Calcium orthophosphate coatings on magnesium and

Acta Biomaterialia xxx (2014) xxx–xxx

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Review

Calcium orthophosphate coatings on magnesium and its biodegradable alloys Sergey V. Dorozhkin ⇑ Kudrinskaja sq. 1-155, Moscow 123242, Russia

a r t i c l e

i n f o

Article history: Received 11 December 2013 Received in revised form 7 February 2014 Accepted 12 February 2014 Available online xxxx Keywords: Hydroxyapatite Calcium orthophosphates Magnesium Biodegradable alloys Coatings

a b s t r a c t Biodegradable metals have been suggested as revolutionary biomaterials for bone-grafting therapies. Of these metals, magnesium (Mg) and its biodegradable alloys appear to be particularly attractive candidates due to their non-toxicity and as their mechanical properties match those of bones better than other metals do. Being light, biocompatible and biodegradable, Mg-based metallic implants have several advantages over other implantable metals currently in use, such as eliminating both the effects of stress shielding and the requirement of a second surgery for implant removal. Unfortunately, the fast degradation rates of Mg and its biodegradable alloys in the aggressive physiological environment impose limitations on their clinical applications. This necessitates development of implants with controlled degradation rates to match the kinetics of bone healing. Application of protective but biocompatible and biodegradable coatings able to delay the onset of Mg corrosion appears to be a reasonable solution. Since calcium orthophosphates are well tolerated by living organisms, they appear to be the excellent candidates for such coatings. Nevertheless, both the high chemical reactivity and the low melting point of Mg require specific parameters for successful deposition of calcium orthophosphate coatings. This review provides an overview of current coating techniques used for deposition of calcium orthophosphates on Mg and its biodegradable alloys. The literature analysis revealed that in all cases the calcium orthophosphate protective coatings both increased the corrosion resistance of Mg-based metallic biomaterials and improved their surface biocompatibility. Ó 2014 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction Metals and their alloys play an essential role as biomaterials which can assist in the repair or replacement of load-bearing bones that have become diseased or damaged [1]. Due to their physical nature, the majority of metals have a high strength and a long service life combined with a low elastic modules and low plasticity at body temperature. In view of their chemical properties (corrosion resistance) and biological compatibility (lack of toxicity), the range of applicable implantable metals is restricted to stainless steels, titanium and its alloys (e.g. Ti6Al4V, Ti6Al7Nb and shape memory Ti–Ni alloys), tantalum, cobalt–chromium-based alloys, as well as some noble metals and their alloys (the latter used mainly for dental restoratives). The limitations of these metallic implants involve a possible release of toxic ions and/or particles through corrosion or wear processes. Furthermore, being xenogenic, all metals evoke a physiological response that results in formation of a fibrous capsule, thus isolating the implants from the body [2,3]. In addition, ⇑ Tel.: +7 4992554460. E-mail address: [email protected]

the mechanical properties of these metals and alloys are not well matched with those of bone, resulting in stress-shielding effects that can lead to reduced stimulation of new bone growth and remodeling, which decreases implant stability [4]. Finally, the above-mentioned metals and alloys are essentially neutral in vivo and remain as ‘‘permanent’’ fixtures. Therefore, if plates, screws and pins made of these metals and alloys are used to secure bone fractures, after healing they will have to be removed by a second surgical procedure [5]. Fortunately, there is a small group of biodegradable (also called bioresorbable or bioabsorbable) metals, which are able to degrade relatively safely within the body. The primary metals in this category are magnesium-based and ironbased alloys, although recently zinc has also been investigated [6]. Among them, magnesium (Mg) and its biodegradable alloys (AZ91, WE43, AM50, LAE442, etc.) appear to be the most promising. They can degrade naturally in the physiological environment by corrosion, and thus appear to be suitable candidates for the construction of temporary implants, including stents [7–13]. A few examples of such Mg implants are shown in Fig. 1 [14]. Metallic implants made from biodegradable materials, such as Mg and its alloys, possess some novel biomedical features. After

http://dx.doi.org/10.1016/j.actbio.2014.02.026 1742-7061/Ó 2014 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

Please cite this article in press as: Dorozhkin SV. Calcium orthophosphate coatings on magnesium and its biodegradable alloys. Acta Biomater (2014), http://dx.doi.org/10.1016/j.actbio.2014.02.026

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being implanted, they will slowly degrade, eliminating the necessity for subsequent surgeries to remove them, thereby accelerating the entire healing process with a simultaneous reduction in health risks, costs and scarring [15]. Nevertheless, to avoid various complications and undesired effects (Fig. 2A,B), a suitable degradation kinetics appears to be critical (Fig. 2C): any biodegradable implant must continue to perform its function(s) until the damaged tissues have been sufficiently recovered or healed [16]. Additionally, the degradation (corrosion) products of such implants must be well tolerated by both the surrounding tissues and the organism as a whole. Fortunately, Mg2+ ions are the fourth most abundant cation in the human body and are stored mainly in bones. They are vital to metabolic processes, a cofactor in many enzymes and a key compo-

nent of the ribosomal machinery that translates the genetic information encoded by mRNA into polypeptide structures [17]. Therefore, contrary to other implantable metals, the wear or corrosion products of which can be potentially toxic or otherwise harmful to patients, those of Mg might be potentially beneficial to patients [18]. As can seen from the above, Mg, its biodegradable alloys and their corrosion products are well tolerated by the human body. However, in the vast majority of the cases, the in vivo corrosion kinetics of Mg and its alloys exceeds that of bone healing (Fig. 2A,B); therefore, it must be slowed down for implant applications. Numerous investigations have shown that both the properties and functional activity of any implantable biomaterial can be influenced by surface modifications, such as polishing, oxidation, passivation, coating deposition, ion-implantation, etc. [19,20]. Of these techniques, the application of synthetic calcium orthophosphate coatings appears to be the most effective way of achieving surface modification, and moreover improves the biocompatibility and osteointegration of metallic implants.

