Surface design of biodegradable magnesium alloys — A review

Guosong Wu, Jamesh Mohammed Ibrahim, Paul K. Chu ⁎Department of Physics and Materials Science, City University of Hong Kong, Tat Chee Avenue, Kowloon, Hong Kong, China

a b s t r a c ta r t i c l e i n f o

Available online 12 October 2012

Keywords:Magnesium alloysBiomaterialsCoatingsIon implantationCorrosion

Biodegradability is a big advantage of magnesium-based materials in biomedical applications such as bonefixation, cardiovascular stents, and even stomach trauma repair. Different from other metals such as stainlesssteels and Ti alloys, the interface between the Mg-based implants and biological environment is dynamic. Inorder to improve the surface properties to allow better and more expeditious adaptation to the physiologicalsurroundings, it is imperative to design and construct a surface to satisfy multiple clinical requirements suchas mechanical strength, biocompatibility, and degradation rate. This paper reviews recent work pertaining tosurface modification of Mg-based biomaterials with emphasis on surface coatings and ion implantation. Thebiodegradation behavior and related mechanism in the physiological environment after surface modificationare also described. Surface modification is a promising means to elevate the performance of Mg-based bioma-terials and expected to be extensively applied to surface design of biomaterials.

© 2012 Elsevier B.V. All rights reserved.

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 22. Coatings and surface engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4

2.1. Design principle for surface coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 42.2. Substrate-involving coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 52.3. Non substrate-involving coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6

2.3.1. Ca–P based coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 62.3.2. Polymer-based coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7

2.4. Composite coatings . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 73. Ion implantation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 84. Degradation behavior of surface-modified Mg-based materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 95. Conclusion . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11Acknowledgments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 11References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12

1. Introduction

Magnesium and magnesium alloys were used in biomedical appli-cations more than a century ago but progress was stifled because cor-rosion of magnesium in vivo could not be solved adequately [1]. Onaccount of the clinical needs for biodegradable metallic implants,there has been resurgence in developing Mg-based biomedical im-plants. In fact, the Young's modulus (E=41–45 GPa) of Mg alloys issimilar to that of bones (E=3–20 GPa) and can reduce the stressshielding effect in orthopedic applications [2]. If Mg alloys with the

desirable surface properties can be successfully incorporated intobone-fixation implants (Fig. 1a), a second surgical process to removethe implant from the patient can be obviated thereby minimizingtrauma to the patients and decreasing medical costs [3]. Natural deg-radation in the physiological environment also bodes well forMg-based stents (Fig. 1b) because if the treatment fails or symptomsrelapse, new stents can be reinserted into the same sites [4,5]. Recent-ly, magnesium alloys have been made into a component of a degrad-able wound-closing rivet for gastrointestinal intervention (Fig. 1c)[6]. Furthermore, Mg-based materials coated with a conducting poly-mer offers a versatile biodegradable and biocompatible platform forcontrolled release of drugs on the therapeutic levels [7] and theserecent revolutionary applications have broadened the biomedicalapplications of Mg-based materials.

Surface & Coatings Technology 233 (2013) 2–12

⁎ Corresponding author. Tel.: +852 2788 7724; fax: +852 2788 7830.E-mail address: paul.chu@cityu.edu.hk (P.K. Chu).

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In order to comprehend the biodegradation and biocompatibility ofMg-based materials, some commercial magnesium alloys such as AZ31[8], AZ91 [9,10],WE43 [11], WE54 [12] and ZM21 [13] have been inves-tigated in vitro and in vivo. Since many commercial Mg alloys are origi-nally designed for other purposes such as automobiles, development ofnovel Mg alloys with the suitable biomedical properties is crucial to the

success. In this respect, some biologically important elements havebeen incorporated to form for instance, binary Mg-based alloys suchasMg–Zn [14,15], Mg–Sr [16], andMg–Ca [17]. Zberg et al. [18] fabricat-ed Mg–Zn–Ca bulk metallic glass and observed no hydrogen evolutionafter implantation in vivo. Recently, in order to obtain uniform biodeg-radation in simulated physiological solutions, a new Mg–Nd–Zn–Zr

Fig. 1. Potential applications of magnesium-based materials to: (a) osteosynthesis [3], (b) cardiovascular stents [5], and (c) wound-closing devices for stomach trauma [6].

Fig. 2. Dynamic interface between the Mg-based materials and bio-environment during surface degradation.

3G. Wu et al. / Surface & Coatings Technology 233 (2013) 2–12

alloy has been developed for bone nails and cardiovascular stents[19,20].

