implants that are usually a layer with a proper thickness are designed to protect the substrate from severe corrosion, particularly at the initial stage after implantation. Since the layers coatings are temporary, they can gradually disappear in vivo without causing harmful effects to the surrounding tissues. Generally coatings can be divided depending on the involvement in the coating formation into three classes to: substrate involving coatings, non-substrate involving coatings and composite coatings. The current progresses of various coatings formation are reviewed as follows and Pros and Cons of different coatings are also discussed. [52] Coating materials are in generally related with coating methods. Even coating methods will be described later, because it is important to present minimal information about these coating methods before make the discussion about coating materials. The coatings method are shown briefly in table 4.

Table 4. Organic or inorganic based coating by different methods

Coating materials which have been widely investigated until now are Ca–P compounds and other bioceramics like bioglass, as well as organic coatings mainly composed of biodegradable polymers.

Inorganic coatings

Inorganic coatings have been applied to improve the corrosion resistance of magnesium and its alloys, such as DLC (diamond like carbon) coatings [62,63], TiO2 coatings[64] and ZrN/Zr bilayered coating[65] . Although these coatings present a very good corrosion resistance, the non-degradability limits their application in biomedical engineering, where the eventually degradation of the coatings is required. [52]

Bioglasses have been widely used as bone cement and scaffold because of their excellent biocompatibility. Taking into consideration their excellent biocompatibility, bioglasses were coated on magnesium sponges. Evaluation of the degradation behaviour and biocompatibility showed that all coated magnesium sponges were tolerated well without causing emissions of hydrogen or severe bone changes. It has been observed that after implantation there are different sized implant reduction and new formations of bone around the implant.[66]

Ca–P-based coatings are the most commonly used and investigated as biodegradable inorganic coatings to improve the corrosion resistance of magnesium and its alloys as well as their surface bioactivity. Apart from the commonly used methods, cold spray process [67] and transonic particle acceleration process [68] have been used to fabricated Ca–P-based coatings on magnesium and its alloys

Organic coatings

Degradable polymers have been applied in various field including biomedical applications [69,70]. In terms of biocompatibility, it is preferable to use these polymers as coatings for magnesium and its alloys. Among those polymers, polylactide (PLLA) and polycaprolactone (PCL) are mostly researched. PLLA is a biodegradable polymer which is characterized by good mechanical properties and high biocompatibility. Moreover, the end products can be removed by the body through fluids and then metabolized by the liver and kidneys. Xu and Yamamoto prepared biodegradable polymer films of PLLA and PCL on magnesium by spin coating in order to improve its early corrosion resistance and cytocompatibility, and after that they compared the two polymers [71]. The results showed that PLLA had a better adhesion strength with Mg substrates than PCL and that more cells were more proliferative on PLLA despite the fact that the amorphous PLLA and semi-crystalline PCL coatings presented uniform and nonporous surface structure on Mg.

Fig.3 Surface morphologies of PLLA, PCL, PEDOT: (a) Surface morphologies of compact polymer coatings on magnesium substrate; (b) Surface morphologies of porous PCL coatings on AZ91 alloy; (c) Surface morphologies of PEDOT conductive coatings fabricated on magnesium with different processes

The results of in vitro dynamic degradation of pure Mg with PLLA and PCL coatings also showed that PCL had better corrosion resistance in modified simulated body fluid solution than PLLA . Wong et al made a pore size controllable PCL coating on magnesium alloy which enhance the performance of AZ91 alloy in orthopaedic applications. The results demonstrated improved corrosion resistance and good cell biocompatibility.[72] Gollwitzer et al also showed that a PLLA coating for orthopaedic implants based on three different alloys had good stability.[73] Abdal-hay et al studied an HA-doped PLLA coating with respect to the bioactivity and corrosion behaviour of AZ 31 alloy as an orthopaedic implant; the coated samples showed a better biocompatibility and bending strength.[74]

Poly(lactic-co-glycolic acid) (PLGA) has been used in drug-delivery systems and in tissue engineering for decades .[75,76,77] PLGA is artificially synthesized polymer which has good biocompatibility and controlled degradation rates with different ratio of PLA/PGA. Ostrowski et al. fabricated coatings of varying thickness on magnesium alloy substrates using PLGA polymer at various concentrations with applications in orthopedics. The results of in vitro studies of cell viability showed improved biocompatibility of polymer-coated substrates even if some tests reported foreign body reactions with PLGA . [78] Because the magnesium substrate plays a role as a supporter, the lack of mechanical strength of these biodegradable polymers is also prevented. Because of the physically adhesion of the polymers to the magnesium substrate, between magnesium substrate and the polymer coatings being an obvious interface, the adhesion strength may not reach the requirement for biomedical applications, causing a potential danger of polymer coatings peeling off during the degradation.

Poly(ether imide) (PEI) has good mechanical properties and is stable at high temperatures and has therefore been explored as a coating for Mg alloys. PEI has been studied as a coating for the Mg alloy AZ31 by da Conceicao[79] and coworkers and Scharnagl et al.[80] . The thin layers of PEI presented high resistance to corrosion at exposing to a 3.5% NaCl solution.

Surface modification of magnesium-based implants by grafting biofunctional molecules is a feasible method to construct a functional surface with good biological performance. Liu et al. fabricated biofunctionalized anti-corrosive silane coatings on magnesium alloys using bistriethoxysilylethane and 3-amino-propyltrimethoxysilane[81]. The layer of densely crosslinked bistriethoxysilylethane coating was immobilized on NaOH-activated Mg surface in order to obtain a better corrosion resistance. Then 3-amino-propyltrimethoxysilane was grafted onto the pretreated surface to impart amine functionality to the surface. Furthermore, heparin was covalently conjugated onto the silane-treated magnesium alloy to render the coating biocompatibility, as indicated by reduced platelet adhesion on the heparinized surface. So by constructing multilayer on magnesium alloys step by step, a multifunctional surface can be easily obtained.

Sebaa et al. used poly(3,4-ethylenedioxythiophene) (PEDOT) which has applications in neurosurgery, as a conductive coating to prevent the degradation and improve the cytocompatibility of magnesium substrate. The coatings presented porous structure and adhesion strength within the classifications of 3B to 4B . The coatings improved significantly the corrosion resistance of magnesium.

For neural applications, Sebaa et al.  used poly(3,4-ethylenedioxythiophene) (PEDOT) as a conductive coating to control the degradation and improve the cytocompatibility of magnesium substrate. The coatings showed porous structure and adhesion strength within the classifications of 3B to 4B . The corrosion resistance was significantly improved by the coatings. Moreover, the PEDOT coatings could load the anti-inflammatory drug dexamethasone during the electrodeposition, which could be subsequently released upon electric stimulation . The surface modification by formation of self-assembling monolayers of nontoxic organic molecules is a facial way to design a functional surface for biomedical applications. [82] Grubac et al. fabricated alkylphosphonate self-assembled films on AZ91D alloy. The existence of well organized and ordered self-assembled alkylphosphonate monolayers showed good protecting properties in physiological solution.[83] Ishizaki et al. used vapor phase method to fabricate alkanoic and phosphonic acid-derived self-assembled monolayers on magnesium alloy. The contact angle hysteresis of SAMs with a carboxylate headgroup is much larger than that of SAMs with a phosphonic acid group. The phosphonic acid-derived SAMs had higher molecular density and better corrosion resistance compared with alkanoic acid-derived SAMs. [84]

Composite coatings

Usually, different kinds of coatings are combined to fabricate composite coatings on magnesium and its alloys because these possess combined properties of enhanced corrosion resistance and biocompatibility.

Magnesium fluoride (MgF2) coating and hydroxyapatite coating have been widely applied on surface modification of magnesium and its alloys. By combining these two types of coatings, Bakhsheshi-Rad et al. synthesized nano-hydroxyapatite/ magnesium fluoride (nano-HA/MgF2) coating and dicalcium phosphate dehydrate/ magnesium fluoride (DCPD/MgF2) composite coating via fluoride conversion process followed by electrochemical deposition on magnesium alloy.[85] The root mean square roughness of the nano-HA/MgF2 and DCPD/MgF2 composite coatings was approximately 395 and 468 nm, respectively, which is higher than that of fluoride treated and untreated samples. The needle-like HA crystals had a diameter of 80–150 nm and a length of about 7 μm and the plate-like DCPD was relatively larger. The experimental results showed that the composite coatings reduced the hydrogen evolution and improved the nucleation site of apatite compared with that of the uncoated sample . Using the chemical conversion coatings as inner layers, Kunjukunju et al. fabricated multilayered coatings of alginate and poly-l-lysine on alkaline- and fluoride-pretreated magnesium substrate using a layer-by-layer technique[86]. Moreover, surface functionalization of these coatings by chemical crosslinking and fibronectin immobilization was realised to control the cellular activity of these multilayered films . The studies showed that is more advantageous to use the fluoride conversion as pretreatment method because better bioactivity and less cytotoxicity were obtained compared with the hydroxide pretreatment. Although imparting good biocompatibility of the modified surface, the multilayered coatings of alginate and poly-L-lysine did not alter the degradation kinetics of the substrates and it is the pretreatment conditions that had a significant impaction on the overall coating degradation behavior. It was proved that the corrosion resistance of magnesium could be significantly improved by MAO coatings. Despite that fact, the porous structure of MAO coatings limits their long-term protection for magnesium substrate. Also considering the inadequate biocompatibility, the MAO coatings may be not proper being used alone as modification surface of magnesium-based implants.. Guo et al. fabricated a composite MAO/poly-l-lactic acid (MAO/PLLA) coating on the surface of WE42 alloy[87] . The PLLA coating effectively sealed the microcracks and micropores on the surface of MAO coating by physical interlocking to inhibit the severe attack of corrosive fluid. Moreover, the MAO/PLLA composite coating endowed magnesium substrate with good cytocompatibility . Also based on MAO coating, Liu et al. formed a calcium phosphate coating on its surface by chemical method to construct MAO/Ca–P composite coatings . The outer calcified coating was composed of calcium-deficient HA and DCPD. Some new apatite formed on the calcified coating surface after SBF incubation. Compared with PEO coating only, the composite coating increased the corrosion potential and decreased the hydrogen gas release to present better corrosion resistance.[88,89]

Compared with single kind of coating, the composite coatings indeed have some advantages . Usually, one coating plays a role as anti-corrosion layer and the other one plays a role as biofunctional layer. The composite coatings have more potential in the biomedical field due to their enhanced multifunction. Although rare investigations have been reported, potential dangers of coating dropping off may be caused because more interfaces are introduced in formation of composite coatings.