2. A brief description of the two major constituents 2.1. Magnesium and its alloys Mg is the eighth most abundant element on the surface of our planet, making up 1.93% by mass of the earth’s crust and 0.13% by mass of the oceans. Mg is an alkaline earth element, which are located in the second group of the periodic table. All alkaline earth elements possess a very high chemical reactivity and form compounds with an oxidation number of +2. Therefore, they are not found free in nature. The first isolation of elemental Mg was performed by Sir Humphry Davy in 1808 [21,22]. In 1852, Robert Bunsen achieved viable commercial production of Mg by electrolysis, and Mg then began to be produced in small quantities in America and Europe, initially for pyrotechnical use and as igniting bands or wires for flashlights of the nascent photographic industry [23]. Mg and its biodegradable alloys are light in weight and low in density (1.738 g cm3 for pure Mg and 1.7–2.0 g cm3 for the alloys—values that are similar to the density of bones: 1.8– 2.1 g cm3). Due to its relatively low melting point (650 °C), Mg is considered a fusible metals. The elastic modulus of Mg is 45 GPa [24], which is much closer to that of bone (trabecular/ cancellous bones: 3–14.8 GPa, cortical bones: 18.6–27 GPa [25,26]) compared to compared to the moduli of other implantable metals: Ti alloys, 110–117 GPa; stainless steels, 189–205 GPa; Co–Cr alloys, 230 GPa. In addition, the numerical values of the compressive yield strength of bones and Mg are 130–180 and 65–100 MPa, respectively, while those of fracture toughness are 3–6 and 15–40 MPa m2, respectively. Hence, by using Mg and its alloys for bone grafting the stress-shielding effect can be mitigated [7,10]. In addition, some antibacterial properties of Mg have been reported [27]. From the chemical point of view, its high reactivity (the standard electrode potential of Mg2+ (aq) + 2e M Mg(s) is 2.37 V, that + of Mg(aq) + e M Mg(s) is 2.70 V and that of Mg(OH)2(s) + 2e M Mg(s) + 2OH is 2.69 V [28]) makes Mg readily soluble in body fluids, which is the primary reason of its in vivo biodegradability. Therefore, when Mg is exposed to aqueous solutions, the following oxidation reaction takes place on its surface [29,30]:

Mg þ 2H2 O ¼ MgðOHÞ2 # þH2 " Fig. 1. Biodegradable orthopedic devices prepared from Mg and its alloys: (top) bone plates; (middle) screws for orthopedic fixation; (bottom) a porous scaffold for bone void filling. Scale bar = 10 mm. Reprinted from Ref. [14] with permission.

This provides a possibility to measure the corrosion kinetics of Mg and its biodegradable alloys by the release kinetics of hydrogen (Fig. 3) [31]. Unfortunately, the oxidized surface layers consisting



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Fig. 2. A model explaining the improvements due to the presence of bioactive calcium orthophosphate coatings on Mg and its alloys. (A) A relatively rapid degradation rate of Mg might lead to formation of gaps at the interface. (B) A typical tetracycline label taken 14 weeks post-operation. (C) Protective calcium orthophosphate coatings can reduce the degradation rate and simultaneously ameliorate biocompatibility. (D) Corrosion-protective effects of calcium orthophosphate coatings measured via the H2 release rate and the change in pH value. Reprinted from Ref. [15] with permission.

Fig. 3. A schematic drawing of a H2 evolution collection set-up to follow the corrosion kinetics of Mg and its alloys. Reprinted from Ref. [31] with permission.

of hydrated forms of MgO and/or Mg(OH)2 are loose in nature and cannot provide sufficient protection to resist the corrosion encountered in the physiological environment, which contains a high level (104 mM) of chloride ions. The corrosion kinetics of Mg is strongly accelerated in the presence of dissolved chloride ions, which are able to convert the insoluble MgO + Mg(OH)2 coatings into a soluble MgCl2, simultaneously decreasing the protected area and promoting further dissolution of Mg [29,30]. Since further details on the corrosion process of both Mg and its biodegradable alloys are beyond the scope of this review, the readers interested in this topic are directed to the specialist literature on this subject

[12,13,29–33] (the same sources also summarize the information on available Mg alloys). For present purposes it is important simply to recognize that, to be feasible for orthopedic applications, the corrosion kinetics of Mg and biodegradable its alloys must be reduced. Ideally, it should be slowed down to allow the mechanical integrity of the metal to remain intact during bone healing. This would also minimize hydrogen production, which was observed as a disadvantageous by-product when using Mg [11,34]. There are a number of ways to improve the corrosion resistance of pure Mg [12,32,33,35]. Briefly, they comprise the following approaches: microstructure tailoring including both grain size [36,37] and texture [38], alloying [13,39–43], preparation of biocomposites [44–47], surface treatment [48–53] and deposition of protective coatings [43,47,54–65]. This review focuses on protective calcium orthophosphate coatings on Mg and its biodegradable alloys. It should be stressed 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 [66–77]. As the use of implantable calcium orthophosphates started in 1920 [78,79], while that of calcium orthophosphatebased biocomposites and hybrid biomaterials started only in 1981 [80], one can conclude that the latter is just at the initial stages and many more publications are expected in the near future. Finally, it should be mentioned that Mg and its biodegradable alloys can be also coated by calcium phosphate glass-ceramics [81,82]; however, that is another story. 2.2. Calcium orthophosphates The main driving force behind the use of calcium orthophosphates as bone substitute materials is their chemical similarity to


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the mineral component of mammalian bones and teeth [83–85]. As a result, in addition to being non-toxic, they are biocompatible, not recognized as foreign materials 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 osteointegration [86]. More to the point, calcium orthophosphates are also known to support osteoblast adhesion and proliferation [87,88]. Even so, the major limitations to the use of calcium orthophosphates as load-bearing biomaterials are their mechanical properties: they are brittle with a poor fatigue resistance [89,90]. That is why, in biomedical applications, calcium orthophosphates are used primarily as fillers and coatings [85,86,91]. A complete list of known calcium orthophosphates, including their standard abbreviations and major properties, is given in Table 1, while detailed information on calcium orthophosphates, their synthesis, structure, chemistry, other properties and biomedical application has been comprehensively reviewed recently [85]. Even more thorough information on calcium orthophosphates can be found in specialist books and monographs [92–97].

