the current trends of mg alloys in biomedical applications ... · biomedical applications,...

Review Article

The current trends of Mg alloys in biomedical applications—A review

Usman Riaz,1 Ishraq Shabib,1,2 Waseem Haider1,2

1School of Engineering and Technology, Central Michigan University, Mount Pleasant, Michigan, 488592Science of Advanced Materials, Central Michigan University, Mount Pleasant, Michigan, 48859

Received 25 April 2018; revised 10 November 2018; accepted 15 November 2018

Published online 00 Month 2018 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.b.34290

Abstract: Magnesium (Mg) has emerged as an ideal alternative

to the permanent implant materials owing to its enhanced prop-

erties such as biodegradation, better mechanical strengths than

polymeric biodegradable materials and biocompatibility. It has

been under investigation as an implant material both in cardio-

vascular and orthopedic applications. The use of Mg as an

implant material reduces the risk of long-term incompatible

interaction of implant with tissues and eliminates the second

surgical procedure to remove the implant, thus minimizes the

complications. The hurdle in the extensive use of Mg implants

is its fast degradation rate, which consequently reduces

the mechanical strength to support the implant site. Alloy

development, surface treatment, and design modification of

implants are the routes that can lead to the improved corrosion

resistance of Mg implants and extensive research is going on in

all three directions. In this review, the recent trends in the alloy-

ing and surface treatment of Mg have been discussed in detail.

Additionally, the recent progress in the use of computational

models to analyze Mg bioimplants has been given special con-

sideration. © 2018 Wiley Periodicals, Inc. J Biomed Mater Res B Part B:

00B: 000–000, 2018.

Key Words: biodegradation, alloy development, surface treat-

ment, computational models

How to cite this article: Riaz U, Shabib I, Haider W. 2018. The current trends of Mg alloys in biomedical applications—A review.

J Biomed Mater Res B Part B. 2018:9999B:9999B:1–27.

INTRODUCTION

Magnesium (Mg) has emerged as a potential biomaterial owingto its biocompatibility, goodmechanical strength, and biodegra-dation. Mg is nontoxic1 in nature with the daily recommendedintake of 240–420 mg/day for adults.2 This value is almost50 times higher than the recommended intake of Iron (Fe) andZinc (Zn), which are other potential implant materials.1 Addi-tionally, Mg and its alloys have shown excellent biocompatibil-ity in physiological conditions.1,3–7 Along with biocompatibility,Mg has suitable mechanical properties for an implant materialsuch as being light weight and having a good strength to weightratio.8,9 Moreover, the elastic modulus of Mg is about 45 GPa,which is closer to the elastic modulus of bone (3–20 GPa),reducing the possibility of stress shielding.1 On the other hand,Fe and Zn have elastic modulus values of 211.4 and 90 GPa,respectively, much higher than that of bone.1 Along with suit-able mechanical properties and biocompatibility, biodegrada-tion is the primary reason for the enhanced interest in Mg as animplant material.7–13 Prolonged interactions of implants in thebiological surroundings can lead to many complexities10,13–15

and are not desirable. Mg alloy implants avoid these long-termincompatible interactions with the body tissues15 eliminating

the possibilities of any complexity. All the above-mentionedproperties make Mg a potential material to replace the conven-tional permanent implant materials.

The research on the potential of Mg as an implant mate-rial is not new as it has been under investigation in biomedi-cal applications since the late 1800s.8 It is reported that Mgin medical applications was first employed by EdwardC. Huse, who used Mg wires as ligatures to stop bleeding in1878.16 Later on, E.W. Andrew considered absorbable metalclips as an alternative to ligatures as they ensured the safetyof hemostasis.16 In 1900, Erwin Payr proposed the idea ofusing Mg as fixator pins, wires, plates, and nails.16 Addition-ally, he did significant work on the biodegradation of Mg andproposed the factors responsible for its in vivo corrosion.16

In 1906, Albin Lambotte treated a boy with a fracture in thelower leg by using a Mg plate with steel screws.16 The degra-dation of Mg inside the body encouraged him to investigateMg as a biodegradable implant. Later on, he successfullytreated four children having supracondylar humerus frac-tures.16 On the basis of successful results, Lambotte recom-mended the use of Mg implants in the treatment of severalfractures and surgeries.16 Jean Verbrugge, an assistant to