Calcium orthophosphates

The main reason for using calcium orthophosphates as bone substitute materials is due to their chemical similarity to the mineral component of mammalian bones and teeth. Furthermore, in addition to being non-toxic, they are biocompatible and don't cause side effects in the body and, most importantly, both exhibit bioactive behavior and integrate into living tissue by the same processes that are active in remodeling healthy bone. This leads to an intimate physicochemical bond between the implants and bone, termed osseointegration. Moreover, they are also known for supporting osteoblast adhesion and proliferation. Despite these facts, the major limitations for using of calcium orthophosphates as load-bearing biomaterials are their mechanical properties: they are brittle with a poor fatigue resistance. Because of that, calcium orthophosphates are used primarily as fillers and coatings for biomedical applications. In table 5 are presented the most known calcium orthophosphates with their major properties. Even more thorough information on calcium orthophosphates can be found in specialist books and monographs. [58]

Table 5. Existing calcium orthophosphates and their major properties

a These compounds cannot be precipitated from aqueous solutions.

b Cannot be measured precisely. However, the following values were found: 25.7 } 0.1 (pH 7.40), 29.9 } 0.1 (pH 6.00), 32.7 } 0.1 (pH 5.28). The comparative extent of dissolution in acidic buffer is: ACP_a-TCP_b-TCP > CDHA_HA > FA.

c Stable at temperatures above 100 _C.

d Always metastable.

e Occasionally called ‘‘precipitated HA (PHA)’’.

f The existence of OA remains questionable

Also, calcium phosphate-containing layers are used in biomedical application as bone substitution and orthopaedic materials, due to the formation of a hydroxyapatite (HA) layer, which is similar to the mineral phase of bone. During processing it is important to adjust phases. The resulting layers are often mainly amorphous, but they contain some crystallized HA and also other calcium phosphate phases. The result of a direct reaction with the base material is represented by conversion coatings which contain Mg compounds. Through the methods of obtaining calcium phosphate-containing coatings is also found the immersion in SBF; this process is often termed bio-mimetic if carried out at 37 șC and a pH of 7.4.Various compositions of surface layers depending on the bath solution were reported by Rettig and Virtanen, including amorphous carbonated calcium/magnesium phosphate layers which formed after immersion in SBF solution for 5 days. Those layers had a thickness of P20 lm, and were seen to be highly permeable [90]. In a related research, Lorenz et al. showed the way is obtained a mixed calcium/magnesium phosphate layer on pure magnesium on soaking in SBF solution, which initially increased the survival of human HeLa cells compared with Mg surfaces after simple soaking in NaOH solution [91]. The main objective of the study is represented by an investigation of the surface roughness, which was increased by the SBF treatment. It should be emphasized that have been developed different composition of SBF solution, and considering that the formation of the calcium phosphate is a precipitation process depending on the solubility of the calcium phosphate phase formed, different types of coating can be formed using different SBF solutions. Ca and P containing layers were also produced via immersion into SBF, and after that were included various oxidizing pre-treatments by Jo et al. [92]. Surface treated samples were tested in direct contact with MC3T3 E1 cells and showed a significant improvement in cell attachment.

Various studies have been carried out to control the phase composition of calcium phosphate coatings. For example, Hiromoto and Yamamoto synthesized HA on a magnesium surface without pre-treatment by immersion in a solution of Ca-EDTA, KH2PO4 and NaOH at various pH values at 110 șC, the immersion times varied between 6 and 24 h to control the coating thickness [93]. Using polarization techniques the authors noted an increase of the corrosion resistance with growing coating thickness, indicated by a reduction in the corrosion current density. Phase analysis of the coatings showed that besides HA other phases, including Mg containing phases, were present, even though an attempt was made to suppress the formation of magnesium hydroxide and magnesium phosphate by adding high concentrations of Ca ions to the conversion solution.

Gray-Munro and Strong produced a coating on Mg alloyed with Al and Zn in three stages [94].The pre-treatment involved firstly, passivation via immersion in NaOH solution, which is followed by a thermal treatment with the aim to reduce corrosion during the following SBF conversion coating process. The main objective of their study was characterization of the coating phases. Again, the coating showed, besides calcium phosphate, Mg phases, and it was mostly amorphous, containing only small amounts of crystallized HA. The results imply that nucleation and growth of HA were catalyzed by dissolution of Mg from the substrate. Coatings of <3 µm thickness exhibited many cracks, as revealed by SEM analysis: after 3 h immersion the coating was seen to contain cracks due to corrosion, while the thicker coatings after 24 and 96 h showed cracking due to dehydration.

In another research, made by Chen et al. it was used calculated equilibrium diagram in order to obtain a stable HA coating using a calcium nitrate and sodium phosphate solution. Nevertheless, a post-treatment in alkaline solution was necessary to develop a HA component within the coating. The HA–Mg(OH)2 coating produced improved the corrosion resistance of the Mg substrate.[95] In vitro and in vivo tests made on Mg–Mn–Zn alloy substrates were included in a research of Xu et al.[96] . Firstly, was applied an alkaline pre-treatment, which was followed by an immersion in calcium phosphate solution. After that, coatings containing mainly CaHPO4•2H2O (DCPD) and were fabricated small quantities of Mg2+ and Zn2+ . The surfaces appeared to be porous. The study compared samples with and without a calcium phosphate coating using titanium as a reference. The number of cells on the coated surfaces was similar to the number of cells on the reference, which were significantly higher than on surfaces without a coating. For in vitro studies 18 rabbits were used to compare the implantation of coated and uncoated Mg alloy samples. A study of the optical cell density on histological cross-sections showed that the coating disappeared after 4 weeks, after which time period the coated and uncoated samples showed no significant differences. This investigation indicates that by adjusting the calcium phosphate phases and by using alloys instead of pure magnesium as the substrate can be achieved different improvements. Without using pre-treatment, the amorphous calcium/magnesium phosphate coatings on AZ31 alloy sample were produced by Yang et al. [97]. It was necessary a thermal post-treatment at 300 șC after immersion in sodium phosphate, sodium carbonate and calcium nitrate solution. The samples, coated and uncoated, as well as a degradable polymer as a control group, were implanted into nine rabbits. After 8 weeks the coated implants showed a slower biodegradation rate, confirming the positive effect of the phosphate coatings.

Pre-deposition procedures

Commonly, the surface of Mg and its biodegradable alloys needs to be prepared before to be coated by calcium orthophosphates. The preparation normally consists of cleaning and/or degreasing to remove any sort of surface contamination arising from manufacturing.

Among the environments that can be performed preparation procedures are the following: acetone, ethanol, mixtures thereof, trichloroethylene, an aqueous solution of Na2CO3, a mixture of 4% nitric acid+96% ethylene glycol or distilled water. In addition, various types of physical modifications of the metallic surface are used; examples include physical grinding and/or polishing, drying, heat treatment and/or autoclaving.

Furthermore, prior to deposition of bioceramics, the surface of Mg and its alloys might be chemically treated e.g. activated alkaline treated , anodized , chemically polished , electrochemically polished , etched, passivized , pre-phosphatized. More to the point, prior to deposition of calcium orthophosphates, the surface of Mg and its alloys can be coated with an interlayer of another compound, such as poly (e-caprolactone) , nicotinic acid , Mg(OH)2 , MgF2 , Ca(OH)2 or titania to enhance the corrosion resistance and coating flexibility. Pre-calcified coatings can also be applied . All these types of treatment are usually performed by dipping, spraying, rinsing and/or soaking, depending on both the quality requirements and the limitations of the product to be coated. Before the deposition of calcium ortophosphates, the surface of Mg and its biodegradable alloys need to be sterilized. Among the the most popular surface pretreatment techniques appear to be grinding and/or polishing of Mg and its biodegradable alloys and cleaning and/or degreasing.[58] There are many types of materials and methods, but the coatings with calcium phosphates (including HA) appear to be more promising . Mg alloys surface pretreatment influence the quality of the coating and adhesion to the substrate. Also, the Mg alloys are used as antibacterial coatings.

Coatings for magnesium alloys-methods

Biomaterial surface can be modified by various methods to overcome or reduce their inherent shortages or disadvantages like their degradation rate or low biocompatibility, such that the surface modification not permanently change the surface structures and properties, such as leading to non-degradability of implants or toxicity to surrounding tissues. For meeting the clinical requirements of bone defect substitution and repair and for improvement of the biomaterials, the cutting edge techniques were introduced to improve surface physical, chemical and biological properties of bone grafts. Therefore, various surface modification techniques were developed to improve their tribological properties. Magnesium and its alloys with no treatment usually cause severe hemolysis due to their rapid degradation. This subchapter covers recent advances in coating, non-coating and patterning techniques for the surface modification of bone graft materials. Some commonly used surface modification techniques and their effects are shown in the following picture.

Fig. 4 Schematic diagrams of surface modification techniques.

In order to control the degradation rate of magnesium-based implants to maintain their mechanical strength as well as reduce the side effects during their service time, many new magnesium alloys have been designed especially for biomedical applications by adding alloying elements. Mg–Ca , Mg–Zn , Mg–Sr and Mg–Ag alloys have been developed by adding single nutrient element or antibacterial element. Mg–Zn–Zr, Mg–Zn–Ag, Mg–Ca–Sr, Mg–Zn–Mn and Mg–Nb–Zn–Zr etc. complex alloys have also been developed for better performance in mechanical properties and biocompatibility [52]. Compared with commercial magnesium alloys, such as AZ31 alloy and WE43 alloy, although these newly developed magnesium alloys possess better mechanical properties, corrosion resistance and biological performance, alloying alone does not reach the requirement of corrosion resistance due to the high electronegative potential of magnesium (2.4V with respect to hydrogen electrode) and its poor passivation tendency. Moreover, the inhomogeneous microstructures in magnesium alloys may cause localized corrosion, which will cause accelerated following corrosion process. So it is critical to minimize the localized corrosion of magnesium based implants, especially during the initial stage of post-implantation, to maintain enough strength to support injured tissues with minimum side effects. Moreover, the response of surrounding tissues to the implants is closely related to their surface properties . So for magnesium-based implants, proper surface corrosion resistance and good biocompatibility for surrounding tissues integration with the implants are critical for their applications. Surface modification of magnesium and its alloys is aimed to construct an anti-corrosion layer with good surface biocompatibility. Advantages and disadvantages of each method are also discussed to give a suggestion for their usage in different situations. Considering the existence of interfaces and failure mechanism of the modification layers, coatings, or called films and ion implantation can be divided into two main classes. Figure 5 depicts the schematic diagram of the failure mechanism of the coated (a) and ion implanted (b) magnesium substrate during corrosion process.

Fig. 5 Schematic diagram illustrating the corrosion failure mechanism of surface modified magnesium and its alloys: (a) coated magnesium substrate and (b) ion implanted magnesium substrate.

For obtaining a modified surface, the most simple and intuitive approaches are the coating techniques. A variety of conventional physical and chemical coating methods (e.g. solvent evaporation, plasma spraying, and physical/chemical vapor deposition (CVD)) have been industrialized, while novel approaches involving recently developed and developing techniques are continuously coming forth. What will be introduced in this section are state-of-the-art physical and chemical methods for creating functional coatings on different biomaterial surfaces (see Figure 6) .

Fig.6 Classification of methods for obtaining coatings

Physics methods

Physical Vapor Deposition

The physical vapor deposition technique (PVD) is based on the formation of vapor of the material to be deposited as a thin film. The material in solid form is either heated until evaporation (thermal evaporation) or sputtered by ions (sputtering). In the last case, ions are generated by a plasma discharge usually within an inert gas (argon). It is also possible to bombard the sample with an ion beam from an external ion source. This allows to vary the energy and intensity of ions reaching the target surface. In Figure 7 is presented the way for obtaining coatings through PVD method.

Fig.7 Process of Physical Vapor Deposition

Ion implantation

Ion implantation represent a process in which ions are accelerated and they affect the modified surface.This method supply the possibility to introduce different species into a substrate independent of thermodynamic restrictions like solubility. This technique introduces a suitable quantity of ions into the near surface of the materials to change the surface properties, such as biocompatibility and corrosion resistance. Unlike surface coatings, an ion implanted layer does not have an sharp interface and layer delamination may be not a serious issue. Ion implanted surface is very thin and because of that it does not provide enough protection for a long period. In spite of that , for delaying initial corrosion, ion implantation is feasible and effective. In conformity with the implanted elements, ion implantation for magnesium modification can be classified as gas ion implantation, metal ion implantation and dual ion implantation.