3. A brief description of the important pre- and post-deposition procedures Prior to be coated by calcium orthophosphates, in the vast majority of the cases, the surface of Mg and its biodegradable alloys needs to be prepared. The preparation normally consists of cleaning and/or degreasing to remove any sort of surface contamination arising from manufacturing. This procedure can be performed in acetone [31,56,57,63 ,64,66,69,70,73–75,98–119], ethanol [59,64,73,113–138], mixtures thereof [139], trichloroethylene [140], an aqueous solution of Na2CO3 [140], a mixture of 4% nitric acid + 96% ethylene glycol [129] or distilled water [31,64,66,68,70,73,74,113,129,131,132, 140,141]. In addition, various types of physical modifications of the metallic

surface are used; examples include physical grinding and/or polishing [31,56,59,63,64,66–75,98–147], drying [31,59,64,74,75, 98–101,110,113,115,117,120,121,123,125,131,132,135–138], heat treatment [120,122,140] and/or autoclaving [114]. Furthermore, prior to deposition of calcium orthophosphates, the surface of Mg and its alloys might be chemically treated (e.g. activated [67,99–101, 125,145], alkaline treated [56,67,129,130,140], anodized [114,136], chemically polished [102–109], electrochemically polished [139], etched [64,75], passivated [59,119,134,148], pre-phosphatized [67,128,136], etc. [120]). 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) [117], nicotinic acid [75], Mg(OH)2 [59], MgF2 [116,149], Ca(OH)2 [150] or titania [77] to enhance the corrosion resistance and coating flexibility. Pre-calcified coatings can also be applied [75,137]. 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. Finally, the surface of Mg and its biodegradable alloys can be sterilized prior to deposition of calcium orthophosphates [119]. As seen from the number of available references, grinding and/or polishing of Mg and its biodegradable alloys appears to be the most popular surface pretreatment technique, followed by cleaning and/or degreasing. In addition, after calcium orthophosphate coatings have been deposited, various types of post-deposition treatment can be used to improve their properties. For example, post-deposition heattreatment (annealing) of calcium orthophosphates leads to conversion of the deposited amorphous (ACP) and non-apatite phases, such as dibasic calcium phosphate dihydrate (DCPD), into either hydroxyapatite (HA) [113] or Ca2P2O7 [130] (depending on the Ca/ P ratio) with a simultaneous increase of coating crystallinity and corrosion resistance, combined with a reduction in the residual stress. Furthermore, for the same purposes, chemical treatment of the coated samples in either aqueous alkaline solutions [69,99, 129,134,142,144,148,151,152] or phosphate-buffered solution (PBS) [66,71,72] can be used (Fig. 4). In addition, the coated samples

Table 1 Existing calcium orthophosphates and their major properties [85]. Ca/P molar ratio

Compound

Formula

Solubility at 25 °C, log(Ks)

Solubility at 25 °C, g l1

pH stability range in aqueous solutions at 25 °C

0.5

Monocalcium phosphate monohydrate (MCPM) Monocalcium phosphate anhydrous (MCPA or MCP) Dicalcium phosphate dihydrate (DCPD), mineral brushite Dicalcium phosphate anhydrous (DCPA or DCP), mineral monetite Octacalcium phosphate (OCP) a-Tricalcium phosphate (a-TCP) b-Tricalcium phosphate (b-TCP) Amorphous calcium phosphates (ACP)

Ca(H2PO4)2H2O

1.14

18

0.0–2.0

Ca(H2PO4)2

1.14

17

c

CaHPO42H2O

6.59

0.088

2.0–6.0

CaHPO4

6.90

0.048

c

Ca8(HPO4)2(PO4)45H2O

96.6 25.5 28.9

0.0081 0.0025 0.0005

5.5–7.0

b

b

5–12d

85

0.0094

6.5–9.5

116.8 120.0 69

0.0003 0.0002 0.087

38–44

0.0007

0.5 1.0 1.0 1.33 1.5 1.5 1.2–2.2 1.5–1.67 1.67 1.67 1.67 2.0

Calcium-deficient hydroxyapatite (CDHA or Ca-def HA)e Hydroxyapatite (HA, HAp or OHAp) Fluorapatite (FA or FAp) Oxyapatite (OA, OAp or OXA)f, mineral voelckerite Tetracalcium phosphate (TTCP or TetCP), mineral hilgenstockite

a-Ca3(PO4)2 b-Ca3(PO4)2 CaxHy(PO4)znH2O, n = 3–4.5; 15–20% H2O Ca10x(HPO4)x(PO4)6x(OH)2x (0 < x < 1) Ca10(PO4)6(OH)2 Ca10(PO4)6F2 Ca10(PO4)6O Ca4(PO4)2O

a a

9.5–12 7–12 a

a

a

These compounds cannot be precipitated from aqueous solutions. 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. b



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Fig. 4. X-ray diffraction patterns of calcium orthophosphate-coated AZ91 Mg alloy before (top) and after (bottom) an alkaline treatment. One can see that the initially deposited coating of DCPD was transformed into that of CDHA. Reprinted from Ref. [148] with permission.

can be kept in boiling water [113]. Finally, to achieve even better protection, other compounds, such as stearic acid [69], can be adsorbed on or deposited over the calcium orthophosphate coatings to form biocomposites.

appears to be new. Nevertheless, the substantial number of publications available clearly indicates its importance.