Correspondence to: W. Haider; e-mail: [email protected]

© 2018 Wiley Periodicals, Inc. 1

mailto:[email protected]

http://crossmark.crossref.org/dialog/?doi=10.1002%2Fjbm.b.34290&domain=pdf&date_stamp=2018-12-07

Lambotte, concluded that Mg implants are not toxic or irri-tant. However, he was concerned about the fast degradationrate of Mg implants and stated that a slow corrosion rate isrequired for the healing of the bones.16 McBride (1938) andMaier (1940) further investigated different aspects of Mg asan implant, including corrosion and compatibility issues.16

All the above-mentioned researchers preferred the use ofPure Mg implants. In 1948, Troitskii and Tsitrin first timeinvestigated the use of an Mg alloy (magnesium–cadmium)in the biomedical field and reported its successful use in thetreatment of 34 cases of pseudarthrosis.16 This research onMg in biomedical applications, though continued on smallscale throughout the 20th century, gained ground recently.The modern era of Mg research only started in the early2000s.1

Although the research on the potential of Mg as animplant material started in 1878, some of the core issueshindering the widespread use of Mg are still present. Therapid degradation of Mg in the physiological conditions andsubsequent loss of mechanical strength is the primary rea-son for implant failure.15,17,18 Another critical issue with therapid degradation of Mg implants is the excessive evolutionof hydrogen (H2) gas.

19,20 The accumulated gas bubbles canblock the regular flow of blood and can damage the tissuesat implant sites.20 Apart from rapid uniform degradation dueto the active nature of Mg, localized corrosion is anotherconcerning point in the case of Mg alloys. The formation ofthe galvanic cell due to the potential difference betweenpure Mg and alloying elements is a primary reason for thelocalized corrosion. The localized corrosion can lead to thefailure of the implant by the degradation of critical points.Therefore, the factor of localized corrosion is an importantconsideration in alloy designs. Hence, this fundamental issueof degradation needs to be addressed before the commer-cialization of Mg implants. The perfect compromise betweenthe mechanical strengths, biodegradation rate, and biocom-patibility is yet to be found. The aim of this article is to focuson the current trends in the research of Mg and its alloys forbiomedical applications, including the improvement of bio-degradation by surface treatment and alloy development,applications of Mg in the biomedical field, and computationalmodels used to simulate Mg and its implants.

Mechanical functionality of the Mg implantsMg and its alloys stand out of other biomaterials because oftheir mechanical properties that are similar to bone. Magne-sium has an elastic modulus of 45 GPa that is relativelycloser to the elastic modulus of bone (20 GPa).1 On the otherhand, the elastic moduli of other prospective biomaterialssuch as Fe, Zn, and Ti are 211.4, 90, and 120 GPa, respec-tively.1 The small difference between the elastic moduli ofMg and bone significantly reduce the possibility of stressshielding of the bones inside the body.1,9,21 Higher the gapbetween the elastic moduli of implant material and bone,higher will be the chances of stress shielding. Stress shield-ing is the reduction in the size and density of bone when itis surrounded by a stiffer material. Moreover, the mechanicalproperties of Mg and its alloys are superior to biodegradable

polymers. The degradation control of polymers is promising,but most of the polymers lack the strength needed to sup-port the implant site.22

Although Mg possesses the ideal mechanical propertiesfor biomedical applications, it has a limitation in terms ofmechanical integrity when it comes to the actual applicationsas an implant. The rapid degradation of Mg and its alloy inphysiological conditions leads to the loss of mechanicalintegrity.23,24 In orthopedics application, the implant shouldbe able to support the implant site until healing but in thecase of Mg, the rapid degradation leads to loss of mechanicalstrength and ultimately the failure of the implant. In case ofstent application, the implant does not need the mechanicalstrength required in orthopedic. However, it should be ableto maintain the implant shape to avoid any premature fail-ure. The stress concentration zones created by the bloodflow in stents can lead to stress corrosion cracking.25 Thestress corrosion can be more detrimental to the life of thestent as compared to uniform corrosion.15 The rapid loss ofmechanical integrity is a concern as it ultimately leads to thefailure of the implant without accomplishing its task.23