In an aqueous solution, dissolution of magnesium proceeds by thereaction: Mg+2H2O→Mg2++2OH−+H2↑, but the physiologicalenvironment complicates the corrosion process. Magnesium dissolutionresults in the emission of hydrogen andbasification, that is, increasedpH.If the OH− concentration adjacent to the surface increases to a certainextent, Ca (or Mg) containing phosphates will precipitate forming a sur-face layer on the surface [21]. It has been found that the degradation rateon most magnesium alloys is too large, particularly in the early stage[22]. Hence, the corrosion process will alter the interface between theMg-based biomaterials and bio-environment in vivo. This dynamic inter-face illustrated in Fig. 2 is different from those onbiomedical Ti alloys andstainless steels in the physiological environment. Generally, the surfacemorphology, microstructure, and composition, which play critical rolesin the efficacy of artificial implants, can alter protein absorption whichmediates adhesion of desirable and undesirable cells [23]. The corrosiondynamics is in fact quite complex and Mg-based implants should bedesigned in such away to control hydrogen evolution, localized basifica-tion, as well as surface degradation, especially in the early stage. Hydro-gen bubbles and surface alkalization can also influence tissue growth andso it is necessary to modify the surface of Mg alloys in order to mitigatedegradation in the initial stage to ensure proper tissue healing andgrowth.

Gray and Luan [24]wrote a review describing the potentialmethodsfor surface modification of magnesium alloys. Since then, techniquessuch as electrochemical deposition [25,26], electroless plating [27–29],plasma electrolytic oxidation [30–33], physical vapor deposition[34–38], ion implantation [39–43], and laser treatment [44] have beeninvestigated. However, in the case of biodegradable Mg-based mate-rials, the aim of the surface design is to alter but not change permanent-ly the surface structure and properties. Hence, the dynamic interfacemust be considered in order to endow the materials with the desirablemechanical properties, corrosion resistance, and biocompatibility dur-ing tissue healing. In this paper, recentwork related to surfacemodifica-tion of biodegradable magnesium-based materials is summarized anddiscussed. The surface design principle is discussed from the perspec-tive of coatings and ion implantation.

2. Coatings and surface engineering

2.1. Design principle for surface coatings

In the development of newMg-based biomaterials, the aim of sur-face design is to construct a temporary surface possessing the propermechanical properties, corrosion resistance, and biocompatibility. Asthe materials degrade naturally, this temporary surface disappearsgradually and should not produce deleterious and toxic effects. Inthis respect, surface coatings are quite useful. However, emphasizingonly the properties of surface coatings is often not enough in the sur-face design and the substrate must also be considered together withthe coating as one entity. For example, diamond-like carbon (DLC)with a high hardness does not adhere to Mg well unless an interlayersuch as Cr is inserted [45]. In addition, since Mg-based materials areused in different bio-environments, the mechanical strength and bio-compatibility of the coated Mg-based materials should be selectivelymodified in accordance with the actual requirements. For instance, ifthey are made into cardiovascular stents, the anticoagulation andmechanical integrity must be assured.

Researches have hitherto paid more attention to the degradationrate in the initial healing stage. Furthermore, sufficient mechanicalstrength and biocompatibility must be maintained on the temporarysurface during degradation and tissue healing. Compared to other tradi-tional engineeringmetals such as Ti alloys, Al alloys, and stainless steels,magnesium is chemically active and has a low standard potential [46].Generally, pores and cracks occur inevitably in the coatings duringpreparation and usage and the electrolyte in the bio-environment pen-etrates the coating via these cracks and pores. Fig. 3 depicts a schematicdiagram of the corrosion failure mechanism of the coated Mg-basedmaterials. If a galvanic cell is formed between the coating and substrate,the galvanic current is given by the following relationship [47]:

I ¼ EC−EMg

Rp Mgð Þ þ Rp Cð Þ þ Rs þ RMg–C;

where EC and EMg are the corrosion potentials of the cathode and anode,respectively, Rp(Mg) is the polarization resistance of the anode, Rp(C) is

Fig. 3. Schematic diagram illustrating the corrosion failure mechanism on coated Mg-based materials.

4 G. Wu et al. / Surface & Coatings Technology 233 (2013) 2–12

the polarization resistance of the cathode, Rs is the electrical resistanceof the electrolyte, and RMg–C is the electrical resistance between theanode and cathode. In a fixed bio-environment, Rp(Mg) and Rs do notchange easily and the current between the anode and cathode dependson EC−EMg, Rp(C), and RMg–C. If the resistance between the coating andsubstrate is very large, for instance, in the case of an insulating coating,each part will corrode independently. On the other hand, if the coatingis conductive, the above formula is valid. Obviously, higher EC−EMg andsmaller RMg–C and Rp(C) accelerate corrosion. A metallic interlayer is al-ways used to improve the adhesion between the top coating and thesubstrate, but Wu et al. [48] have found that addition of a metalliclayer can lower the corrosion resistance in cell culture media due tothe aforementioned galvanic effect. Therefore, it is better to avoidusing metal or conductive ceramic coatings and in past studies, some

Ca–P based and polymer-based coatings produced better results. All inall, it is necessary to select the coating with the proper characteristicsin the surface design but the substratemust also be considered simulta-neously. Here, we divide the coating processes into three classes, name-ly substrate-involving coatings, non-substrate-involving coatings, andcomposite coatings. The following section presents an overview ofthese processes.