4. Deposition techniques

From the preparation point of view, spontaneous precipitation upon exposure to supersaturated solutions is the simplest way to deposit calcium orthophosphate coatings on Mg and its biodegradable alloys. Depending on the experimental conditions, spontaneous precipitation can be divided into biomimetic deposition and wet-chemical precipitation. In order to mimic natural conditions, biomimetic deposition is carried out from artificially prepared simulating solutions such as Hank’s balanced salt solution (HBSS) [114,156], simulated body fluid (SBF) [122,137,155,157–159] and modifications thereof [59,141,150,160], while wet-chemical deposition is performed from solutions both of simpler compositions [31,54,56,109,115,119,122,125,126,129–131,136,137,140,143–145, 147,161,162] and those containing non-biomimetic ions [98,121, 128,133]. Furthermore, chemical deposition can also be performed from non-aqueous solutions, such as alcohol [64]. Interestingly, the aforementioned biomimetic solutions HBSS [31,48,59,73,74,77, 105,117,119–121,125,131–135] and SBF [54,55,64,68–72,75,76,

4.1. A brief historical background Very briefly, the history of biodegradable Mg metallic implants started in 1878, when the physician Edward C. Huse used Mg wires as ligatures to stop bleeding from vessels of three human patients: once in a radial artery and twice in an operation for varicocele. He had already observed that corrosion of Mg was slower in vivo and that the time taken for complete degradation was dependent on the size of the Mg wire used [153]. The history of biomedical applications of calcium orthophosphates is much longer, because attempts to treat rickets (rachitis) by these compounds have been known since, at least, 1797 [154]. Nevertheless, in spite of the long biomedical histories of both major constituents, the earliest papers on protective calcium orthophosphate coatings on Mg were not published until 2007 [54,155]. Therefore, the subject of this review

4.2. Biomimetic deposition and wet-chemical precipitation


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100,110–112,116,118,121,124,127,129–131,136,138,139,141–144, 147,152,162–169], as well as Dulbecco’s Modified Eagle’s Medium (DMEM) [119], minimum essential medium (MEM) [43] and aqueous solutions (0.8–3.5 wt.%) of NaCl [59,106,113,130], are commonly used to study the in vitro corrosion of Mg and its alloys [11,170], and thus the protective properties of calcium orthophosphate coatings [62]. The latter property can be measured in vitro by hydrogen evolution, increase in solution pH or by following the concentration of released Mg2+ ions (Fig. 5) [116,138]. It is much more difficult to measure the in vivo corrosion kinetics; nevertheless, typical results are presented in Fig. 6 [138]. To continue the differentiation between the biomimetic deposition and wet-chemical precipitation, the former experiments are always performed under physiological conditions (temperature and solution pH), while the latter can be performed at elevated temperatures and non-physiological solution pH [120]. Nevertheless, in all cases, calcium orthophosphates are precipitated and grown on the surface of Mg and its alloys. Both techniques are simple to set up and perform and are a cost-effective way of creating homogeneous coatings on several samples simultaneously. In addition, they do not require line of sight and thus allow complex shapes to be coated [171,172]. 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 precipitated on Mg and its alloys [115,147]. 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 [115]. Similar 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 gradually increased [147]. In addition, precipitation of b-tricalcium phosphate (b-TCP) on Mg was detected [120,128]. Since simulating solutions often contain a number of various ions, ion-substituted calcium orthophosphates are always deposited via biomimetic approach. For example, amorphous carbonated calcium–magnesium orthophosphate coatings were formed in SBF [155,157–159]. After incubation in SBF for 5 days, the deposited coatings had a thickness exceeding 20 lm and appeared to be highly permeable [157,158]. In addition, the presence of highly reactive Mg results in the formation of Mg-enriched calcium orthophosphates [140] with a simultaneous decrease in crystallization kinetics because Mg2+ ions are known inhibitors of apatite nucleation and growth [173]. Various investigations have indicated that prolonged immersion in supersaturated solutions results in formation of thicker coatings only if the supply of calcium and orthophosphate ions is plentiful. Under these circumstances, biomimetic deposition requires regular renewal of the solutions [59,114,122,137,141,155– 159]. For example, controlling the biodegradation rate of Mg by biomimetic apatite coating was investigated [141]. 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. The findings demonstrated that two immersions resulted in formation of thicker CDHA coatings with increased corrosion protection in SBF comparative to the controls. The authors concluded that the degradation rate of Mg could be tailored by controlling the thickness of apatite coatings [141]. Other researchers demonstrated comparable findings for corrosion protection of calcium orthophosphate-coated AZ31 and AZ91D Mg alloys using both immersion tests in SBF [163] and electrochemical tests [122], respectively. In support of this, still other investigators illustrated the corrosion protection of biomimetic CDHA-coated pure Mg with subsequent increase in cell adhesion [58,159]. Furthermore, the biomimetic precipitation of calcium orthophosphates on Mg can be performed under a magnetic field

Fig. 5. In vitro determination of the protective properties of calcium orthophosphate coatings on Mg and its alloys in SBF (corrosion tests). Variation of pH (top) and H2 evolution (middle) for bare and HA-coated Mg. Reprinted from Ref. [138] with permission. Concentration of released Mg2+ ions (bottom) for bare Mg, MgF2coated Mg and HA/MgF2-coated Mg samples. Reprinted from Ref. [116] with permission.