Gas ion implantation

Gas ion implantation can introduce inorganic elements, such as nitrogen and oxygen into magnesium substrate. For example, Wan et al. used oxygen plasma immersion ion implantation for control of the degradation rate of magnesium substrate[141]. The results of their investigation showed that although the treated sample presented increased corrosion resistance against neutral PBS, they could not resist the more aggresive chloride ion enriched PBS. The high level of corrosion resistance is considered to ascribe to increased Mg–O bonding states created on the surface layer of magnesium and more homogenous surface morphology due to the ion bombardment effects. The intensification of corrosion resistance was reduced in chloride ion-enriched PBS because the Mg-O bond can be dissolved asily in Cl− enriched and more acidic environment. Tian et al. utilized nitrogen plasma ion implantation on AZ31B alloy [142]. The improvement of corrosion resistance after nitrogen implantation was considered to be attributed to the compactness of the loose natural oxide layer and ion irradiation effect. Severe surface sputtering and possible formation of a small amount of Mg3N2phase might have an adverse effect.

For improvement of corrosion resistance, Tian et al. also used water plasma ion implantation and oxidation for magnesium alloys[143] . The oxide layer composed of a native oxide layer and oxidization layer caused by water implantation. The oxygen content in the layers increased because of increasing treatment time and voltage. By using the proper water implantation and oxidation conditions, the corrosion resistance could be easily improved. Thus, the formation of a compact oxide is due to the improved corrosion resistance. Regarding the modified surface is mostly composed of magnesium oxide, water implanted surface may be also easily affected by chloride ion rich environment.

Xu et al. recently investigated the influence of carbon dioxide plasma immersion ion implantation on the electrochemical properties of AZ31 magnesium alloy in physiological environment[144]. Carbon dioxide PIII has formed in the surface of AZ31 alloy a surface layer with carbon in the graphite state and an oxide film composed of magnesium oxide and aluminum oxide. Thus, the corrosion resistance was improved especially in Dulbecco's Modified Eagle Medium (DMEM).

Metal ion implantation

Gas ion implantation enhances the corrosion resistance of magnesium and its alloys especially by passivating their surface to form an oxygen- or nitrogen-rich layer. With distinct mechanism, metal ion implantation can introduce metallic elements into the magnesium substrate to form surface alloying. In biomedical field, the biological properties and toxicity of the alloying elements must be considered. As it is known, aluminum (Al) and zinc (Zn) are frequent used as alloying element in AZ-based alloys which have achieved a success in industry. However, after Zn implantation, the degradation rate in simulated body fluids was increased significantly. The degradation rate in simulated body fluids was increased significantly after Zn implantation. The X-ray photoelectron spectroscopy (XPS) results showed that a thin Zn rich surface layer with Zn existing in the metallic state was produced by ion implantation, which assigned to the reduction of corrosion resistance due to the galvanic effect [145]. Opposite with the results of Zn ion implantation, Al ion implantation appreciably enhanced the surface corrosion resistance of pure Mg as well as AZ31 and AZ91 alloys. This increase could be attributed to the constitution of a gradient surface structure with a gradual transition from an Al-rich oxide layer to Al-rich metal layer revealed by XPS depth profile[146]. Besides, considering the potential danger of Al, the Al ion implantation must be carefully chosen for biomedical application. Zirconium (Zr),tantalum (Ta) and titanium (Ti) are biologically friendly to the human body as their metal or oxide implants have been applied clinically. So Ti, Zr and Ta have been experienced as implanted ions into magnesium and its alloys. For example, Liu et al. conducted Ti and Zr ion implantation for AZ91 magnesium alloy[147]. The surface layers revealed a typical intermixed layer composed of a external surface mostly composed of titanium or zirconium oxide with a small quantity of MgO and Mg(OH)2, an intermediate layer containing metal oxide and metallic implanted particles, and a base layer rich in metallic elements. The corrosion resistance of AZ91 alloys was improved with the implantation of Zr and Ti ions. Wang et al. discovered that Ta ion implantation could also increase the corrosion resistance of AZ31 alloy[148]. Ta2Al was found to be prepared in the modified layer. The formation of a pre-oxidation layer with a double structure of the dense MgO layer and the protective Ta2Al border could describe the mechanism for a better corrosion resistance the implanted samples.

Dual ion implantation

In general, it is hard to avoid oxidation when the samples are exposed to air or an oxygen-containing ambiance. Some results also indicate that metal ion implantation can results in O-rich outer layer which may help to the improvement of corrosion resistance for magnesium substrate. Considering the higher corrosion resistance of metal oxide, production of an oxide layer on magnesium surface is achievable and can be fabricated by metal and oxygen dual implantation. Zhao et al. employed aluminum and oxygen dual ion implantation to change the surface of magnesium alloy[149] . The results indicated Al and O ion implantation produced an Al2O3-containing protection layer. The modified layer enhanced the corrosion resistance of the substrate and localized corrosion became the principal corrosion mechanism instead of general corrosion . Because of the protection of the modified Al2O3 layer, the plasma-treated implant deteriored more slowly and simultaneously stimulated bone constitution in vivo in a minimal invasive way without leading at post-operative complications [150]. Oxygen and titanium dual ion implantation produced a TiO2-containing film which also notably improved the corrosion resistance of magnesium alloy[151] . ZrO2-containing surface film was produced on magnesium alloy by zirconium and oxygen dual ion implantation. Corrosion resistance, in vitro biocompatibility and even antimicrobial properties were improved. [152] Xu et al. also developed a thicker surface oxidized layer composed of chromium oxide by chromium and oxygen dual ion implantation[153]. The layer which was obtained could successfully delayed the surface degradation of pure magnesium [154]. In simulated body fluid and sodium sulfate, the chromium and oxygen dual ion implanted magnesium both had a reduced corrosion rate and exhibited less pitting corrosion .

Ion implantation and plasma immersion ion implantation

Ion implantation is a physical surface modification process that supposes the accelerated injection of high-energy ions into the surface of a material with purpose to modify its physicochemical and biological properties. Ion implantation is available for almost all elements from the periodic table. Ion implantation could be used on biomaterial surface for enhancing corrosion resistance, to reduce wear debris, regulate hardness and for improvement of biocompatibility and bioactivities. For example, to enhance corrosion resistance, iridium was implanted into TI-6Al-4V surface [155]; for reducing surface wear is used nitrogen ion implantation into Ti–6Al–4V and UHMWPE [156]; ion implantation of silver into surfaces of titanium, 317L stainless steel and Ti–Al–Nb alloy has role in increase their anti-bacterial natures [157]; graphene could attain a good cytocompatibility by NH2 ion implantation [158].

Ion implantation has also been used for surface modification properties of polymers for biomedical polymers. Among advantages of ion implantation methods are following: material of the substrate is unlimited, because of forcibly injection of the high-energy ions into substrate surface; the implanted ions are dispersed within a certain depth of substrate surface without forming a new layer, avoiding disadvantages (e.g. cracking and detachment) of traditional coatings; low performing temperature (sometimes at room temperature) did not affect the substrate material. Despite these advantages, ion implantation is a line-of-sight processing technique, and is not suitable for the treatment of those bone implants with complex shape and internal structure, like hip joint prosthesis with complex curved surface [159].

Rapid prototyping

Rapid Prototyping (RP) by layer-by-layer material deposition, started during early 1980s and comprises a series of techniques using three dimensional computer-aided design (CAD) data to quickly fabricate a model or duplicate a same part. Some of RP techniques were used for construction of coatings for biomaterials, especially for metallic biomaterials. Laser engineered net shaping (LENS) is an additive RP manufacturing technique that uses a focused, high-energy laser beam to melt metallic powders directly injected to the focused laser beam spot to form a new layer. By using a LENS process, Balla et al coated titanium with tantalum to improve the osseointegration property [100]. Graded Co–Cr–Mo alloy coating was also successfully created on porous Ti6Al4V surface by LENS to obtain a high hardness interface . LENS is also able to prepare ceramic coating, not just metallic coating. Roy et al. successfully fabricated calcium phosphate coating on titanium without phase transition of the ceramic coating. [101]

Pulsed laser deposition

One of physical vapor deposition (PVD) methods is pulsed laser deposition (PLD), which is a popular method for fabricating calcium phosphate coating on metallic substrate, since it is able to stoichiometrically transfer material from target to substrate and could obtain a ultra-thin coating layer (thickness of several atoms) [102]. Although PLD has been introduced to the surface modification of biomaterials for nearly 20 years, this technique is continuously developing in the field. For example, in its recent development, PLD was used to fabricate calcium phosphate coating on porous Ti6Al4V substrate produced by selective laser melting (SLM, one of RP techniques); water assisted PLD was developed to improve coating-substrate binding strength [103]. In recent years, PLD has also been introduced to the surface coating on polymers. Prosecka et al. fabricated thin layer of hydroxyapatite (HA) on caprolactone/polyvinyl alcohol composite nanofibers [104]. Besides calcium phosphate coating, bioceramic coating composed of akermanite (Ca2MgSi2O7) was successfully created on both non-biodegradable polysulfone and bioresorbable polylactic acid (PLA) surface by PLD. [105]

Ion beam-assisted deposition

Ion beam-assisted deposition (IBAD), which is also called ‘ion beam enhanced deposition’ (IBED) is a vacuum deposition surface modification technique that combines PVD and ion implantation (described in the section ‘Ion implantation and plasma immersion ion implantation’).

During the process of IBAD, the ion beam bombardment is continuous such that to realize cleaning of the substrate surface prior to the deposition and control depositing film properties during the deposition. A significant advantage of IBAD is that such technique is able to create a gradual transition layer mixed with substrate material and depositing material between the substrate and the deposited film, thereby the coating adheres strongly to the substrate. It is important to know the difference between IBAD and some other surface modification techniques that also use ion beam and have similar names, including ion beam deposition (IBD), ion beam induced deposition (IBID) and ion beam sputtering deposition (IBSD). IBD is a direct beam deposition (DBD) process that directly applies an ionized particle beam onto substrate surface to fabricate thin film [106]. The significant difference between IBD and ion implantation is the difference between the energies used by methods such that the ionized particle beam in the IBD has low energy, and the particles arrive at substrate surface with a ‘soft landing’ .[106]

IBID is a CVD technique which for decomposing gaseous molecules and for depositing non-volatile component onto substrate surface uses focused ion beam (usually Ga+ ion beam). IBSD is a PVD process that an ion beam bombards a target and ejects particles in atomic scale from the target to form thin film on nearby substrate surface . To create thin film coating on material surface can be used one of IBD, IBID and IBSD methods, even if they are hardly to create gradual transition between the substrate and the deposited film as IBAD, thus obtaining relatively lower adhesive strength. IBAD has been used for the surface modification of biomaterials for decades and is still in development. [107] Cui et al. fabricated HA coating on Ti–6Al–4V substrate with an atomic intermixed coating/substrate interface by IBAD. The substrate surfaces were cleaned by Ar+ ion beam bombardment, before the deposition [108]. After the deposition, a composite target containing tricalcium phosphate and HA was sputtered by Ar+ ion beam to form the coating on the substrate, which was simultaneously bombarded by another energetic Ar+ ion beam. In the deposition process, for producing atomic intermixed layer of the coating and the substrate, at first the bombardment energy of Ar+ ion beam was relatively higher, and then the bombardment energy was reduced to increase the thickness of the coating and reinforce the compactness.

The fact that the temperature of the substrate was below 100ș C during the deposition, will not affect the substrate. The results of the testing of coating fabricated by IBAD showed that the adhesive strength was nearly twice to that prepared by IBSD with the same processing environment. With IBAD method, Chen et al. Prepared calcium phosphate thin film coating on pure titanium and further eated biomimetic apatite precipitation layers by immersing the coating in Dulbecco’s phosphate buffered saline solutions containing calcium chloride, as well as biomolecules to modulate precipitation processes and enhance bioactivities[109,110]. IBAD was also applicable for the fabrication of metallic, bioceramic and composite thin film coating for many varieties of biomaterials for bone graft (e.g. titanium, stainless steel and ultra-high molecular weight polyethylene [UHMWPE])[111-113]. A drawback for IBAD is that it is a line-of-sight modification technique, so it is difficult for IBAD to treat an irregular surface with non-line-of-sight regions.