2.2. Substrate-involving coatings

A substrate-involving coating process is defined as follows. A portionof themagnesium substrate participates in the preparation process andis eventually converted into the deposited coating. Conversion coatings

Fig. 4. (a) Surface and cross-sectional morphologies of a micro-arc oxidation (MAO) treated Mg–Ca alloy [49]. The inset shows the EDS result. (b) Surface and cross sectional mor-phologies of hydroxyapatite (HA) coated AZ31 magnesium alloy [62]. (c) Surface morphologies of porous polymer membranes on an AZ91 magnesium alloy [68]. (d) Surface andcross-sectional morphologies of dicalcium phosphate dihydrate (DCPD) and dicalcium phosphate dihydrate/polycaprolactone (DCPD/PCL) coatings on an Mg–Zn alloy [73].

5G. Wu et al. / Surface & Coatings Technology 233 (2013) 2–12

such as those produced by oxidation, anodization, and phosphate con-version belong to this class.

Microarc oxidation (MAO) or plasma electrolytic oxidation (PEO),one of the simple ways to fabricate porous and robust coatings onMg-basedmaterials, has been usedwidely in the development of bioma-terials. Gu et al. [49] utilized this technique to prepare a MAO coating onthe Mg–Ca alloy (Fig. 4a). They studied the effects of different appliedvoltages on the surface morphology, phase constituents, hydrogen evo-lution, and pH variation in a solution and the in vitro biocompatibilityof the MAO film. It was found that the thickness and pore size of theMAO coating increased with increasing applied voltages. Immersion inHank's solution for 50 days revealed that the 360 V MAO coating hadthe best long-term corrosion resistance. On account of reduced Mg ionrelease and pH value variations in the culture medium, adhesion, prolif-eration, and differentiation of MG63 cells were promoted by the MAOcoatings. The pulsing frequencywas also a critical parameter in the prep-aration process. The MAO coatings were deposited on an AZ31 Mg alloyin the pulsed DCmode using four pulsing frequencies of 300 Hz, 500 Hz,1000 Hz, and 3000 Hzwhile the pulse ratiowas constant. TheMAOcoat-ing produced at 3000 Hz exhibited the best corrosion [50].

Titania sol was introduced into the MAO process in an alkaline sili-cate electrolytic bath by Wang et al. [51]. By increasing the titania solconcentration from 0 to 10 vol.%, the coating thickness decreasedfrom 22 to 18 μm. The TiO2 modified coatings formed in electrolytescontaining 5 and 10 vol.% titania sol had progressively worse corrosionresistance in SBF than the unmodified coating. Ryu et al. [52] developedan approach to fabricate MAO coatings on the AZ31 magnesium alloyusing an AgNO3-containing electrolyte. The Ag-containing MAO coat-ings had higher corrosion resistance than the Ag-free MAO coatingsand also exhibited excellent antibacterial activity of over 99.9% againsttwo strains of bacteria, Staphylococcus aureus and Escherichia coli.

The blood compatibility of theMAO-treatedMgalloywas investigat-ed by Wang et al. [53]. The MAO-treated Mg–1.0 wt.% Zn–1.0 wt.% Caalloy exhibited favorable blood compatibility. The MAO group showeda decreased hemolytic ratio (2.25%) compared to the untreated Mgalloy group (24.58%) (pb0.001). The MAO group also showed signifi-cantly shorter prothrombin and thrombin times and significantly longeractivated partial thromboplastin time than the untreated Mg alloygroup. Arachidonic acid- and adenosine diphosphate-induced plateletaggregations were significantly reduced by the untreated Mg alloyextract, but they were less affected by the extract of the MAO-treatedMg alloy.

Compared to plasma electrolytic oxidation (PEO), another oxida-tion method is simpler and favored by researchers because it can re-duce the formation of micropores and microcracks. Gu et al. [54]soaked Mg–1.4 wt.% Ca alloy in three alkaline solutions (Na2HPO4,Na2CO3 and NaHCO3) for 24 h and subsequently annealed them at773 K for 12 h. Magnesium oxide layers with thicknesses of about13, 9, and 26 μm were formed on the surfaces of Mg–Ca alloy afterthe different alkaline and heat treatments. The in vitro corrosiontests performed in simulated body fluids indicated that the corrosionrates of Mg–Ca alloy were effectively retarded after the treatment andthe efficacy showed the following order: NaHCO3+annealingbNa2-HPO4+annealingbNa2CO3+annealing. Cytotoxicity evaluation fur-ther disclosed that none of the alkaline and heat treated Mg–Caalloy samples induced toxicity to L-929 cells during culturing for7 days. Lei et al. [55] employed an anodic electrodeposition processto prepare MgO coatings on magnesium alloy surfaces. The MgO coat-ings were produced by anodic electrodeposition in a 10 M KOH alka-line solution at a constant potential of 1.0 V (SCE) for 2 h, followed byannealing at 450 °C for 6 h in air. The MgO coating exhibited superiorstability and improved the corrosion resistance of the Mg alloy inHanks' solution.

Besides the aforementioned techniques, fluoridation is one of thesubstrate-involved coating processes that can effectively improve thesurface properties of magnesium alloys. Recently, Drynda et al. [56]

developed a binary fluoride-coated magnesium–calcium alloy that pos-sessed improved mechanical attributes, decreased degradation kinetics,as well as good biocompatibility with vascular cells.