[115,174]. A difference in particle morphology and crystal texture of precipitates near the two poles was observed. In the presence of a magnetic field, an increase in crystallite sizes in the (0 2 0) and



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technique involves 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 precursor, typically P2O5 or triethylphosphate, is dissolved in ethanol. The selected calcium precursor, most often calcium nitrate, is also dissolved in either water or ethanol and then both solutions are mixed. The obtained 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 dipped into the sol–gel several times to acquire a calcium orthophosphate coating, which is then cured at high temperatures to increase the coating/substrate adhesion and to accomplish apatitic structures within the applied coatings. When adapting this coating system for Mg substrates and their alloys, the curing temperatures cannot exceed the melting point of pure Mg (650 °C) to avoid affecting the surface integrity of the Mg substrate. Therefore, curing of sol–gel coatings on Mg and its alloys has been reported in the range 25– 400 °C by various authors [64,77,176–181]. A typical experimental example looks as this. 3.94 g of Ca(NO3)24H2O and 0.71 g of P2O5 were dissolved separately in 10 ml of ethanol. Ca-precursor was added dropwise into the P-precursor to produce a mixture containing a Ca/P ratio of 1.67. The obtained suspension was stirred at 400 rpm for 5 h at 26 °C in a closed beaker. Afterwards, the samples of a Mg alloy were dipped vertically into the suspension and withdrawn at a constant speed of 0.1 mm s1 using an electric dip coater. The coated substrates were then maintained at room temperature for 24 h in order to complete aging. Then, they were gradually heated to 60 °C and maintained at this temperature for 24 h. Afterwards, the coated samples were calcined at 400 °C for 6 h [64]. To date, sol–gel preparations combined with dip coating have been established as cost-effective and simple systems to set up, as well as possessing the ability to coat irregular shapes, similar to spontaneous precipitation techniques. This approach shows the potential of using relatively low temperatures and short incubation times to achieve thick coatings on Mg and its alloys. Their primary advantage over the aforementioned spontaneous precipitation techniques is in the strength of the coating/substrate adhesion [182,183]. 4.4. Electrodeposition

Fig. 6. In vivo determination of the protective properties of calcium orthophosphate coatings on Mg and its alloys. (Top) Optical images of the bare Mg and HAcoated Mg specimens after 2, 4, 6, 8 and 12 weeks of implantation. (Bottom) Ultimate tensile strength of the bare Mg and HA-coated Mg specimens after 2, 4, 6, 8 and 12 weeks of implantation. Reprinted from Ref. [138] with permission.

(0 4 0) planes was observed for precipitated DCPD, which allowed the authors to vary crystallinity of the coatings [115]. To conclude this part, spontaneous precipitation techniques have proven to be popular methods for coating Mg and its alloys. However, there are reports 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 [175]. 4.3. Sol–gel preparation and dip coating Sol–gel preparation combined with the dip-coating technique has been widely investigated to coat Mg and its biodegradable alloys for both corrosion protection and increased adhesion. The

Electrodeposition is a broad range of deposition techniques involving electrical current. It comprises electrochemical deposition (or cathodic deposition), electrophoretic deposition and some other techniques. In general, electrodeposition is a low-cost and simple process that can be carried out at room temperature to form uniform coatings, and has also been applied for deposition of calcium orthophosphate coatings on Mg and its biodegradable alloys [31,69,75,99,118,129,134,142,148,152,167,184–186]. This technique is commonly performed from aqueous solutions similar to those used in wet-chemical deposition (see Section 4.2 above). 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. Afterwards, a transformation step was carried out in 1 M aqueous NaOH for 2 h to convert the biphasic mixture into uniform CDHA coatings [142]. Similarly, another research group conducted cathodic deposition of calcium orthophosphate coatings on AZ31 alloy at 85 °C for a period of 4 h and obtained a combination of CDHA and DCPD phases. Afterwards, the coated samples were immersed in 0.25 M NaOH solution for 4 h at 60 °C to transform the biphasic (CDHA + DCPD) coatings into the single-phase CDHA ones (Fig. 7) [59]. The authors of those studies


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Fig. 7. (Left) A cross-sectional SEM image and (right) an EDS spectrum (performed in the cross on the left image) of a CDHA-coated specimen of Mg-based AZ31 alloy. The EDS shows that the coating consists of Ca, P and O elements with a Ca/P atomic ratio of 1.45. Reprinted from Ref. [99] with permission.

demonstrated that the deposited CDHA coatings both improved corrosion protection (namely, the numerical value of corrosion potential (Ecorr) of AZ31 substrate increased from 1.6 to 1.42 V [99]) and improved the bioactivity of the samples. Furthermore, the CDHA-coated samples exposed to SBF showed 20% improvement in mechanical strength as compared to that of the uncoated samples [148]. The electrodeposition process of calcium orthophosphate coatings on Mg and its alloys can be modified to include 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 [100,101,139, 165,168,187]. Adjustment of both the pulse current parameters and electrolyte solutions were suggested as effective ways to

control the coating structure. By this means, CDHA was successfully deposited on Mg–Zn–Ca alloy directly without the need for a transformation step [100]. In addition, dense and uniform fluorine-doped HA coatings were electrodeposited the same way but with addition of H2O2 to the electrolyte to reduce H2 evolution by oxidizing this species [139]. Furthermore, pulse electrophoresis at 85 °C for 30 min revealed greater adhesion of calcium orthophosphate coatings to Mg–Zn–Ca alloys, whilst pulse treatments of the coated samples of AM50 alloys at 20 °C for 15 min displayed a higher corrosion resistance [187]. Other researchers utilized cathodic deposition to coat DCPD onto Mg–Zn alloys [15,124,188]. The DCPD coatings were deposited at room temperature by adding 10 ml l1 H2O2 into the electrolyte and adjusting the pH value to 4.4. The authors stated that the DCPD phase could effectively decrease the degradation rate of the alloy as well as increase its biocompatibility. Further research