Plasma coating

Plasma is a state of matter that is partially or fully ionized, and contains charged particles of free ions, electrons, radicals, as well as neutral particles of atoms and molecules. Plasma could be thermal (high-temperature/hot/equilibrium) one and non-thermal (low-temperature/cold/non equilibrium) one.

Regarding the thermal plasma, it is nearly fully ionized and the electrons and the heavy particles have the same temperature . Usually, the temperature used to generate thermal plasma ranges from 4000 to 20 000K [114]. High temperatures like these is destructive for biomaterials, especially for those polymers. For non-thermal plasma, only a small fraction of the gas molecules are ionized, and ions and neutrals are at a much lower temperature (may as low as room temperature), although the temperature of electrons could reach several thousand degrees Celsius. For the surface modification of biomaterials is using the nonthermal plasma, which can be generated by different sources, including corona discharge, dielectric barrier discharges, radio frequency discharges and so on [115]. Plasma surface engineering is a series of economic and effective approaches for the surface modification of biomaterials and has been applied to commercialized products [116].To modify material surfaces via different processes like etching (or ablation), sputtering, polymerization, grafting and spray are used plasma treatments. Furthermore, to produce coating on a surface could be used plasma spray, plasma sputtering and plasma polymerization. Other plasma processes for the surface modification will be attributed to non-coating techniques.

Plasma spray

Plasma spray is a coating process that sprays melted or partially melted coating material onto substrate surface, and this technique has been applied to commercially available bone implants. Plasma surface engineering techniques descripted above has the energy and temperature of the plasma environment, relatively higher than plasma spray. Plasma spray operate at high temperatures. Due to this fact, it is usually applied to fabricate various coatings on metallic biomaterials, such as calcium silicate coating [117,118], apatite and its derivatives coating [119-121], bioglass coating [121-123], titanium coating [124], zirconia coating [125] and composite coating [126,127]. Moreover, plasma spraying was used to fabricate bioceramic coatings on polymer substrate, for example HA coating on polyetheretherketone (PEEK) or carbon fiber-reinforced polyamide 12 , and titanium coating on carbon fiber reinforced PEEK. However, the influences of the high-energy plasma on the polymer substrate were not discussed by these studies. The thickness of plasma sprayed coatings was usually more than 100 µm, and the interface could be clearly observed between the substrate and the coating. Therefore, the interface binding strength produced by plasma spray is relatively lower than those by IBAD with a gradient transition layer.

Plasma polymerization

Plasma polymerization is a process that ionizes monomer gas into plasma state and induces radical polymerization to create polymer coating on a substrate, so as to enhance corrosion resistance of metallic biomaterials or improve biocompatibility and bioactivity of relatively inert materials [128-130]. Lewis et al. applied a fluorocarbon film on 316L stainless steel and the results showed that the corrosion rate was significantly decreased compared with those uncoated [128]; Liu et al. used plasma polymerization to modify surfaces via generating different functional groups (amine, carboxyl, methyl and hydroxyl) and found that the plasma polymerization of allylamine on the surface promoted osteogenic differentiation of human adipose-derive stem cells. [131]

Chemical methods

Chemical covalent bonding

Commonly, for constructing chemical coating are used functional groups on material surface to form covalent bond between the substrate and the coating. The reaction is specific and binding effect is stable. Silanization is a low-cost and effective covalent coating method to modify material surface that are rich in hydroxyl groups, such as HA, bioglass, titania and many other metal oxide surfaces. There are many types of commercially available silane coupling agents, which are easy to react with hydroxylated surface and introduce active groups (e.g. amino group and carboxyl group) to the surface. Silanized surface can easily be modified by further grafting. Zhang et al. labeled nanometer HA with fluorescein isothiocyanate (FITC) by modifying HA with 3- aminopropyltriethoxysilane (AMPTES), and then grafting FITC via reaction with the amino group . Although the silanization is simple and effective, the reaction conditions such as concentration of the silane and reaction time must be carefully controlled to prevent from forming thick polymerized silane network on the surface. Otherwise, the bond between silane and the surface can also subject to hydrolysis in some conditions. [132,133] Chemical conversion treatment is an adequate method to form coatings on magnesium and its alloys to improve their corrosion rate. During the treatment process, the whole or partial contents of treatment solution can interact with the magnesium substrate to form magnesium compounds, which form a protective layer retarding the subsequent corrosion of magnesium substrate.

Because Mg(OH)2 can be formed as a mainly protective layer, for modification of magnesium-based materials can be used alkaline treatment. Zhu et al. fabricated a Mg(OH)2 film on AZ31 alloy using a treatment by 5.66 wt.% NaOH solution at 160°C. The corrosion rate of the magnesium alloy was inhibited by the protection of Mg(OH)2 film. During corrosion in Hank’s solution, it was observed that amorphous calcium apatite has been deposited on the surface of the film. The tape test showed a strong adhesion between the substrate and the film. The cytotoxicity test also revealed that no signs of changes on cell morphology or inhibitory effect on cell growth appeared for this kind of filmFor Mg-CA alloy treatment were also applied three weak alkaline solutions (Na2HPO4, Na2CO3 and NaHCO3), aside from alkaline solution containing NaOH. The corrosion rates of Mg–Ca alloy in simulated body fluid were effectively reduced after they were soaked in these solutions and subsequently heat treated, as suggest the following sequence: NaHCO3 heated < Na2HPO3 heated < Na2CO3 heated. Furthermore, cytotoxicity evaluation revealed that none of the alkaline heat-treated Mg–Ca alloy samples produced toxicity to cells. Without any chemical additions, for fabrication of Mg(OH)2 coating on AZ31 alloy, Ishizaki et al. only used ultrapure water. Vertically self-aligned nano- and microsheets with color expression were created at a temperature of 120°C with different treatment time . The color-tuned magnesium alloy revealed damping capacity and anti-corrosive performance . With subsequent modification with n-octadecyltrimethoxysilane, color-tuned superhydrophobic surfaces were successfully produced . Because the coatings formed by alkaline treatment mainly composes of Mg(OH)2, which is easily attacked by chloride ions to convert into highly water soluble MgCl2, the coatings formed by this method may not reach the required corrosion resistance in long-term service in chloride-rich physiological environment.

Fig.7 Surfaces of different coatings obtained through chemical covalent bonding

Surface and cross-sectional morphologies of Mg(OH)2 film by NaOH treatment for 3 h

(b) Digital photographs and FE-SEM images of color-tuned surfaces on AZ31 alloy by water treatment

(c) Surface and cross-sectional morphologies  of fluoride-treated AZ31 alloy for 72 h. The insert shows the high magnification

(d) Surface and cross-sectional morphologies of the MAO-coated Mg–Zn–Zr alloy

(e) Surface morphologies of Mg–Fe–CO3 LDH layers on pure magnesium by different treatment process

In comparation with Mg(OH)2, MgF2 is a much stable phase in physiological environment. Hydrofluoric acid immersion is frequently used to fabricate MgF2 coating on magnesium alloys by surface fluoridation (Fig. 2c). The effect of HF treatment of powder metallurgy Mg, cast Mg and AZ31 alloy was compared by Carboneras et al. . The MgF2 coating which has been formed on their surface, delayed the degradation rate of all the alloys, especially for cast Mg and AZ31, retarding the corrosion process in cell culture medium for at least a week. Over the degradation process, the fluoride-treated magnesium alloy preserved a better mechanical strength evaluated by three-point bending test, showing a promise for application as biodegradable implants. In in vivo test, a new bone created at the edges of the MgF2-coated magnesium implants and the degradation induced on the implant surface a calcium and phosphorous rich layer , which was coated by an incomplete layer containing fluoride The MgF2-coated implant proved a easy decrease in volume and better strength after 6 months implantation in comparation with uncoated implants. MgF2 coating reac hes a big success in modification of magnesium-based implants by in vitro and in vivo evaluations for biomedical field. Although, the hydrofluoric acid is usually used in MgF2 formation, which may be dangerous for operators and harmful to environment. So fluoride salt may be a favorable substitute and proper treatment parameters need to be researched. For improving the corrosion resistance of magnesium-based materials is used surface phosphorylation as well as surface bioactivity to form a phosphate-containing coating. For fabrication of a conversion coating on WE43 alloy by simple immersion treatment , Ye et al. used phytic acid (PA). Their research showed that the PA treatment could improve the corrosion resistance of the magnesium substrate and the cytocompatibility of the PA-coated WE43 alloy was significantly enhanced. Furthermore, the hemolysis ratio of PA-coated WE43 alloy was lower than 5%, which met the hemolysis standard of biodegradable materials. Taking into consideration that phosphorous is a type of nutrient element, especially for bone growth in human body, surface phosphorylation of magnesium-based materials is more indicated for orthopedic application. Except the conversion coatings mentioned above, rare earth conversion treatment on the surface of magnesium alloys is an environmentally friendly technique. Besides, a small quantity of rare earth element have not infuence on the human health based on the practice application of the rear earth containing magnesium alloys mentioned above. The corrosion behavior of cerium conversion coating on AZ31 magnesium alloy was investigated by Cui et al. in physiological solution. The formed dense Ce-based conversion coating is composed of a mass of trivalent and tetravalent cerium oxides. The coating could furnish obvious protection for magnesium substrate to efficiently reduce the degradation rate in Hank’s solution. Unlike the traditional technique, Levy et al. discovered a diffusion coating of Nd on Mg-1.2%Nb-0.5%Y-0.5%Zr-0.4%Ca alloy and examinated the effect of this coating on corrosion behavior of the alloy in simulated physiological electrolyte. By the protection of the diffusion coating, the corrosion resistance of the alloy was significantly enhanced. The obtained intensification of corrosion resistance was due to the formation of a relatively continuous network of passive Mg41Nd5 intermetallic at grain boundaries and the enrichment of the oxide film with Nd and Nd-oxide. Rare earth conversion coatings are mostly composed of the corresponding compounds. Because these compounds show better corrosion resistance, they can easily separate the magnesium substrate from corrosive fluids. Although the rare earth elements used have not been related to present obvious toxicity, the potential danger of these elements in long term should be observed. [52]

Plasma electrolytic oxidation coating

Plasma electrolytic oxidation (PEO), also called micro-arc oxidation (MAO), has been widely used to produce porous and robust coatings on biodegradable magnesium and its alloys; it is a combination of plasma-chemical and electrochemical processes.. This process is used to improve the corrosion resistance of magnesium-based implants. PEO combines an electrochemical oxidation with a high-voltage spark treatment carried out in aqueous electrolytic baths, which also contain modifying elements in the form of dissolved salts (e.g. silicates, borates) to be incorporated into the resulting coatings. This technique was found to be suitable for depositing ceramic coatings on Mg and its alloys with simultaneous corrosion resistance.