2.3. Non substrate-involving coatings

Contrary to the substrate-involving coating processes described inthe previous section, most conventional coating techniques such asphysical vapor deposition (PVD), electrochemical deposition, andspraying do not require the participation of the substrate during film for-mation. These processes are thus categorized at non-substrate-involvingcoating ones. ZrN [57], DLC [58] and TiO2 [59] have been deposited onmagnesium alloys for biomedical applications. Most of these coatingsare non-biodegradable and inflammation may occur if they are brokenand remain in the human body for a prolonged period of time. Toenhance biodegradation and biocompatibility, Ca–P coatings and biode-gradable polymer-based coatings are potential candidates. Here, wedescribe two common coating techniques and review the recentresearch progress.

2.3.1. Ca–P based coatingsCalcium and phosphorus are the main elements in bone tissues and

Ca–P coatings, especially osteoconductive minerals such as hydroxyapa-tite (HA) and tricalcium phosphate (TCP), and have been widely used toconstruct new bones and promote osteointegration on biomedicalimplants. Song et al. [60] used electrodeposition to synthesize a bioactiveCa–P coating on an AZ91D magnesium alloy. This as-deposited coatingconsisted of dicalcium phosphate dehydrate (DCPD, CaHPO4·2H2O)and β-tricalcium phosphate (β-TCP, Ca3(PO4)2). After immersion in a1 M NaOH solution for 2 h, it was transformed into a uniform hydroxy-apatite (HA, Ca10(PO4)6(OH)2) coating. Correspondingly, the biodegra-dation rate of the coated AZ91D magnesium alloy was retarded in SBFs.Another type of bone-like nanowhisker HA coating was electro-deposited on an AZ31 substrate by Wen et al. [61]. Similar to Song'sresults, the as-deposited plate-like HA coating contained other calciumphosphates such as DCPD. The post NaOH alkaline treatment results inthe incorporation of Na+, Mg+, HPO4

2−, and CO32− and the resulting

structure had a composition similar to that of a natural bone. Thispost-treated HA coating retarded the degradation rate on the AZ31 sub-strate and induced deposition of Ca–P–Mg apatite in SBF more effective-ly. Hydrothermal treatment in a 250 mmol/L C10H12N2O8Na2Ca aqueoussolution with a pH of 8.9 was adopted by Hiromoto et al. [62] to preparehydroxyapatite (HA) coatings (Fig. 4b) on an AZ31magnesium alloy andpure Mg. The HA coatings consisted of an inner dense layer and outercoarse layer. The inner layer on the AZ31 was composed of dome-shape and densely packed precipitates whereas the outer layer com-prised rod-like crystals growing from each dome in the radial direction.With the HA coatings, magnesium ion release and the corrosion currentdensity were remarkably reduced. It was postulated that the protectionoffered by the HA coating was related to the inner layer and the efficacydid not depend significantly on the Mg substrate type.

Wang et al. [63] prepared a soluble Ca-deficient hydroxyapatite(Ca-def HA) coating on an Mg–Zn–Ca alloy substrate by pulseeletrodeposition. The Ca/P atomic ratio of the as-deposited coating wasabout 1.33 (within the range between 1.33 and 1.65). By regulating thepulse amplitude and width, the Ca-def HA coating showed better adhe-sion to the Mg–Zn–Ca alloy and the lap shear strength increased to41.8±2.7 MPa. The Ca-def HA coating improved the corrosion resis-tance significantly in SBF. The ultimate tensile strength and time of frac-ture measured from the coated Mg–Zn–Ca alloy were larger than thosefrom the uncoated one, thus offering benefits in supporting fracturedbone healing for a longer time. The way to control the ratio of Ca/P inthe coatingwas investigated by Yao et al. [64]. In their study, the ceramiccoatings containing Ca and P were prepared on the AZ91D Mg alloy byplasma electrolytic oxidation in NaOH and Na2SiO3. Adjustment of theCa2+ concentration in the electrolyte was an effective method to alter

6 G. Wu et al. / Surface & Coatings Technology 233 (2013) 2–12

theCa/P ratio in the coating. The reaction time andworking voltageweremore suitable for tailoring the Ca/P ratio in NaSi2O3 than NaOH.

Guan et al. [65] produced a hydroxyapatite (HA) coating on theMg–4.0Zn–1.0Ca–0.6Zr (wt.%) alloy using a process combining alka-line heat pretreatment, electrodeposition, and alkaline post heattreatment. The hemolysis rates of the HA coated and uncoated Mg–4.0Zn–1.0Ca–0.6Zr (wt.%) alloy samples were both b5%, which metthe requirements for implant materials. The HA-coated and uncoatedMg–4.0Zn–1.0Ca–0.6Zr (wt.%) alloy samples had the same cytotoxic-ity score as the negative control. However, the HA-coated samplesshowed a slightly larger relative growth rate (RGR%) of fibroblaststhan the uncoated samples. Both the HA-coated and the uncoatedMg–4.0Zn–1.0Ca–0.6Zr (wt.%) alloys provided evidence of acceptablecytocompatibility suitable for medical applications.