Narrowing down the problem reveals that degradation isthe prime reason behind the rapid loss of mechanical integ-rity. The rapid degradation has two reasons: (1) Mg is anactive material in physiological conditions and (2) stress cor-rosion cracking. This article will discuss in detail the solu-tions to counter the first problem. The second issue of stresscorrosion can be effectively solved by modifying the designof the implant. It has been proved in the literature that thedesign of the implant can significantly affect the stresses thatare induced by the concentrated zones.15 The Mg implantsare not mechanically suitable for the practical use withoutaddressing the above-mentioned limitations.

CURRENT BIOMEDICAL APPLICATIONS OF MG

The recent research in the field of Mg and its alloys revolvesaround two main applications: vascular stents26–32 andorthopedics.33–40 The first successful implantation of a bio-degradable Mg stent became possible in 2005 when it wasimplanted in treating a left pulmonary artery of a 6 week oldbaby.29 Erbel et al.30 studied the stenting procedure on63 patients that had a plaque in their arteries and the resultsproved the safe nature of Mg stents. Moreover, the angio-graphic results were similar to other metal stents.30 Simi-larly, Haude et al.31 proved the feasibility and safetyprospects of the drug-eluting absorbable metal scaffold bydeploying them in 46 patients with lesions. The studyshowed good clinical results and angiographic performanceup to 12 months, suggesting metal scaffolds as an alternativeto polymeric scaffolds with an advantage of mechanicalstrength.31,32 Heublein et al.26 investigated the possibility ofthe Mg alloy AE21 (2 wt % aluminum [Al] and 1 wt % rareearth elements [REEs]) as stent material and found noadverse effects in the body during implantation, but theissue of fast degradation was a concern. The study provedMg a potential stent material after improvements in degrada-tion time.26 Similarly, the in vivo study of the Mg alloy WE43

2 RIAZ, SHABIB AND HAIDER CURRENT TRENDS OF MG ALLOYS IN BIOMEDICAL APPLICATIONS

(3.7–4.3 wt % yttrium [Y], 2.4–4.4 wt % REEs, 0.4% zirco-nium [Zr]) stent design showed encouraging results in termsof mechanical strengths.27 Zhu et al.41 compared the siliconcoated biodegradable Mg stents with plastic stents byimplanting these stents in rabbits. Silicon covered Mg stentsshowed better degradation behavior and biocompatibility ascompared to the plastic stents.

The stress and strain analysis of stents during thedeployment and in the artery is an important considerationfor the performance of stents. Galvin et al.42 studied theeffect of plastic strains on the degradation of Mg alloy WE43stent. Li et al.43 optimized the design of Mg alloy stent usingfinite element analysis (FEA). Gastaldi et al.22 discussed thedegradation of Mg alloy stents using continuum damagemechanics (CDM). Bartosch et al.44 concluded that the use ofMg stents is safe but the research on Mg stents should beexpanded to large stent designs for large vessels and appli-cations other than just coronary stents.44 There are a num-ber of studies on the mechanical and degradation analysis ofMg-based stents15,45–51 which will be discussed in detail inthe “Computational Models” section.

Apart from cardiovascular applications, Mg has beeninvestigated in the orthopedics applications in the form ofpins, screws, rods, and plates.29–36 Biodegradation of Mgalong with its good mechanical strength makes Mg one ofthe ideal materials for the orthopedic applications. The studyof Mg screws in treating the anterior cruciate ligament (ACL)rupture proved Mg a suitable alternative to other metals andpolylactic polymer, due to its good mechanical strength, bio-degradation, and bone generation capability.37 The perma-nent implant materials can lead to graft damage in treatingACL, which could be avoided using biodegradable Mgscrews.37 The in vivo study of Mg pins in the femur of rab-bits showed that these pins retained their shape until the24th week of implantation with a uniform degradation. Abone–implant interface resulted in the fourth week. Addi-tionally, the Mg pins degradation showed an increase in Mgcontent in body tissues and urine but had no significanteffect on the feces. The study proved that degraded Mg canbe excreted through regular body channels without anyadverse effects.38 The implantation of the uncoated and mag-nesium phosphate (Mg3(PO4)2) coated Mg plate/screws sys-tems in the ribs of pigs showed less formation of gas andslow degradation in case of coated systems. The studyproved no harmful effect of Mg implants on bone formationand healing. The Mg plate/screw system exhibited goodmechanical stability with only a few screws broke on 12thand 24th weeks.39 Figure 1 shows the implanted plate/screws system in the ribs of the pig along with its schematicpresentation. Figure 1 also presents the histological sectionsof bone healing in osteotomized ribs after 24 weeks implan-tation of titanium, Mg3(PO4)2 coated Mg, and uncoated Mg.39