Fig. 8. Representative micro-CT reconstruction images of the rabbit femora containing both uncoated (top) and CDHA-coated (bottom) Mg–Zn–Ca alloy implants at 8, 12, 18 and 24 weeks postoperatively. Reprinted from Ref. [101] with permission.



indicated that the electrophoretically deposited DCPD on Mg–Zn alloys (which were then converted to CDHA via alkali treatment) is more stable and effective in corrosion resistance compared to the originally deposited DCPD. Similar results were obtained by other researchers [134]. However, coatings of fluoridated HA were found to possess even better corrosion protective properties [15,124]. Further research analyzed the effect of different deposition times (20, 60, 120 or 240 min) on cathodic deposition. The authors concluded that the deposition time had a great influence on the morphology of the coatings; however, it did not influence the conversion of DCPD to CDHA or other calcium orthophosphates [189]. Normally, electrodeposition of calcium orthophosphate coatings on Mg and its biodegradable alloys is performed from aqueous solutions (see above); however, to improve certain properties of the coatings, studies are available on the addition of organic solvents, such as alcohol [75,118,168]. For example, to reduce 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 proportions, i.e. 10%, 30%, 50%, and 70% (v/v), and electrochemical deposition was carried out on AZ91 alloy using a constant-potential method. This research revealed a significant decrease in hydrogen bubble bursting during deposition from ethanol-containing solutions, which resulted in denser packing of the precipitated DCPD crystals and, thus, a higher degradation resistance of the DCPD coatings, as compared to those formed from ethanol-free solutions. The optimum results were obtained for 30% (v/v) ethanol-containing solutions, while further a increase in ethanol content in the solution produced thinner DCPD coatings offering poorer protection [118]. Representative micro-computed tomography (micro-CT) reconstruction images of in vivo biodegradation studies of electrochemically deposited calcium orthophosphate coatings are presented in Fig. 8 [101]. At 8 weeks of implantation, the surface morphology of the residual bare Mg alloy implant exhibited obvious corrosion pits (indicated by red arrows in Fig. 8a), while there was only slight corrosion with superficial pits for the CDHA-coated specimen (indicated by red arrows in Fig. 8e). With increasing implantation time, corrosion of both samples gradually became severe due to the local

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failure of the CDHA coatings (shown in Fig. 8c, d, g and h). 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]. To conclude this part, calcium orthophosphate coatings obtained by electrochemical methods have a uniform structure since they are formed gradually through a nucleation and growth process at relatively low temperatures. 4.5. Hydrothermal Hydrothermal treatment is a simple process and one of the most cost-effective techniques for coating deposition on metallic surfaces. It is rather similar to the aforementioned biomimetic deposition and wet-chemical precipitation; however, since the hydrothermal treatment is performed at elevated (>90 °C) temperatures over a relatively prolonged period of time (>1.5 h), the calcium orthophosphate deposits are usually crystalline. However, it is difficult 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. To the best of my knowledge, until recently only one research group, located in Japan, was actively pursuing this approach [57,102–109]. However, in 2012 and 2014, research groups from China [132] and Korea [138], respectively, also published papers on this technique. Using a hydrothermal treatment for 2 h at 363 K, the Japanese researchers succeeded in forming 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 authors, the formation of highly crystalline coatings was achieved primarily due to using Ca-EDTA solutions, which could supply 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

Fig. 9. Surface (a,b) and cross-sectional (c,d) SEM images of calcium orthophosphate coatings deposited on Mg synthesized under the hydrothermal conditions in the treatment solutions with a pH value of 5.9 (a,c) and 8.9 (b,d). As seen in (c) and (d), the outer porous layer and the inner continuous layer are located in the left and right sides of the images, respectively. Reprinted from Ref. [109] with permission.


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Fig. 10. A schematic illustration of the formation and growth mechanism of HA coatings on Mg under hydrothermal conditions. Reprinted from Ref. [104] with permission. Another illustration of this mechanism is available in Ref. [57].

microstructures of the coatings were found to vary with the pH of the treatment solutions. Namely, in weak acidic (pH 5.9) solutions, a dual-layer structure was formed: an outer coarse layer consisted of plate-like OCP crystals and an inner dense layer consisted primarily of HA crystals. In weak alkaline (pH 8.9) solutions, a duallayer structure was also formed: an outer coarse layer consisted of rod-like HA crystals and an inner dense layer consisted of well-packed HA crystals (Fig. 9). In strong alkaline (pH 11.9) solutions, needle-like HA crystals were formed. Both layers were found to grow with an increase in the treatment period. A thin Mg(OH)2 layer was also formed at the boundary between the calcium orthophosphate coatings and Mg substrates (Fig. 10). The HA and OCP coatings were found to improve the corrosion resistance of both pure Mg and Mg–Al–Zn alloys in both HBSS and a 3.5 wt.% NaCl solutions; however, the corrosion resistance of HA coatings was always higher than that of OCP ones [57,102–109]. Similar results were obtained by other researchers [132,138]. In addition, such coatings showed good adhesive properties with slight 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 [108]. The authors revealed that the level of protection afforded 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. 4.6. Aerosol deposition In addition, calcium orthophosphate coatings can be put down on Mg and its biodegradable alloys by an aerosol deposition technique. This technique has been used to deposit HA onto the surface of both pure Mg and Mg previously covered by either poly(e-caprolactone) [117] or MgF2 [116]. To perform aerosol deposition, HA powder was sprayed onto Mg samples in a deposition chamber using oxygen carrier gas at a flow rate of 5 104 m3 s1 under a pressure of 9.2 Torr. Scanning electron microscopy (SEM) observations 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 [117]; however, when HA was deposited onto Mg with the MgF2 interlayer, no composite-like structures were observed [116]. Corrosion tests

performed in SBF revealed that such coatings had good corrosion resistance (Fig. 5, bottom). 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 [117]. 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 [68]. 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 min1. The 5 lm 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 s1. 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 [68]. 4.7. Spin coating Only two publications on spin coating could be found, and both of these were devoted to deposition of calcium orthophosphatebased biocomposites on Mg alloys [73,76]. Initially, biocomposites of HA/collagen (HAC) [73] and HA/poly(lactic-co-glycolic acid) [76], 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 [73]. Both the deposition technique and the obtained corrosion results appeared to be similar for Mg alloys