Generally, pure hydroxyapatite coatings suffer from relativelyhigh dissolution rates in the biological environment and do nothave long-term stability in vivo. Recently, fluoridated hydroxyapatite(Ca5(PO4)3(OH)1−xFx, FHA) has become a potential candidate asa substitute for HA in medical devices. Li et al. [66] preparedbone-like fluoridated hydroxyapatite (FHA) coatings on Mg–6 wt.%Zn electrochemically. The bioactive FHA coating improved the inter-facial bioactivity including biodegradation as well as cellular prolif-eration and differentiation. Because of the unfavorable effectsconcerning polarization of the concentration difference and H2 evo-lution, fluorine-doped hydroxyapatite coatings prepared by tradi-tional cathodic electrodeposition are loose and porous. In order toensure long-term stability of the Mg alloy implants, Meng et al.[67] introduced pulsed electrodeposition and also incorporatedH2O2 into the electrodeposition procedures. The dense and uniformnano-FHA coating produced effectively protected the Mg alloy sub-strate from corrosion. Since the nano-phase had a comparativelybig specific surface area, it also induced the precipitation of Mg2+,Ca2+ and PO4

3− more effectively in comparison with other tradition-al electrodeposited coatings.

2.3.2. Polymer-based coatingsBiodegradable polymers approved for human clinical applications

constitute a promising option to improve the initial corrosion resistanceand cell compatibility on Mg-based materials to meet the healingrequirement. Wong et al. [68] prepared porous polymeric membranescomposed of polycaprolactone (PCL) and dichloromethane (DCM) onan AZ91 magnesium alloy (Fig. 4c). The polymeric membranes reducedthe degradation rate while preserving the bulk mechanical propertiesupon degradation. The polymer-coated samples also showed bettercytocompatibility with eGFP and SaOS-2 osteoblasts than the uncoatedsamples and higher volumes of new bone were observed on the coatedsamples by micro-CT. Histological analysis indicated no inflammation,necrosis, and hydrogen gas accumulation during degradation. Owingto the excellent biocompatibility, a poly (lactide-co-glycolide) (PLGA)coating was produced on Mg–6Zn alloy samples by dipping [69]. Thecoating degraded by simple hydrolysis of the ester bonds into lacticand glycolic acids and the products were excreted by the normal meta-bolic pathways. The coated magnesium alloy showed excellent degra-dation resistance and enhanced cell attachment. In order to furtherunderstand the effects of coating thickness, adhesion strength betweenthe coatings and substrates, polymer molecular weight, and differentpolymers on the corrosion resistance and differentiated cell functionson the coated Mg-based materials, Xu et al. [70] adopted spin coatingto fabricate uniform, nonporous, amorphous poly (L-lactic acid)(PLLA) and semi-crystalline poly (ε-caprolactone) (PCL) films onextrudedMg substrates. The PLLAfilm showedbetter adhesion strengthto the Mg substrate than the PCL film. For both PLLA and PCL, the lowmolecularweight (LMW)filmswere thinner and exhibited better adhe-sion strength than the high molecular weight (HMW) ones. The corro-sion resistance of the Mg substrate was improved by the polymer filmsto a different degree according to the pH measurements of the cell

culture medium and quantification of released Mg2+ during the cellculture. All in all, all the polymeric films enhanced the cytocompatibilityduring incubation for 7 days.

2.4. Composite coatings

Composite coatings are becomingmore attractive as advanced coat-ings. With two or more constituents in a typical composite coating, theproperties can be fine-tuned to address specific requirements. In prac-tice, the materials consist of a layered or mixed structure. In the courseof developingMg-based biomaterials, various combinations such as ce-ramic/ceramic and ceramic/polymer have been proposed recently.

Aerosol deposition (AD) offers the advantage of room temperaturedeposition of ceramic–polymer composite materials. Hahn et al. [71]prepared a dense and well-adherent HA–chitosan composite coatingon an AZ31 Mg alloy. The composition of the coatings could be tailoredby adjusting the HA and chitosan concentrations in the powder mix-tures. All the coatings exhibited high adhesion strength ranging from24.6 to 27.7 MPa and better corrosion resistance than the bare Mgalloy. Moreover, the biocompatibility of the coated alloy was improvedappreciably. Generally, Ca–P ceramics have favorable biocompatibilityand osteoconductive properties but they also induce slow bone forma-tion in vivo. Compared to Ca–P ceramics, CaSiO3 ceramics can promoteproliferation and differentiation of osteoblast-like cells and acceleratethe formation of hydroxyapatite (HA) in SBF but unfortunately, CaSiO3

degrades rapidly in the physiological environment. To improve theanti-corrosion properties and cell compatibility, Du et al. [72] produceda microporous calcium silicate and calcium phosphate (CaSiO3–

CaHPO4·2H2O) composite coating on an Mg–Zn–Mn–Ca alloy using achemical reaction. The layer was mainly composed of CaHPO4·2H2Owith a small amount of CaSiO3. In vitro cell experiments indicated thatthe surface cytocompatibility of the coated Mg alloy was significantlyimproved, for instance, more cell adhesion, growth, and proliferation.