Yang et al. used this technique to fabricate a Mg2SiO4 containing coating on ZK60 magnesium alloy[98]. Besides the increase of the corrosion resistance of magnesium alloy, the MAO coating, also improve its in vitro biocompatibility. The results of investigation of MAO-coated alloy revealed no cytotoxicity and conduct to an increase of alkaline phosphatase level in comparation with that of naked alloy, indicating that the release of Mg and Si ions from the coating was beneficient for the differentiation of bone marrow stromal cell (BMSCs). The surface of MAO coating revealed a better cell adhesion and affinity for cells which directly grew on various surfaces, the surface of MAO coating exhibited a better cell adhesion and affinity. It was observed the major difference between the hemolysis ratio of MAO-coated alloy (1.04%) which was significantly decreased compared with the naked alloy (28.89%), which means a great improvement of the hemocompatibility. By adjusting the compositions of electrolytes, ZrO2 containing coating and Ca–P containing coating were also succesfully fabricated[99]. Ca- and P-containing coatings were also produced on AZ91D alloy in both NaOH- and Na2SiO3-containing electrolytes with addition of sodium hexametaphosphate and calcium hypophosphite. The results of energy-dispersive spectroscopy (EDS) showed that the coatings prepared in the NaOH system mainly contained oxides of Mg, Al, P and Ca, while those prepared in the Na2SiO3 system also were composed by a substantial amount of Si oxide. The results of X-ray diffraction revealed MgO and Mg2SiO4 to be the preponderant phases in the coatings prepared in the NaOH and Na2SiO3 systems, respectively. Constitution of calcium orthophosphates was not noticed in the coatings, except for Ref, and thus the micro-arc oxidation technique alone appears to be unsuitable for deposition of calcium orthophosphates on Mg and its biodegradable alloys. It might, nevertheless, be considered as a pre-deposition technique. Although, using other procedures, such as chemical precipitation or electrodeposition, calcium orthophosphate coatings could be deposited over micro-arc oxidative coatings. Mg–Zn–Ca alloys coated with such composite coatings were found to produce rapid precipitation of calcium orthophosphates from simulated solutions with simultaneous increase in the corrosion resistance of Mg and its alloys. Moreover, combinations of micro-arc oxidation (to improve the rate of corrosion resistance) with any suitable deposition technique for calcium orthophosphates appear to be effective and promising methods for constructing biodegradable Mg-based metallic bone grafts offering improved patient outcomes. In comparation with others techiques, PEO may be more favorable to produce coatings on magnesium-based materials because the obtained coatings are hard and highly adhered to the substrate. Furthermore, the easy reglation of electrolytes which produces the adjusting of the coatings , is promising in biomedical applications because antibacterial elements or nutrient elements can be introduced into the coatings. In spite these advantages, the existence of some cracks or surface pores formed during PEO process makes the PEO coatings not to be so satisfactory in anti-corrosion in long-term service as corrosive fluids can penetrate into the holes and cracks,and thus, significantly reduces the corrosion resistance of the coating. The main phase composition of PEO coating is MgO, which is not favorable for cells growth. So is needed to be improved the surface biocompatibility of this kind of coating. So the PEO coating is required to be further modified to obtain enhance corrosion resistance as well as surface biocompatibility. So the PEO-based composite coatings are promising in practice and the progresses will be also reviewed in the following part.

Photografting and radiation grafting

For obtaining a stable surface modification results for biomaterials has been widely used the chemical grafting method. The active groups (e.g. –OH,–COOH and –NH2) which are exposed to the surface are necessary to acquire high chemical reactivity for the grafting. Because on the surface of the bioinert materials are only a few or no active groups exposed to their molecular surface, it is difficult to realize chemical grafting.

Despite that, many relatively inert materials are being used as bone implants. In order to perform grafting on the surface of these biomaterials, must be introduced to the grafting reaction an extra energy. As the name imply, radiation grafting and photografting, make use of radiations, including UV radiation (photografting), high-energy electron beam and gamma radiation. By breaking chemical bonds on material surface, radiation form free radicals. After this, the reactive surface will be exposed to monomers for initiate surface graft polymerization.[134] In the field of biomaterials, the use of photografting and radiation grafting, is concentrate on surface modification of hydrophobic and bioinert polymers. Many materials commonly used in the preparation of bone substitutes, like PEEK, UHMWPE and some biodegradable polymers, have been studied with purpose to modify chemical and physical properties and also improve biocompatibility and osseointegration using radiation grafting and photografting . For improvement of the tribological properties of UHMWPE, increasing hydrophilicity of PEEK and biodegradable polymers and to adjust biodegradation rate of PLA was used photografting method by UV radiation. Gamma radiation grafting was reported to graft poly(N-isopropylacrylamide), a polymer had low critical solution temperature onto the surface of polystyrene Petri dish to control attachment and detachment of cells. Also, Cho et al. used gamma radiation for the modification of surface of UHMWPE by the graft polymerization of methyl methacrylate (MMA) monomer for improvement of interfacial strength with poly (methyl methacrylate) (PMMA) bone cement. [134-138]

Plasma etching and grafting

Besides preparing coatings on biomaterials, plasma is also used for conducting various non-coating surface modification processes, like plasma etching and plasma grafting. By plasma etching a surface is modificated by shooting a high-speed stream of plasma onto the substrate. Plasma etching has an important role in improvement of surface activity for bioinert polymers, and less effect on surface topography than chemical etching process. For modifying surface chemical properties of biomaterials is used plasma grafting method by grafting active groups on the surface, for example, for obtaining an antibacterial surface is realized plasma grafting of zinc oxide onto polypropylene .[139,140]

Photolithography with mask

Photolithography was the first surface patterning method introduced with purpose to make patterns for controlling cell behaviors [160]. The fabrication of DNA arrays represents a well-known application of photolithography from the field of biology [161,162]. A photolithography process commonly comprises following steps :

preparing of a flat and clean substrate;

onto the substrate is added a light sensitive polymer (called photoresist);

exposure of the photoresist under a mask (usually quartz or metal) to form a desired pattern;

transfer of the pattern to the substrate by an etching process (development process)

remove the photoresist.

The photoresist can be positive or negative: at the positive one, that areas exposed to the light beam can be dissolved in the development process and the negative one works conversely . The resolution of the pattern created by photolithography is limited to about half of the wavelength of the light source, usually several hundred nanometers , because of the optical diffraction of the focused light beam. The finest resolution used by photolithography is approximately 1µ. Therefore, photolithography is convenient to create patterns with the size comparable to a cell. [163]

Direct-write photolithography

Compared to classic photolithography with mask, direct-write photolithography is an indirect method for the fabrication of surface patterns. Over the years, a direct-write mode without using mask has been revealed via using a focused light beam to fabricate pattern directly on material surface [164-166]. In this type of photolithography , laser is more often used as the light source to supply high-intensity light beam for the fabrication of surface patterns. For fabrication of patterns, the light beam may induce two types of reactions on material surface: for photoactive surfaces are a photochemical one, and the other is a physical one that implies ablation, melt or deformation of the substrate caused by the high energy of the laser [167-169]. In the next paragraph are presented some examples of using direct-write photolithography: By using laser photolithography combined with UV radiation, Pfleging et al. created micro-patterns on polystyrene surface for enhancing L929 cells adhesion and protein adsorption[170,171].

Rebollar et al. used this method to fabricate submicro-patterns on polystyrene substrate to guide the alignment and improve cell proliferation [172].

With pulsed laser, Ahrem et al. made 3D channels on bacterial cellulose hydrogels without chemical strength lose or chemical modifications, and observed inward migration of chondrocytes in the channels, as well as good matrix production and phenotypic stabilization[173].

Electron beam lithography

Electron beam lithography (EBL) has a similar principle and fabrication process as photolithography and is used to create nano-sized patterns on material surfaces. Because the electron beam is considered as a de Broglie wave, the wavelength of the electron is much shorter than that of light beams (e.g. a typical electron beam provided by electron microscope is accelerated by 100kV electric field, and the wavelength is 0.003 nm)[174]. With EBL are made patterns on electron sensitive material surfaces, which would be chemically transformed, cross-linked, polymerized and so on under the electron beam.

Idota et al. obtained patterns using polymerized and grafted N-isopropylacrylamide onto a hydrophilic polyacrylamide-grafted glass surfaces, so as to control cell attachment directions and detachment. Besides all the advantages, EBL has its own limitations like high cost and low throughput. EBL is now more studied in laboratory rather than industrialization applications [175,176]. Also, special patterns which gave outstanding functions to material surfaces. Using EBL technique, Wang et al. constructed a patterned surface with sub-micron sized polyethylene glycol (PEG) microgels [177]. Because the PEG microgels were non-adhesive to both cells and bacteria, the patterned surface would be non-adhesive to bacteria with comparable size to the PEG microgels, as well as did not affect adhesion and behaviors of normal cells. Such special effects were further researched and many similar multi-functional patterned surfaces were developed. [178-180]

Scanning probe lithography

Scanning probe lithography (SPL) is a direct-write method that shift a micro- or nano-stylus on material surface to mechanically ‘write’ patterns. SPL can be classified into two different types according to the patterning manners: the one is constructive that matters are transferred to the surface from the stylus (such as dip-pen nanolithography (DPN)); and the other one is destructive so that the surface is deformed (such as nano-imprinting/engraving). To obtain patterns by directly writing on the surface utilizing a variety of molecular inks (solutions of molecules), DPN uses a tip of atomic force microscope. In the biomedical field, DPN has been used to create patterns with biomolecules (including proteins, peptides, enzymes, lipids and DNA) [181], polymers [182], nano-particles [183], also living cells [184], onto different substrates. The throughput of DPN is relatively low comparing with other nanopatterning techniques mentioned above . To overcome this real disadvantage, a high-throughput DPN was evolved by parallel operating a 2D probe arrays consisted of 55 000 tips [185].

For create patterns, nano-imprinting and engraving uses a hard stylus to indent or scratch material surface. Thus, the process is destructive to the surface. By using this process, nano-patterns were created on thin film (such as self-assembly monolayer) coated surface. Moreover, nano-patterns of proteins were fabricated by grafting protein molecules onto the exposed substrate after the scratching .

Patterning with master

To replicate patterns on a substrate, patterning with master technique uses a template with patterns. The process is also called ‘microcontact printing’ (µCP). Commonly, a mold made of elastomeric polymer like polydimethylsiloxane (PDMS) , is created according to the master, and then used for printing patterns on the substrate with molecular inks. In this patterning technique, the mold has a connection with the substrate for realize the transfer of the patterns onto the surface of the molecular inks. Since 1990s, Singhvi et al. used a PDMS stamp to create patterns on a golden substrate and to control cell distribution and shape. Thus, they demonstrated that such spatial restriction is useful to maintain albumin secretion, which is an important physiological function of hepatocytes. The regulation of cell fate is another example for patterning with master. Using µCP, Kilian et al. created patterns with different shapes on a glass substrate coated by gold . Thus, the cells displayed different adipogenesis and osteogenesis profiles, after a period of growing mesenchymal stem cells (MSCs) on the surface, indicating that modulation of cell shape was able to direct cell differentiation. Patterning with master has also been applicated to patterning proteins , DNA , and cells for biomedical field.

Another process that can be utilized to create patterns on the substrate is imprinting with master, which uses a template made of hard material. Therefore, the master for imprinting must be hard enough to produce patterns on a relatively soft substrate. By combination between imprinting with master and lithography, nanoimprint lithography (NIL) creates nano-patterns on a substrate by imprinting a mold with the nano-patterns into a resist coated on the substrate and subsequent etching process. NIL technique was firstly developed in 1996 by Prof. S. Chou and his coworkers . A mold of nano-patterned silicon dioxide was prepared by EBL and etching, and after that it was realized the transfer of the pattern to a PMMA resist coated on a silicon substrate by identing to form patterned resist with distinct thickness, followed by anisotropic etching to finally transfer the pattern onto the substrate. By using of a nano-template to create pattern , NIL process has relatively higher throughput. NIL technique has been used to produce nano-patterns on various substrates for biomedical applications, or fabricate polymers, DNA and proteins patterns on the substrates. [186]

Self-assemble of molecules or nano-particles

Self-assemble is a process used to prepare patterns by an energy-saving mode. During the self-assemble process, intermolecular or inter-particle forces make molecules or nano-particles to arrange in a regular pattern, in order to minimize total free energy of the entire surface. Nano-particles or molecules can self-assemble in an area ranging from nano-scale to micro-scale. Various types of materials were used in self-assemble fabrication of patterns, like nano-spheres [187,188], nano-particles [189], biomolecules [190,191], block copolymers [192] and others.