The effect of implant’s degradation on the hardness of bonesin the femur of rats was studied by comparing different Mgalloys with Ti alloy Ti–6Al–7Nb, polymeric PLGA, and a sam-ple with a drilled hole in the bone. An unimplanted bonewas used as a reference. The study revealed that the hard-ness of the bone was affected by the initial degradation of

implants and then reached at the same level as a referenceafter the complete degradation confirming no adverse effectson the mechanical strength of bones.40

Mg could be the ideal material for orthopedic applica-tions because of its properties like good mechanical strength,biodegradation, and biocompatibility but the rapid degrada-tion of Mg implants is an existing barrier toward its practicaluse and needs extensive investigation. Furthermore, the Mgcannot be used in load supporting implants because of itsbrittleness. Therefore, the focus of the future researchshould address the issues of rapid degradation and highbrittleness.

MG AND ITS ALLOYS

Alloying is a commonly used technique to enhance the prop-erties of different elements. The properties of these elementsare optimized by adjusting the content of impurities init. The alloying of Mg has been in practice extensively toimprove the mechanical properties and corrosion resistance.The most commonly used alloying elements are aluminum(Al), calcium (Ca), zinc (Zn), zirconium (Zr), strontium (Sr),and REEs such as yttrium (Y), gadolinium (Gd), lanthanum(La), and dysprosium (Dy). Mg-based bulk metallic glasses(BMGs) are also under investigation because of their abilitiesto control the concentrations of alloying elements above thesolubility limits. The effects of different alloying elements onthe properties of Mg have been presented in Table I. Thissection will cover a brief review of current trends in Mgalloying.

Mg–Al-based alloysAl is a commonly used alloying element in Mg and its con-centration ranges between 2 and 9 wt % in commerciallyused Mg–Al alloys.52 Mg is alloyed with Al to increase boththe strength and corrosion resistance.53,54 Alloying of Mgwith Al promotes both the solid solution strengthening andprecipitation strengthening, thus increases the strength ofMg–Al alloys.55 Along with improved strength, Mg–Al alloysalso show very good castability.54 The reason for improvedcastability of Mg–Al alloys is the lowering of temperatures ofsolidus and liquidus lines as the percentage of Al increases.55

The content of Al in Mg directly affects the corrosion rate ofMg–Al alloys and increasing the Al content typically resultsin lower corrosion rates.56–58 The reason for improved cor-rosion resistance is the formation of an insoluble Al2O3 layer,instead of Mg(OH)2, which is soluble in chloride solution.8

Recent studies have proved the enhanced corrosion resis-tance of Mg–Al alloys as compared to pure Mg in simulatedbody fluids (SBFs),57 in phosphate buffer saline (PBS),59 andsodium chloride (NaCl) solutions.56,58,60,61 Figure 2 repre-sents the comparative Tafel plots of pure Mg and threeMg–Al alloys, and results exhibit the improvement in the cor-rosion potential of all the three Mg–Al alloys as compared topure Mg. Among the three alloys, AZ31 showed the highestpotential followed by AZ61 and AZ91, respectively, afterboth the 16 and 24 days of immersion.57