A composite coating composed of dicalcium phosphate dihydrate(DCPD) and polycaprolactone (PCL) was fabricated on an Mg–Zn alloy(Fig. 4d) [73]. The DCPD coating was synthesized in a 0.042 mol/LCa(NO3)2·4H2O and 0.025 mol/L NH4H2PO4 solution by electrodeposi-tion. The DCPD-coated samplewas immersed in a 2 wt.% PCL chloroformsolution and subsequently dried in air. Compared to the DCPD coatedalloy, the DCPD/PCL coated alloy had higher corrosion resistance asmanifested by the elevated corrosion potential, reduced corrosion cur-rent, and less hydrogen released.

Microarc oxidation (MAO) is a facile and effective means to fabricateprotective coatings on magnesium-based materials and many MAO-based composite coatings have been produced on biodegradablemagne-sium alloys. Gao et al. [74] produced a porous coating on an Mg–Zn–Caalloy byMAO and bymeans of electrochemical deposition (ED), fabricat-ed rod-like nano-hydroxyapatite (RNHA) on the MAO coating. The HArods were deposited inside the pores. The bonding strength betweenthe HA film and MAO coating of 12.3 MPa was almost twice as high asthat of the sample produced by direct electrochemical deposition(6.3 MPa). Owing to the enhancement in the bonding strength and de-position of RNHA inside the pores, the corrosion resistance of thecomposite-coated alloy was much higher than that of the MAO-coatedor uncoated alloys. In addition, according to the immersion test, theRNHA induced more rapid precipitation of calcium orthophosphatethan the conventional HA coating in SBF. Shi et al. [75] also adopted elec-trodeposition to make a composite coating consisting of an oxide layerand a top layer comprising dicalcium phosphate dihydrate (DCPD,CaHPO4·2H2O) on an AZ80magnesiumalloy. TheMAO/DCPD compositecoating reduced the corrosion rate on the AZ80 and at the same time en-hanced the deposition of apatite on the coating. Zhang et al. [76] pre-pared a porous Mg phosphate coating on an Mg–Zn–Ca alloy by MAOat a voltage between 120 and 140 V. A TiO2 layer was produced on theporous MAO layer by sol–gel dip coating followed by annealing. Similarto the results obtained by Gao et al. [74], the pores in the Mg phosphate

7G. Wu et al. / Surface & Coatings Technology 233 (2013) 2–12

layer provided accommodation sites for the subsequent TiO2 sol–gelcoating which sealed the pores and significantly enhanced the corrosionresistance. In comparison, a single MAO coating only led to limitedimprovement in the anticorrosion ability.

3. Ion implantation

Ion implantation involves a process in which ions are acceleratedand impinge into the surface. The technique provides the possibility ofintroducing different species into a substrate independent of thermody-namic limitations such as solubility. Ion implantation introduces a suit-able amount of ions into the near surface of the materials to alter thesurface properties such as biocompatibility and cytotoxicity. Unlike sur-face coatings, an ion implanted layer does not have an abrupt interfaceand layer delamination is not a serious issue. In particular, plasmaimmersion ion implantation (PIII) which can process samples with acomplex shape without complex sample manipulation or ion beamrastering has attracted much interest [77]. In PIII, the specimens aresurrounded by a plasma and pulse-biased to a high negative potentialrelative to the chamber wall. Ions in the overlying plasma are accelerat-ed across the sheath formed around the specimens and implanted intothe surface conformally. Together with high efficiency and ease of oper-ation, PIII excels as a surface engineering technique for biomedical im-plants with a complicated geometry [77].

Metal ion implantation can introduce alloying elements into themagnesium substrate. In biomedical applications, the biological proper-ties and toxicity of the alloying elements must be considered. Zn, one ofthe vital elements in the human body, has been implanted intomagnesium-based materials (Fig. 5). After plasma implanting 2.5×1017ions·cm−2 of Zn using a cathodic arc source into pure magnesium

at 35 kV, the degradation rate in SBF increased significantly. X-rayphotoelectron spectroscopy (XPS) revealed galvanic effects betweenthe metallic Zn-rich surface and magnesium matrix beneath [78]. Crplays a critical role in the corrosion resistance of stainless steels.When the Cr concentration in stainless steels exceeds 12%, a surfacelayer consisting of chromium-rich oxide forms on the surface and actsas a barrier against oxidation. Chromium plasma ion implantation intopure magnesium was found to induce rapid degradation in SBF similarto Zn ion implantation [79]. In comparison, Al has an electrical potentialclose to that of magnesium in aqueous solutions and was implanted totailor the surface corrosion resistance of pureMg, AZ31, and AZ91mag-nesium alloys. The corrosion resistance in SBF improved appreciablyand this enhancementwas attributed to the formation of a gradient sur-face structure involving a gradual transition from an Al-rich oxide layerto an Al-richmetal layer (Fig. 6). Compared to the high Al-contentmag-nesium alloy (AZ91), a larger reduction in the degradation rate was ac-complished from pure magnesium and AZ31 [80]. However, the use ofAl in biomedical applications is suspected because Al may be involvedwith Alzheimer's disease and may also cause muscle fiber damage[81,82].