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covered by HA/poly(lactic-co-glycolic acid) coatings [76]. Thus, the spin-coating technique appears to be a convenient tool for deposition of composite coatings. 4.8. 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 [74], while the other was devoted to deposition of HA coatings on AZ51 alloy [112]. 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 °C for 12 h and then at 67 °C for 1.5 h [74]. In the second paper, the deposition technique was almost the same but the spraying process was performed onto AZ51 substrates heated to 400 °C [112]. 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) [74]. Similar trends were obtained in another study [112]. In addition, cytocompatibility studies with MC3T3 cells revealed a continuous increase in cell growth on the coated samples [74]. 4.9. Ion-beam assisted deposition Just one publication was found on ion-beam assisted deposition of calcium orthophosphate coatings on Mg and its alloys, and that publication was devoted to deposition of calcium orthophosphates with a Ca/P ratio >1.67 on AZ31 alloy [113]. To achieve this, calcium orthophosphate evaporants were prepared by adding 37% CaO powder to HA powder. Prior to deposition, the surface of AZ31 substrates was cleaned by 120 V and 2 A Ar+ beam bombardment for 20 min. Afterwards, vapor fluxes of evaporants, generated by an electron beam evaporator, were deposited on the rotating AZ31 substrates. The substrate temperature during deposition was kept below 100 °C. To convert amorphous phases (ACP) into HA, after deposition some specimens were subjected to annealing at 250 °C for 2 h and then kept boiling in deionized water for 30 min. Compared with the uncoated AZ31 samples, the numerical values for both hardness and elastic modulus of the coated samples were found to increase 77% and 55%, respectively. In addition, in degradation tests performed in 3% aqueous solutions of NaCl, the mass loss of the coated samples appeared to be just one-fifth of that of uncoated samples (Fig. 11), indicating that the ion-beamdeposited calcium orthophosphate coatings decreased the degradation rates of AZ31 significantly [113]. However, there is another study, in which nanostructured ‘‘HA coatings were deposited on polished Mg using the patented transonic particle acceleration (Spire Biomedical) process’’ [133]. Unfortunately, no information on the deposition process was disclosed. A search on the Spire Biomedical website revealed the following information: ‘‘Deposited at low temperature and offering a wide range of calcium phosphate formulations, IonTite™ coatings are ideal for promoting bone/implant interfacial bonding. These adherent coatings can be applied to dental, joint replacement and fixation products made of biomaterials such as Co–Cr, Ti, stainless steel and most other metals.’’ Further investigations have revealed that the

Fig. 11. A comparison of the degradation behavior of the uncoated and calcium orthophosphate-coated samples of AZ31 alloy performed in 3% aqueous solution of NaCl at different immersion times. Reprinted from Ref. [113] with permission.

company uses an ion-beam-assisted deposition technology; therefore, with a reasonable level of probability, Ref. [133] might be considered as the second publication on this technique. The results revealed that in comparison to the non-coated Mg samples, the coated samples showed significantly decreased degradation rates, indicating that the HA-coating was protecting the Mg samples from rapid degradation [133]. 4.10. Direct laser melting Again, just one publication was found on this topic and that publication was devoted to deposition of HA on AZ31B alloy [146]. To perform the deposition, a precursor was first prepared by mixing a HA powder with an aqueous solution containing a binder and a reducer. The obtained slurry was then deposited onto the surface of AZ31B alloy substrates using an air-spraying gun, followed by air-drying for 24 h to remove moisture. The precursor-deposited substrates were subsequently subjected to laser processing using an ytterbium (IPG YLS-3000) fiber laser (1064 nm) of continuous wave with Gaussian power distribution and a beam focal spot of 1 mm. The results of further investigations revealed improvements of both corrosion resistance (+48%) and biodegradability (+180%) for HA-coated AZ31B samples for the selected laser process variables, such as power, scanning speed and beam indexing [146]. 4.11. Double-layered capsule hydrothermal hot pressing Just one publication was found on this topic and that publication was devoted to deposition of HA on AZ31 alloy [123]. First, an Mg alloy rod and a powder mixture of DCPD and Ca(OH)2 were placed into a polyethylene tube. The powder mixture was loaded into the tube such that the Mg alloy rod was concentrically positioned with respect to the tube axis. Both ends of the tube were fastened with paper staples. The entire tube was then further encapsulated using a poly-vinylidene-chloride film, and alumina powder was placed between the tube and film. The entire construction was put into a batch-type high-temperature and highpressure vessel for hydrothermal treatment. Then, the vessel was heated up to 150 °C for 3 h, while the pressure was kept at 40 MPa using a pressure regulator. After the treatment, the vessel was cooled down to room temperature and the HA-coated AZ31 samples were removed. Afterwards, pullout tests were conducted in order to measure the adhesion of the HA coating to the AZ31 substrates. The average value of the maximum shear stress was determined to be 6.1 ± 1.0 MPa. In addition, it was revealed that HA remained on the surface of the Mg alloy after the pullout fracture tests. Thus, by means of this technique, HA coatings could be bonded to Mg and its alloys with good adhesive properties [123].


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Fig. 12. Cross-sectional SEM images of micro-arc oxidized (MAO) Mg substrates: initial (left) and with deposited calcium orthophosphate coatings (right). The line scanning of element distribution across the coatings is depicted. It can be clearly seen that the MAO layer consists of magnesium silicates. Reprinted from Ref. [143] with permission. Additional pictures of the distributions of the elements across the coatings are available in other publications [57,169].