3D patterning

3D patterns are suitable on biomaterial surfaces because the cells live and act in a 3D physiological environment in natural tissues , and such would provide spatial arrangements closed to physiological conditions and helpful for tissue reconstruction and repair. Therefore, 3D patterning technique is becoming an area of interest for surface modification. Two-photon lithography (TPL) and multiphoton lithography (MPL) are direct-write technique that is able to create 3D patterns on polymeric surfaces by laser beam. During a TPL or MPL process, two-photon or multiphoton is absorbed at a photosensitive surface by attaining energy from the laser beam, thereby chemical reactions (usually polymerization) take place to form 3D patterns at the laser spot. , Nielson et al. fabricated 3D patterns with bovine serum albumin by photocrosslinking using MPL on a coverglass according to high resolution X-ray computed tomographic data. By using a dynamic mask and a resolution as high as submicron (~0.5µ), the 3D patterns were exactly replicated. Using MPL could also be produced replicated patterns, cross-linked protein with unstrained structures in micro-scale. In biomedical field, MPL can be used to obtain micro-patterns or 3D structures with hydrogels, proteins, bioabsorbable polymers, gelatin and others [193-202].

Interference lithography (IL) uses interference patterns formed by two coherent laser beams to build periodic 3D patterns. It is also called holographic lithography or interference holography. The features of the interference patterns could be modified by adjusting parameters of the coherent laser beams like phase, amplitude, polarization . The interfering laser beams can be used to lead polymerization reactions to crate 3D patterns on a substrate. The simplicity of patterning process (without mask) is a major advantage and the throughput is relatively higher than other patterning processes. However, the alignment of the coherent beams is complex, and any modifications for the patterns need to simultaneously adjust both beams.

IL technique has been used to fabricate patterns on materials for biomedical applications even if in recent years it was relatively widely used in the field of microelectronics or optoelectronics . For example, using laser IL , Prodanov et al. produced nano-grooved surfaces with different features on titanium .

Magnesium-based layered double hydroxide

Magnesium-based layered double hydroxide (LDH) is also introduced into surface modification of magnesium and its alloys . For example, Lin et al. fabricated highly oriented Mg–Fe–CO3 LDH coatings directly on pure Mg by a two-step treatment, i.e. treatment in pH 5.6 aqueous Fe3+/HCO−/CO32− at 50°C and then immersing it in pH 9.5 aqueous HCO−/CO32− at 50°C [203]. The first step was performed to yield Mg2+ in aqueous solution by corroding the Mg substrate. A two-layered slim film was thus formed on Mg substrate with a external layer comprised fine plate-like Mg–Fe–CO3 LDH. After the latter treatment in pH 9.5 aqueous HCO−/CO32− at 50°C, the fine LDH platelets grown into a strong-oriented Mg–Fe–CO3 LDH. The Mg–Fe–CO3 LDH-coated sample had a much higher corrosion resistance than the pure Mg substrate. Moreover, the LDH coating showed a better cell spreading and cell–cell interaction behavior than the pure Mg substrate. Because the composition of Mg containing LDH is restricted by choice of trivalent metal anions and cations , magnesium-based LDH is recently tested on surface modification of magnesium-based materials. In spite of that , studies about magnesium-based LDHs are rarely reported until now, the application of this method is promising as the compositions of LDHs have potential to be adjusted by the exchange of anions, which is beneficial for enhancing bio-functional surfaces.

Biomimetic deposition and wet-chemical precipitation

The easiest way to deposit calcium orthophosphate coatings on Mg and its biodegradable alloys, from the preparation point of view, is spontaneous precipitation upon exposure to supersaturated solutions. Regarding the experimental conditions, spontaneous precipitation can be classified into biomimetic deposition and wet-chemical precipitation. With the purpose to imitate natural conditions , biomimetic deposition is realized from artificially prepared simulating solutions such as Hank’s balanced salt solution (HBSS) , simulated body fluid (SBF) and modifications thereof , while wet-chemical deposition is produced from solutions both of simpler compositions and those containing non-biomimetic ions[58] . Furthermore, chemical deposition can also be carried out from non-aqueous solutions, such as alcohol . The afore mentioned biomimetic solutions HBSS and SBF , as well as Dulbecco’s Modified Eagle’s Medium (DMEM) , minimum essential medium (MEM) and aqueous solutions (0.8–3.5 wt.%) of NaCl , are usually used to study the in vitro corrosion of Mg and its alloys , and thus the protective properties of calcium orthophosphate coatings . The latter property can be measured in vitro by hydrogen evolution, enhance in solution pH or by following the concentration of released Mg2+ ions . It is harder to measure the in vivo corrosion kinetics. To continue the distinctness between the biomimetic deposition and wet-chemical precipitation, the former experiments are always effectuated under physiological conditions (temperature and solution pH), while the latter can be performed at elevated temperatures and non-physiological solution pH . However, in all cases, calcium orthophosphates are precipitated and grown on the surface of Mg and its alloys. Both methods are simple to set up and perform and are a cost-effective way of developing homogeneous coatings on several samples simultaneously. In addition, they do not need line of sight and thus allow complex shapes to be coated . Depending on both the Ca/P ratio, temperature and the solution pH, coatings of DCPD, octacalcium phosphate (OCP) or calcium-deficient hydroxyapatite (CDHA) can be deposited on Mg and its alloys. For example, variations in the ionic composition of the initial solutions were found to lead to deposition of various phases such as DCPD, DCPD + CDHA, and CDHA .The sane behavior is observed for temperature: with increasing temperature, the intensities of the DCPD diffraction peaks in calcium orthophosphate coatings were found to decrease, while those of CDHA progressively increased . In addition, precipitation of β-tricalcium phosphate (β-TCP) on Mg was identified. Since simulating solutions often contain a number of different ions, ion-substituted calcium orthophosphates are always deposited via biomimetic approach. For example, amorphous carbonated calcium–magnesium orthophosphate coatings were formed in SBF and after incubation in SBF for 5 days, the coatings obtained had a thickness exceeding 20 µm and appeared to be highly permeable . Furthermore, the presence of highly reactive Mg results in the formation of Mg-enriched calcium orthophosphates with a concomitant reduction in crystallization kinetics because Mg2+ ions are known inhibitors of apatite nucleation and growth . Many investigations have showed that extended immersion in supersaturated solutions results in constitution of thicker coatings only if the supply of calcium and orthophosphate ions is plentiful. Under these conditions, biomimetic deposition requires regular renewal of the solutions . For example, managing the biodegradation rate of Mg by biomimetic apatite coating was examinated. The authors used both single-coated (one immersion for 24 h) and dual-coated (two immersions for 24 h each) Mg samples in modified SBF, as well as uncoated Mg as a control, followed by corrosion tests performed in standard SBF[58]. The disclosures demonstrated that two immersions resulted in formation of thicker CDHA coatings with increased corrosion protection in SBF comparative to the controls. The researchers has deduced that the degradation rate of Mg could be adapted by controlling the thickness of apatite coatings . Zhang et al. formed a homogenous bone-like apatite coating successfully on pure magnesium using a biomimetic method without the need of heat treatment. The corrosion rate of the magnesium implants could be closely tailored by adjusting the apatite coating thickness[204] . In biomimetic process, variation of preparation parameters could lead to a big difference in finally formed coating. For example, the variation of ionic composition of the initial solution led to the deposition of coatings with various phase composition, i.e. DCPD, DCPD + HA, HA and the magnetic field influenced the particle morphology and crystal texture of the precipitates [205].

Other scientists proved comparable discoveries for corrosion protection of calcium orthophosphate-coated AZ31 and AZ91D Mg alloys using both immersion tests in SBF and electrochemical tests , respectively. Also other investigators illustrated the corrosion protection of biomimetic CDHA-coated pure Mg with further increase in cell adhesion. Moreover, the biomimetic precipitation of calcium orthophosphates on Mg can be achieved under a magnetic field. A distinction in particle morphology and crystal texture of precipitates near the two poles was noticed. In the presence of a magnetic field, an increase in crystallite sizes in the (020) and (040) planes was remarked for precipitated DCPD, which permitted the authors to fluctuate crystallinity of the coatings . To conclude this part, spontaneous precipitation techniques have proven to be acknowledged methods for coating Mg and its alloys. However, there are reviews on the formation of non-uniform and porous calcium orthophosphate coatings on Mg substrates caused by coating formation occurring around hydrogen bubbles formed on the Mg surface during immersion .

Sol–gel preparation and dip coating

Sol–gel preparation combined with the dip-coating technique has been widely explored to coat Mg and its biodegradable alloys for both corrosion protection and enhanced adhesion. The technique requires immersing (dipping) a substrate into a liquid, which is a concentrated calcium orthophosphate solution with a gel-like texture. Requirements for the sol preparation are calcium and phosphorus precursors and one or two solvents, often ethanol (as the only solvent) or water and ethanol (if two solvents are used). The phosphorus predecessor, usually P2O5 or triethylphosphate, is dissolved in ethanol. The chosen calcium precursor, most frequently calcium nitrate, is also melted in either water or ethanol and then both solutions are mixed. The resulted mixture is then refluxed at various temperatures and solvents are evaporated off until a more viscous solution is obtained to achieve a sol–gel. Samples of Mg and its alloys to be coated are then immersed into the sol–gel several times to achieve a calcium orthophosphate coating, which is then cured at high temperatures to rise the coating/substrate adhesion and to perform apatitic structures within the applied coatings. When adapting this coating system for Mg substrates and their alloys, the preservation temperatures cannot exceed the melting point of pure Mg (650 șC) to avoid affecting the surface integrity of the Mg substrate. Therefore, conservation of sol–gel coatings on Mg and its alloys has been related in the range 25– 400 șC by various authors . A typical experimental example looks as this. 3.94 g of Ca(NO3)2•4H2O and 0.71 g of P2O5 were dissolved separately in 10 ml of ethanol. Ca-predecessor was added dropwise into the P-precursor to obtain a mixture containing a Ca/P ratio of 1.67. The resulted suspension was agitated at 400 rpm for 5 h at 26 șC in a closed beaker. Subsequently, the samples of an Mg alloy were dipped vertically into the suspension and withdrawn at a constant speed using an electric dip coater [58]. The coated substrates were then kept at room temperature for 24 h in order to complete aging. Then, they were progressively heated to 60 șC and maintained at this temperature for 24 h. Then, the coated samples were calcined at 400 șC for 6 h . To date, sol–gel preparations combined with dip coating have been established as cost-effective and simple systems to set up, as well as having the ability to coat irregular shapes, similar to spontaneous precipitation techniques. This method shows the potential of using relatively low temperatures and short incubation times to realize thick coatings on Mg and its alloys. Their primary advantage over the before mentioned spontaneous precipitation techniques is in the strength of the coating/substrate adhesion .