Both Ti and Zr are biologically friendly to the human body and havebeen implanted into Mg-based materials. Wu et al. conducted Ti ion im-plantation to modify the surface of AZ31. In their study, the corrosionrate was accelerated in the NaCl solution [83]. Liu et al. used plasma im-mersion ion implantation and deposition (PIII&D) to introduce Ti and Zrinto an AZ91 magnesium alloy. After this treatment, the corrosion resis-tance was improved in SBF and the implanted layer was composed of atri-layered microstructure with the outer layer composed of mainlymetal oxide with a small amount of MgO andMg(OH)2, an intermediatelayer containing metal oxide and metallic implanted particles, and a

Fig. 5. (a) XPS depth profiles of Zn implanted magnesium [78]. (b) High-resolution XPS spectra of Zn implanted magnesium [78]. (c) Polarization curves of pure magnesium and Znimplanted magnesium in SBF [78]. (d) Surface and cross-section pictures of the samples after immersion in SBF for 18 h [78].

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bottom layer rich in metallic elements [84]. Generally, it is difficult toavoid oxidation when the samples are exposed to air or an oxygen-containing environment including that in a non-ultra-high-vacuum(non-UHV) processing chamber. Therefore, some of the improvementmay stem from the O-rich outer layer.

Wei et al. [85] performed oxygen PIII at different voltages to modifythe surface of an AZ31 magnesium alloy. The modified layer was com-posed of mainly MgO and some MgAl2O4. At higher sample voltages,the oxygen implant fluences and layer thickness increased. At 60 kV,the effects were more pronounced due to the heating effects from ionbombardment. Their study revealed that oxygen PIII using the optimalconditions improved the corrosion resistance of the AZ31 magnesiumalloy. However, a voltage that was too high created corrosion pits onthe surface and excessive oxygen diffusion led to a loose surface oxidewhich showed decreased corrosion resistance. Wan et al. [86] used oxy-gen PIII to treat pure magnesium using oxygen fluences in the range of2.5×1015ions·cm−2 to 2.0×1016ions·cm−2. The corrosion propertieswere investigated in the neutral phosphate buffer solution (PBS,pH 7.4) and chloride ion enriched PBS (145 mM Cl−, pH 6.4). Signifi-cantly enhanced corrosion resistance against PBS was observed fromthe Mg samples implanted with oxygen at fluences above 1×1016ions·cm−2. However, all the treated samples could not withstandthemore aggressive PBS (Cl−) attack and no improvement in the corro-sion resistance in this solution was observed even from the sampleimplantedwith 2.0×1016ions·cm−2 of oxygen. The enhanced corrosionresistance against neutral PBS was ascribed to more Mg\O bondingstates formed in the near surface by a PIII process whereas the Mg\Obonds could not withstand a Cl− enriched andmore acidic environment.Wu et al. also applied oxygen PIII to modify the Mg–Nd–Zn–Zr alloy but

different fromAZ31, no significant improvementwas observed from thisMg alloy in SBF [87].

The above results demonstrate that the composition of the alloysplays a critical role in the oxidation process. Magnesium corrosion is af-fected by the thin surface oxide film and dissolution typically occurs inthe oxide-free areas [88]. Generally, the native surface oxide filmformed upon exposure to air consists of mainly MgO, but according tothermodynamics, MgO is not stable in an aqueous solution and will beconverted into magnesium hydroxide. Cl− can substitute OH− formingchloridewhich expedites dissolution of the surface structure. In order toconstruct a temporary surface on biodegradable Mg-based materials toretard the initial corrosion process, some chemically stable phases suchas Cr2O3 and Al2O3 can be formed in the surface structure. Cr and Al ionimplantation was used to modify the surface of an Mg–Nd–Zn–Zr alloybefore oxygen ion implantation [87]. After the dual treatment, theAl-rich or Cr-rich oxide film is formed on the magnesium alloy toimprove the surface corrosion resistance in SBF. Xu et al. [79] performedoxygen PIII on pure magnesium after Cr ion implantation and surfacedegradation in SBF was retarded due to the formation of Cr-containingoxide on the surface. Similarly, Zhao et al. [89,90] found that the surfacecorrosion resistance on aWE43 alloy in SBF was significantly improvedafter oxygen PIII in conjunction with Al and Ti ion implantation (Fig. 7).

4. Degradation behavior of surface-modified Mg-based materials

After surface modification, the composition, phase structure, andmorphology of the surface are altered and so it is important to inves-tigate the degradation behavior on the biodegradable Mg-basedmaterials with the modified surface. Better understanding of the

Fig. 6. (a) XPS depth profile of pure magnesium after Al ion implantation [80]. (b) High resolution XPS spectra of Al-implanted pure magnesium at different sputtering time. Thenumber in the figure denotes the sputtering time (min) [80]. (c) Polarization curves of pure magnesium without and with ion implantation [80]. (d) SEMmicrographs of pure mag-nesium before and after the polarization test [80].