4.12. Micro-arc or plasma electrolytic oxidation Micro-arc oxidation (also called plasma electrolytic oxidation, or anodic spark deposition, or micro-arc discharge oxidation) is a combination of plasma-chemical and electrochemical processes. It combines an electrochemical oxidation with a high-voltage spark treatment performed 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 [48,66,70–72,110,126,143,147,164,190,191]. Therefore, the protective coatings consist mainly of MgO combined with magnesium silicates, borates, etc., depending on the chemical composition of the electrolytic bath. In the case of Mg alloys, such coatings also contain oxides, silicates, borates, etc., of the alloying elements. In the

Fig. 13. X-ray diffraction patterns of the calcium orthophosphate coatings formed on the surface of micro-arc oxidized (MAO) Mg substrates soaked in calcium orthophosphate-containing solutions. It can be seen that the initial surface of the MAO Mg covered by MgO (bottom pattern) was covered initially by CDHA (patterns at 20–45 min) and finally by a mixture of CDHA with DCPD (top pattern). Reprinted from Ref. [126] with permission.

presence of Ca- and P-containing salts, such coatings also contain various Ca- and P-containing compounds (but not calcium orthophosphates!) [127,135,169,192,193]. Ca- and P-containing coatings were prepared on AZ91D alloy in both NaOH- and Na2SiO3-containing electrolytes with addition of sodium hexametaphosphate and calcium hypophosphite. According to the results of energy-dispersive spectroscopy (EDS), the coatings prepared in the NaOH system were mainly composed of oxides of Mg, Al, P and Ca, while those prepared in the Na2SiO3 system also contained a substantial amount of Si oxide. The results of X-ray diffraction revealed MgO and Mg2SiO4 to be the predominant phases in the coatings prepared in the NaOH and Na2SiO3 systems, respectively [192]. Similar results were obtained in other studies [127,135,169]. However, among other phases, formation of CaNaPO4 was detected in one case because the electrolyte used in that study contained CaCO3, Ca3(PO4)2 and 20 g l1 of orthophosphate ions [193]. Since, except for Ref. [193], formation of calcium orthophosphates was not observed in the coatings, 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 (see Section 3 above). However, using other techniques, such as chemical precipitation [72,126,143,147] or electrodeposition [66,70,71,110,

Fig. 14. Tafel curves to measure the corrosion potential in SBF for bare AZ91D alloy, micro-arc oxidized AZ91D alloy (MAO-AZ91D) and MAO-AZ91D alloy covered by calcium orthophosphate coating. The more positive the value of the corrosion potential, the better the corrosion resistance. Reprinted from Ref. [71] with permission.



127,164], calcium orthophosphate coatings might be deposited over micro-arc oxidative coatings (Figs. 12 and 13). Mg–Zn–Ca alloys coated with such composite coatings were found to induce rapid precipitation of calcium orthophosphates from simulated solutions with simultaneous increase in the corrosion resistance of Mg and its alloys [127]. Similar results were obtained in other studies [110,143,164]. In addition, micro-arc oxidation of Mg surface can be performed simultaneously with electrodeposition of calcium orthophosphates, which results in the formation of composite MgO/calcium orthophosphate protective coatings on Mg and its alloys [70]. More to the point, calcium orthophosphate-based biocomposite coatings can be deposited over microarc oxidative coatings [66,71,72]. Therefore, combinations of micro-arc oxidation (to provide good corrosion resistance) with any suitable deposition technique for calcium orthophosphates (Fig. 14) appear to be effective and promising methods for constructing biodegradable Mg-based metallic bone grafts offering improved patient outcomes. 5. Conclusions Due to their unique properties, the utilization of Mg and its biodegradable alloys as bone-graft substitutes appears to be very promising but, as identified above, for successful clinical applications the corrosion of these metals must be slowed down. Development of non-toxic, biocompatible and corrosion-resistant protective coatings seems to be the most promising route to solve this problem. Since the use of calcium orthophosphates is well established in the field of biomaterials, they appear to be the excellent candidates to construct the required coatings on Mg and its biodegradable alloys. An additional advantage of using calcium orthophosphates as protective coatings is to provide the implants with surface biological properties for adsorption of proteins, adhesion of cells and bone apposition. Therefore, the knowledge available on calcium orthophosphate coatings on Mg and its biodegradable alloys has been summarized in this review. The following conclusions can be drawn from the above discussion. Although the combination of calcium orthophosphates and Mg metal is a new area of research that has still to be thoroughly explored, the number of the available publications on the topic clearly indicates that the use of calcium orthophosphate coatings on Mg and its biodegradable alloys is rapidly becoming common practice. However, in many cases, satisfactory results have not yet been achieved, typically due to an insufficient amount of research. That is to say, according to the available literature, calcium orthophosphate coatings, films and layers can be deposited and/or created on various substrates by means of, at least, 22 different techniques and each of those techniques has both advantages and shortcomings of its own [91]. In addition, new deposition techniques are continuously being elaborated. However, as seen from this review, only 12 deposition techniques have been used to form calcium orthophosphates on Mg and its biodegradable alloys, and for five of these just one or two publications are available, which is not enough. Even less information is currently available on in vivo investigations. Therefore, although the development of appropriate techniques for coating Mg and its biodegradable alloys with calcium orthophosphate holds great potential for the development of novel biomaterials, the information currently available on the subject appears to be insufficient to draw serious conclusions. Appendix A. Figures with essential colour discrimination Certain figures in this article, particularly Figs 2, 5–8, 10, 12 and 14, are difficult to interpret in black and white. The full colour

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images can be found in the on-line version, at http://dx.doi.org/ 10.1016/j.actbio.2014.02.026.

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