Electrodeposition

Electrodeposition is a wide field of deposition techniques involving electrical current. Electrodeposition is usually used to deposit Ca–P-based coatings on magnesium and its alloys. It contains electrophoretic deposition, electrochemical deposition (or cathodic deposition), and some other techniques. Usually, electrodeposition is a low-cost and simple procedure that can be performed at room temperature to obtain uniform coatings, and has also been applied for deposition of calcium orthophosphate coatings on Mg and its biodegradable alloys. This procedure is regularly carried out from aqueous solutions similar to those used in wet-chemical deposition . For example, an AZ91D alloy was successfully coated by a biphasic combination of DCPD and b-TCP using cathodic deposition at room temperature for 2 h. Then, a transformation step was carried out in 1 M aqueous NaOH for 2 h to transform the biphasic mixture into uniform CDHA coatings . Other scientists conducted cathodic deposition of calcium orthophosphate coatings on AZ31 alloy at 85 șC during 4 h and obtained a combination of CDHA and DCPD phases. Subsequently, the coated samples were dipped in 0.25 M NaOH solution for 4 h at 60 șC to convert the biphasic (CDHA + DCPD) coatings into the single-phase CDHA ones . In these studies was demonstrated that the deposited CDHA coatings both enhanced corrosion protection (namely, the numerical value of corrosion potential (Ecorr) of AZ31 substrate increased from ~1.6 to ~1.42 V) and raised the bioactivity of the samples. Moreover, the CDHA-coated samples exposed to SBF showed ~20% improvement in mechanical strength in comparation to that of the uncoated samples . The electrodeposition process of calcium orthophosphate coatings on Mg and its alloys can be changed to comprise pulse currents as opposed to constant currents. This resulted in better performance of the pulse-potential coatings, which was attributed to a more closely packed morphology of the protective coatings, decreasing the anodic dissolution of Mg and its alloys. Regulation of both the pulse current parameters and electrolyte solutions were suggested as effective ways to maintain the coating structure. Thus, CDHA was successfully deposited on Mg–Zn–Ca alloy directly without the need for a transformation step. Moreover, dense and uniform fluorine-doped HA coatings were electrodeposited the same way but with addition of H2O2 to the electrolyte to decrease H2 evolution by oxidizing this species . In addition, pulse electrophoresis at 85 șC for 30 min showed bigger adhesion of calcium orthophosphate coatings to Mg–Zn–Ca alloys, whilst pulse treatments of the coated samples of AM50 alloys at 20 șC for 15 min exposed a better corrosion resistance . Other investigators utilized cathodic deposition to coat DCPD onto Mg–Zn alloys. The DCPD coatings were deposited at room temperature by adding 10 ml l_1 H2O2 into the electrolyte and adjusting the pH value to 4.4. The authors established that the DCPD phase could surely reduce the degradation rate of the alloy as well as enhance its biocompatibility. Subsequently, the researches indicated that the electrophoretically deposited DCPD on Mg–Zn alloys (which were then transformed to CDHA via alkali treatment) is more stable and efficient in corrosion resistance in comparation to the originally deposited DCPD. Almost the same results were obtained by other researchers . Besides, coatings of fluoridated HA were found to have even better corrosion protective properties [15,124]. Following research analyzed the effect of different deposition times (20,60, 120 or 240 min) on cathodic deposition. The authors deducted that the deposition time had a high impact on the morphology of the coatings; still, it did not affect the conversion of DCPD to CDHA or other calcium orthophosphates [58]. Normally, electrodeposition of calcium orthophosphate coatings on Mg and its biodegradable alloys is achieved from aqueous solutions (see above); however, to amend certain properties of the coatings, studies are available on the addition of organic solvents, such as alcohol . For example, to decrease the conductivity of the coating solutions (0.1 M Ca(NO3)2 and 0.06 M of NH4H2PO4 dissolved in water), ethanol was added in different percentages, i.e. 10%, 30%, 50%, and 70% (v/v), and electrochemical deposition was carried out on AZ91 alloy using a constant-potential method. This investigation showed a considerable decrease in hydrogen bubble bursting during deposition from ethanol-containing solutions, which followed in denser packing of the precipitated DCPD crystals and, thus, a higher degradation resistance of the DCPD coatings, as compared to those obtained from ethanol-free solutions. The best results were achieved for 30% (v/v) ethanol- containing solutions, while further an increase in ethanol content in the solution produced thinner DCPD coatings offering poorer protection [58]. At 8 weeks of implantation, the area morphology of the residual bare Mg alloy implant displayed obvious corrosion pits , while there was only slight corrosion with superficial pits for the CDHA-coated specimen . While the implantation time get bigger , corrosion of both samples slowly became severe due to the local failure of the CDHA coatings. Therefore, both uncoated and CDHA-coated Mg alloy implants were corroded in the rabbit femora; however, during the same implantation period, the bare Mg alloy implants experienced more serious corrosion than the CDHA-coated implants [101]. Song et al. also used this procedure to obtain brushite (DCPD, CaHPO4·2H2O), hydroxyapatite (HA, Ca10(PO4)6(OH)2) and fluoridated hydroxyapatite (FHA, Ca5(PO4)3(OH)1−xFx coatings on Mg–Zn alloy and compared their corrosion behavior in modified simulated biological fluid. All of these coatings reduced the degradation rate of Mg–Zn alloy[58]. The precipitates on the uncoated and DCPD-coated Mg–Zn alloy in modified simulated biological fluid had low Ca/P molar ratio, which delayed bone-like apatite formation, while both HA and FHA coating could promote the nucleation of osteoconductive minerals for 1 month. Li et al. [52] discovered that the bone-like FHA also showed better cellular proliferation and differentiation than Mg–Zn substrate for human BMSCs (hBMSCs). In comparation with classical cathodic electrodeposition method, a modified pulse electrodeposition has also been applied to fabricate Ca–P-based coatings. Wang et al. [52] developed the deposition of soluble Ca-deficient hydroxyapatite (Ca-def HA) coating on Mg–Zn–Ca alloy substrate by pulse electrodeposition. With this coating protection, the corrosion resistance was significantly improved, thus the ultimate tensile strength and time of fracture for the coated Mg–Zn–Ca alloy were higher than those of the untreated one. Xu et al. [52] studied the porous and netlike DCPD coating formed on the surface of the Mg alloy. In vitro cell evaluation showed that L929 cells showed remarkably good adherence and proliferation on the Ca–P-coated magnesium alloy. In vivo evaluation results developed that the Ca–P coating provided magnesium with a significantly good surface bioactivity and promoted early bone growth at the implant/bone interface. Pure β-tricalcium phosphate (β-TCP) was also investigated by this method [52]. Tomozawa et al. [52] formed HA coatings uniformly on pure Mg by a hydrothermal treatment using a C10H12N2O8Na2Ca (Ca-EDTA) solution. This HA coating remarkably reduced the corrosion rate of the Mg in simulated body fluid (SBF). Moreover, the biological responses, including cell attachment, proliferation and differentiation, of the HA-coated samples were enhanced considerably compared with the pure magnesium. Preliminary in vivo experiments showed that the biodegradation of Mg implant was significantly retarded by this HA coating [52].

Hydrothermal

Hydrothermal treatment is an easy procedure and one of the most cost-effective techniques for coating deposition on metallic surfaces. It is almost similar to the afore mentioned biomimetic deposition and wet-chemical precipitation; however, since the hydrothermal treatment is carried out at elevated (>90 șC) temperatures over a relatively prolonged period of time (>1.5 h), the calcium orthophosphate deposits are usually crystalline. Nevertheless, it is hard to form coatings of pure calcium orthophosphates on Mg and its biodegradable alloys because aqueous solutions at elevated temperatures cause heavy corrosion of Mg, while the released Mg2+ ions might both form a surface layer of Mg(OH)2 and substitute Ca2+ ions in the structure of the coatings. [206]. besides, in 2012 and 2014, research groups from China [207] and Korea [208], also published papers on this procedure. Using a hydrothermal treatment for 2 h at 363 K, the Japanese researchers succeeded in obtaining well-crystallized HA and OCP coatings on both pure Mg and Mg–Al–Zn alloys from a 0.25 M Ca-EDTA and KH2PO4 treatment solution over a wide pH range from 5.9 to 11.9 According to the researchers, the constitution of highly crystalline coatings was achieved primarily due to using Ca-EDTA solutions, which could provide a sufficiently high concentration of Ca2+ ions to cause precipitation. Both HA and OCP coatings were found to consist of an outer porous layer and an inner continuous layer, while both the crystal phases and the microstructures of the coatings were found to vary with the pH of the treatment solutions. Namely, in small acidic (pH 5.9) solutions, a double-layer structure was obtained: an inner dense layer consisted primarily of HA crystals and an outer coarse layer composed of plate-like OCP crystals. In weak alkaline (pH 8.9) solutions, a dual layer structure was also created: an outer coarse layer consisted of rod-like HA crystals and an inner dense layer consisted of well-packed HA crystals. In strong alkaline (pH 11.9) solutions, needle-like HA crystals were found. Both layers were found to grow with an enhance in the treatment period. A thin Mg(OH)2 layer was also formed at the boundary between the calcium orthophosphate coatings and Mg substrates. The HA and OCP coatings were found to increase the corrosion resistance of both pure Mg and Mg–Al–Zn alloys in both HBSS and a 3.5 wt.% NaCl solutions; although, the corrosion resistance of HA coatings was always higher than that of OCP ones [58]. Other researchers achieved the same results [207,208]. Moreover, such coatings proved good adhesive properties with small plastic deformation under cyclic stress below the fatigue limit. Neither cracks nor detachment was microscopically observed under 5% static elongation and under 3% cyclic elongation [58]. The authors discovered that the level of protection allowed by the calcium orthophosphate coatings could be varied by changing their crystal phase, microstructure and thickness. Thus, optimization of the microstructure of the coatings is necessary to adjust the corrosion resistance of the coated Mg to the desired values.

Aerosol deposition

Furthermore, calcium orthophosphate coatings can be put down on Mg and its biodegradable alloys by an aerosol deposition technique. This procedure has been used to deposit HA onto the surface of both pure Mg and Mg previously covered by either poly(e-caprolactone) [210] or MgF2 [209]. To carry out aerosol deposition, HA powder was sprayed onto Mg samples in a deposition chamber using oxygen carrier gas at a flow rate of 5 x10-4 m3s-1 under a pressure of 9.2 Torr. Scanning electron microscopy (SEM) examinations showed that when HA was deposited onto Mg with the poly (e-caprolactone) interlayer, it was partially embedded into this interlayer, forming composite-like structures [210]; still, when HA was deposited onto Mg with the MgF2 interlayer, no composite-like structures were observed [209]. Corrosion tests performed in SBF revealed that such coatings had good corrosion resistance. In addition, HA coatings on Mg with the poly (e-caprolactone) interlayer were found to have better stability during deformation compared to HA coatings on Mg without the interlayer. These results revealed that coating Mg by an HA/ poly (e-caprolactone) double layer might prove a promising approach to reduce the corrosion rate of Mg and improve the coating flexibility [210]. Furthermore, using aerosol deposition, Mg and its biodegradable alloys can be coated by calcium orthophosphate-based biocomposites. For example, HA/chitosan biocomposites have been deposited on AZ31 alloy. The authors employed a slit-type nozzle with a 10×0.5 mm2 rectangular opening and air as a carrier gas with a flow rate of 30 l min-1. The 5 µm thick HA/chitosan coatings were deposited over the entire surface of the AZ31Mg alloy substrates by scanning the substrates on the motorized X–Y stage for 1 min at a scanning speed of 1 mm s-1. The biocomposite coatings were found to exhibit high adhesion strengths ranging from 24.6 to 27.7 MPa and showed good corrosion resistance. Although addition of chitosan lowered the corrosion resistance of the HA coatings, their biocompatibility was improved.

Spin coating

Only two publications on spin coating could be found, and both of these were devoted to deposition of calcium orthophosphate based biocomposites on Mg alloys [211,212]. Initially, biocomposites of HA/collagen (HAC) [211] and HA/poly (lactic-co-glycolic acid) [212], respectively, were prepared. Afterwards, to deposit HAC onto the surface of AZ31 alloy chips, a mixture of 2 g of poly(L-lactic acid) (PLLA) and HAC (at various PLLA/HAC ratios) was dissolved in 20 ml of dichloromethane and magnetically stirred for 30 min followed by ultrasonic dispersion for 15 min. The prepared suspension was spin coated on pretreated AZ31 alloy chips for 30 s at a rotational speed of 2000 rpm. The coated surface was immediately dried by blowing at room temperature and, in order to obtain thick coatings, the procedure was repeated five times. Corrosion studies performed in HBSS revealed that the biocomposite coatings suppressed the sharp increase in pH value and Mg2+ release from the substrates, while the degradation behavior of the alloy was correlated to the microstructure of the coatings [211]. Both the deposition technique and the obtained corrosion results appeared to be similar for Mg alloys covered by HA/poly (lactic-co-glycolic acid) coatings [212]. Thus, the spin-coating technique appears to be a convenient tool for deposition of composite coatings.