9G. Wu et al. / Surface & Coatings Technology 233 (2013) 2–12

Fig. 7. (a) XPS depth profiles of Ti implanted WE43 and Ti–O implanted WE43 [90]. (b) High-resolution XPS spectra acquired from the Ti-implanted WE43 and Ti–O implantedWE43 at different sputtering time. The numbers in the figures denote the sputtering time (min) [90]. (c) Polarization curves of un-implanted WE43, Ti-implanted WE43 and Ti–O implanted WE43 in SBF [90].

Fig. 8. (a) Schematic diagram illustrating the degradation performance of the WE43 samples after different heat treatments, (b) initial surface conditions on the annealed andpolished samples featuring a large reactive surface, (c) initial surface conditions on the oxidized samples with an oxide layer, and (d) penetrated oxide layer upon degradation.Xi represents the reactive area fraction [91].

10 G. Wu et al. / Surface & Coatings Technology 233 (2013) 2–12

effects of the temporary surface on biodegradation is imperative tothe optimal design and fabrication of Mg-based biomaterials. Recent-ly, some studies on the in vitro and in vivo biodegradation of surface-modified Mg-based materials have been performed.

Hanzi et al. [91] studied the biodegradation behavior of anMg–Y–REalloy (WE43)with various surface conditions in SBF. They employed thefollowing heat treatment protocols tomodify the surface: (1) annealingat 525 °C (plus artificial aging at 250 °C in one case) followed bypolishing and (2) polishing followed by annealing at 500 °C in air. Asshown in Fig. 8a, the degradation performance of the samples can bedivided into two groups. Hydrogen evolution from the annealed andpolished samples could be described by a parabolic law whereas theoxidized samples revealed a sigmoidal degradation behavior. On theannealed and polished samples, the high initial degradation rate wasdue to the large reactivemetallic surface initially exposed to the immer-sionmedium (Fig. 8b). Over time, degradation sloweddue to the forma-tion of corrosion products on the surface and the ensuing barrier action.On the oxidized samples, a very small degradation rate was observeddue to the protective effect of the oxide layer that covered the entiresurface of the samples (Fig. 8c). After the oxide layer was penetrated,the reactive surface area increased (Fig. 8d) and degradation accelerat-ed. The barrier action rendered by the deposited corrosion products re-duced degradation in these sites with time. At the same time, a “fresh”reactive surface in other locations was exposed to the liquid to triggeraccelerated degradation. This situation persisted until most of theoxide layer was removed and the corrosion products covered the de-graded surface. Afterwards, since degradation slowed, the degradationbehavior became similar to that of the annealed and polished samples.

Compared to the in vitro studies, in vivo evaluation can better ad-dress clinical needs and provide more appropriate information. Wanget al. [92] fabricated calcium-deficient hydroxyapatite (Ca-def HA) onanMg–Zn–Ca alloy and investigated the biodegradation in vivo. Magne-sium alloy rods were implanted into the rabbit femora and evaluatedover a course of 24 weeks after surgery. As shown in Fig. 9, in vivo deg-radation of the Ca-def HA coating and magnesium substrate occurred

almost simultaneously, and the in vivo lifetime of the coating wasabout 8 weeks, after which the degradation rate of the coated implantincreased obviously. Degradation of the Ca-def HA coatingwas attribut-ed to its reaction with body fluids and substitution of Mg2+ ions in theCa-def HA. Histological and pathological examinations showed that theCa-def HA coating had good osteoconductivity and favored new boneformation on the surface of the magnesium alloy implant. All in all,the Ca-def HA coating not only retarded the degradation of the magne-sium implant but also improved bone response in vivo.

5. Conclusion

An overview on common surface modification methods for Mg-based biomaterials is presented. Coatings and ion implantation as thetwomain techniques to fabricate the temporary surface at the dynamicinterface are described and recent work is summarized. Coatings can becategorized into three groups: substrate-involving coatings, non-substrate-involving coatings, and composite coatings. By means ofmicroarc oxidation, Ca–P based and polymer-based coatings servingas temporary surfaces and having good biological properties can be syn-thesized on Mg-based materials. Ion implantation is also a promisingtechnique and is reviewed from the perspective of surface corrosion ofmagnesium-based materials. Finally, the degradation behavior of thesurface-modified Mg-based materials and associated mechanism arediscussed. As surface modification techniques become more effectiveand economical, temporary surfaces that can better meet clinicalneeds are expected to emerge and their application to biodegradablematerials such asmagnesium-based ones will be acceptedmorewidelyby the biomedical community and industry.

Acknowledgments

This work was financially supported by the Hong Kong ResearchGrants Council (RGC) General Research Fund (GRF) nos. CityU 112510and CityU 112212 and the City University of Hong Kong Applied

Fig. 9. Representative micro-CT 2D reconstructed images of the femora containing magnesium implants [92].

11G. Wu et al. / Surface & Coatings Technology 233 (2013) 2–12

Research Grant (ARG) no. 9667066. The authors would like to thankDr. Ying Zhao, Prof. Shuilin Wu, and Prof. Guping Tang for helpfuldiscussions.

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