Spray coating

Only two publications were found on the topic of spray coating too. One of them was devoted to deposition of HA-doped poly(lactic acid) porous coatings on AZ31 alloy [213], while the other was devoted to deposition of HA coatings on AZ51 alloy [214]. In the former paper, the authors dissolved poly(lactic acid) in dichloromethane solvent. Afterwards, colloids were prepared by adding nanosized HA particles to the poly(lactic acid) solutions, which were stirred for more than 24 h. Then, these colloids were sprayed onto the surface of AZ31 samples at room temperature and 50% humidity. The air pressure (400 kPa) was optimized to produce fine droplets. The coated samples were dried in a vacuum oven at 40 șC for 12 h and then at 67 șC for 1.5 h [213]. In the second paper, the deposition technique was almost the same but the spraying process was performed onto AZ51 substrates heated to 400 șC [214]. The results of both studies revealed that the coated Mg alloy samples had better corrosion resistance compared to that of uncoated samples. During immersion tests performed for 15 days in HBSS, the numerical values of pH increase, weight loss and bending strength decrease were found to be lower for the coated samples (with average values of 8.5%, 7.2% and 10%, respectively) than the similar values for the uncoated samples (10.5%, 15.5% and 25%, respectively) [213]. Similar trends were obtained in another study [214]. In addition, cytocompatibility studies with MC3T3 cells revealed a continuous increase in cell growth on the coated samples [213]. There are a lot of methods, but some seem to have promising results and are intensely studied yet. Also, exist a great variability of tested material and because that, not always is used the same evaluation methods.

Coatings properties

Critical factors for the application of biodegradable coatings on Mg and Mg alloys for biomedical field are given below.

Surface chemistry

Specific surface chemistry necessities hinge on the targeted application, e.g. stents, orthopaedic implants, or tissue engineering scaffolds. The greatest interest considered by literature has been for calcium phosphate-containing and organic based coatings. There are also patents for calcium phosphate- containing coatings on medical implants. In spite of the great number of researches to date, the results in the literature proved difficulties in adjusting the required phases in the coatings.

Chen et al. [215] demonstrated in a study of chemical conversion coatings on Mg and its alloys that the coating pre-treatment seems to be more important than the choice of coating technology itself. Our analysis of the literature proposes that this can be applied to organic based coatings as well. Pre-treatment is necessary to operate the surface and to control the coating processes, mainly in aqueous solutions, as dissolution during the coating process can be a strong side-reaction.

Corrosion rate

Corrosion researches were performed in about 75% of the reviewed papers. The specific test methods, most frequently used were immersion tests, polarization studies and impedance spectroscopy. During the process, were varied the composition, concentration and volume of the electrolyte, as well as the test time and other parameters.

Analogy of the results is very difficult because of this variation in experimental conditions. All coatings were usually found to decrease the corrosion rate to a certain measure, as expected. Gu et al. [216] resembled the results of immersion tests, and concluded that when is applied PEO technique on Mg and Mg alloys, seems that it produce coatings more effective than others. Gu et al. [216] developed that PEO resulted in a more than 90% decrease in corrosion rate, whereas other coatings, e.g. alkaline and fluoride treated surfaces and organic based coatings, caused between about a 50% and 80% reduction. In spite these facts, even PEO coatings did not show a 100% reduction in corrosion rate. Indeed, the key function of coatings on Mg for biodegradable applications is to cause a short-term inhibition but not complete suppress corrosion of the material in a physiological environment. The most often used means to produce low corrosion rates of the coatings was the potentio-dynamic polarization method. This electrochemical procedure was carried out in SBF or NaCl solution, and corrosion protection in the presence of the coatings was determined on the basis of the corrosion current density icorr. The reduction in icorr for both types of coatings using SBF as the corrosion test electrolyte was in the range 1–3 orders of magnitude. These data propose that the corrosion rate is regulable. However, there is preoccupation regarding the uniformity and long-term nature of the corrosion process. Information on these issues cannot be revealed by carrying out polarization measurements alone.

In the articles of F. Witteet he analyzed corrosion of four types of magnesium alloys in the body. The alloy composition has important effect on the corrosion resistance of magnesium alloys, the corrosion resistance of AZ31 and AZ91 alloy which comprise elements such as Al; Zn is smaller than that of LAE442 alloy which contains rare earth element such as La, Ne. In his study it is showed that the corrosion resistance of LAE442 and AZ91D alloy had contrary behavior in vivo and in vitro: in vivo LAE442 had higher corrosion resistance and in vitro had an opposite, which showed that the effects of the alloy composition on alloy corrosion resistance in vivo and in vitro remained to be further clear.

The results of the studies of Yang Ke group also revealed that the addition of Mn and Zn can stimulate magnesium phosphate layer to produce in simulated body fluid, enhancing the corrosion resistance of magnesium substrate.

Pking University Zheng Yufeng group used the features of Ca element in human body, for it is one of the essential mineral elements and has light weight, to develop the Mg – Ca duplex alloy, and to control the corrosion resistance of the alloy by adjusting the content of Ca element.

In a research of Zhang et al. from 2009 it was reported that the corrosion rate of magnesium implants can be adequately controlled by covering them with well-adjusted thick apatite coatings .

The degradation behavior of calcium magnesium alloy in vitro SBF (simulated body fluid) was studied by M. Bobby Kannan and others and the results of the research showed that calcium magnesium alloy has higher corrosion resistance. Over the years were carried out investigations with the purpose of improving the osseointegration of implants by altering the chemical composition at the implant/surrounding tissue interface. The most related materials often used are tricalciumphosphate and hydroxyapatite. In addition to accelerating bone growth, tricalcium phosphate can enhance the corrosion properties of magnesium. In a study of Zhang et al. from 2009 it was related that the corrosion rate of magnesium implants can be pertinently controlled by covering them with well-adjusted thick apatite coatings.

During the same period Song et al. studied the transformation of β- tricalciumphosphate into uniform hydroxyapatite coating after immersion in 1M NaOH solution. Their results showed a high improvement in the corrosion resistance and bioactivity on the surface of AZ91D magnesium alloy immersed in simulated body fluid (SBF).

In order to overcome this limitation which is given by the corrosion rate of magnesium alloys, various techniques have been carried out on the surface protection of magnesium alloys, e.g. inorganic conversion coatings or plasma electrolytic anodizations, physical vapor deposition or thermal spraying.

There are generally two possible ways to improve the corrosion behaviour of Mg and Mg alloys:

adjustment of the composition and microstructure, including the grain size and texture of the base material;

carry out surface treatments or form coatings , which produce protective ceramic, polymer or composite layers.

With these coatings, biodegradation may be delayed; the delay is essential for a biodegradable implant that requires a fully functionality for a certain period of time time before the surgery region start healing.

Fig. 8 Corrosion layer of Magnesium

Until now they have not been developed coatings for magnesium alloys with applications in the medical field, most of coatings being developed for aircrafts components or mobile phone industry. The main objective of this chapter is the development of biocompatible and biodegradable coatings for Mg and Mg in order to reduce corrosion and to increase their initial biocompatibility. Infections that can occur in bone-implant interface are another reason to worry about. Given the antibacterial properties of silver, silver coatings are capable of addressing bacterial colonization on implants and other nosocomial infections during surgical procedures. The efficiency of silver coatings in controlling the spread of infections around an implant is dependent on the balanced of activities between the silver cation which kill bacteria and the concentration of silver ions released from the coating. A very high concentration of silver cations released from the coating can be toxic to cells . Silver coating on magnesium will form a galvanic pair which will accelerate corrosion of the substrate. Therefore, doping TCP with silver can control the galvanic corrosion of magnesium, while enhancing the osseointegration properties. Also, the right balanced of silver is required to maintain the corrosion protection of TCP, as well as control the toxicity effect due to high concentrations of silver. In a study from 2009, Won-Hoon et al. investigated the antibacterial properties of Ag containing calcium phosphate coatings formed by micro-arc oxidation on titanium implants. From their researches resulted that the calcium phosphate coatings obtained in the low Ag concentration electrolyte exhibited in vitro antibacterial activity but no cytotoxicity, in contrast to the coatings with high quantity of Ag. One of the directions of this research was to improve osseointegration and antibacterial properties of magnesium using TCP and silver, but also analyze the influence of Ag doped TCP coating on the corrosion properties of magnesium in simulated body fluids.

In the other studies were developed degradable and elastomeric polymer coatings applicable to blood-contacting, magnesium-based devices. The polymer coating had not only improved corrosion resistance and enhance the thromboresistance in acute blood contact, but also offered controlled release of bioactive agents to inhibit vascular hyperplasia. PCUU-coated magnesium stents exhibited improved corrosion resistance in vitro and reduced thrombotic deposition versus uncoated stents or those coated with PLGA (poly lactic-co-glycolic acid). Usually, in vitro and in vivo tests have related that most of the coatings developed can delay the start of corrosion. In spite of coating for the lifetime of the medical device, the right alloy choice is also significant as the coating will disappear with time. It has been presented that the accumulation of subcutaneous gas bubbles on the phases and microstructure of the magnesium alloy used and on the geometry of the samples rather than on the coating itself.

Adhesion

It is notable that studies on calcium phosphate-containing layers have not included adhesion tests. Even if there are processes proposed consisting of more than one step, the pre- and post-treatments are not taking in consideration as promoting adhesion. Those treatments were effectuated firstly to adjust the calcium phosphate phases, often with the aim of developing HA layers. However, some SEM studies have revealed the presence of cracks. Cracking can expand during coating due to corrosion taking place, and in other cases during dehydration during the drying process. Cracks are a denotation of low adhesion and may bring to sigmoidal degradation behavior of the coating. Roy et al. [52] analyzed the degradation behavior of calcium phosphate films, presenting that the layers were only stable for 3 days and did not supply sufficient protection against degradation due to the presence of pores and cracks. For organic based coatings pre-treatments are often performed to create a bond between layers. Often the bonding is merely physical interlocking of the phases. However, adhesion tests for polymer or composite coatings have rarely been carried out.

Coating morphology

The thickness of calcium phosphate based coatings extends over three orders of magnitude (from 0.2 to 200 µm). The thicknesses of organic based coatings are theoretically from couple nanometers up to hundreds of microns. Indeed, the suitable thickness for each application has to be meticulously selected. The surface topography will have a considerable impact on the corrosion behavior, as well as on cell adhesion. Most of the subjects discussed in this review used SEM to notice the surface. However, there is still a lack of information on the correlation between surface roughness, surface morphology, corrosion behavior and cytotoxicity. Nanostructuring and functionalization using proteins to enhance cell response, as is known for other materials, represent an attractive subject for future research to improve the biocompatibility of coatings on Mg and Mg alloys. Furthermore, specific requirements of the surface modification approaches depend on the targeted application, e.g. stents, orthopaedic implants, or tissue engineering scaffolds. For example, the mostly porous and rough surfaces after anodization are not appropriate for stents.

Conclusions

It clearly highlights that is recommended to use methods for evaluating the release of hydrogen, SEM and AFM for coatings morphology and the size of the deposited layers, XRD for structural characterization and the evaluation of the corrosion by electrochemical methods.

The major problem seems to be at this time, the evalution in time of corrosion products which are formed on the surface affter the interaction with the biological environment and the evaluation method of these.

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