JOURNAL OF BIOMEDICAL MATERIALS RESEARCH PART B: APPLIED BIOMATERIALS | MONTH 2018 VOL 000B, ISSUE 0 3

REVIEW ARTICLE

Another important consideration in the corrosion mecha-nism of Mg–Al alloys is the content of critical contaminants(Fe, Ni, and Cu).62 The content of these impurities needs to bekept under a tolerance limit to avoid any active catalyst thatcan initiate an electrochemical reaction.62 These elementsform efficient cathodes and lead to the effective galvanic cellif they are above the tolerance limit. The formation of galvaniccells results in the localized corrosion which is detrimental tothe mechanical integrity of the implants.63 For example, if theFe content is above the tolerance limit, then it can deviate thetrend toward decreasing corrosion resistance with increasingAl. The corrosion rate can increase rapidly if the Fe content isabove the tolerance limit.64 The tolerance values of above-mentioned elements vary with the type of alloys and produc-tion methods.62 In the case of Mg–Al alloys, the tolerance limitof Fe decreases with the increase in the Al content.62 There-fore, increasing the Al content can reduce the corrosion resis-tance by increasing the possibility of localized galvaniccorrosion. The tolerance limits of Fe, Ni, and Cu for Mg–Alalloys are reported in the literature.62,64

Although the mechanical strength and corrosion resistanceare positive attributes of Mg–Al alloys, there are some con-cerns regarding the biocompatibility of Al.65 The higher con-centration of Al is undesirable1 as the accumulation of Al inthe body may lead to the neuronal injuries under specific con-ditions.66 It has also been proved that Al can be harmful toosteoblasts.67 Although the higher concentration of Al is a

point of concern, the biocompatibility of Mg–Al alloys is stillunder research. A recent in vivo study of Mg–1.0Al alloy in arat for 4 weeks showed good biocompatibility of thesealloys.68 Similarly, short-term biocompatibility of Mg–Al alloyshas been proven,3,59 but long-term biocompatibility needsdetailed in vivo studies. Moreover, the safe content limit of Alin the Mg–Al alloys should be defined by more in vivo studies.

The addition of other elements like, for example, Ca, Zn, Mn,Si, and REEs can improve some of the properties of Mg–Alalloys.52,69 The addition of Ca to the Mg–Al alloy AZ91 refinedthe microstructure and increased the thermal stability of theβ-phase, thus resulting in the improvement of yield strengthand creep resistance.52 Further addition of lanthanum-rich mis-chmetal (34.22 wt% La and 65.63 wt % Ce70) to the AZ91 alloyalong with Ca enhanced the yield strength and creep resistanceat elevated temperatures.52 The addition of Zn to Mg–Alresulted in the enhanced corrosion resistance.57 The addition ofSi to AZ91 showed an increase in the creep resistance and ten-sile strength at elevated temperature, while the content of Alshowed improvement in mechanical properties only at roomtemperature.69 Studies have proved that the additions of theREEs to the Mg alloys enhance the creep resistance of thesealloys.52,69

There are few studies that discuss the in vivo response ofMg–Al alloys, particularly AZ31.71,72 Liu et al.71 investigatedthe corrosion and biosafety of AZ31 biliary stents in rabbits.After 1, 3, and 6 months of implantations, the computed

FIGURE 1. Representative histological sections of bone healing in osteotomized ribs 24 weeks after implantation (a1) titanium (a2) Mg coated

(a3) Mg uncoated. (b) Coated plate/screw system after rib fixation on the (left) osteotomized rib and (right) intact rib. (c) Schematic diagram of plate

i on the osteotomized seventh rib and plate j on the intact seventh rib. Screws i1, i3, i5, j1, j3, and j5 were selected for CT and histological

analysis.39


tomography (CT) scans were conducted to observe the corro-sion analysis and healing process. After CT scans, the stentwas removed from the body and underwent length andweight analysis. The stent maintained its shape and size after1 month of implantation. After 3 months, the stent degradedseverely and lost significant weight. After 6 months, the stentdegraded completely, and the only residue was observed.From biocompatibility point of view, the concentration ofwhite blood cells changed at the initial stages of implantationbut came back to normal after 6 months. Similarly, both theserum magnesium and urea nitrogen had an increase duringimplantation but came back to normal after 6 months. Thestudy concluded the biosafety of Mg alloy AZ31 stent butpointed out the need to improve the corrosion resistance.71 Inanother study,72 Mg and AZ31 craniofacial bone screws werecompared by implanting them in the rabbits. After 4 weeks,pure Mg screws showed the signs of degradation, and afterweeks 8 and 12, major parts of screws were degraded. On theother hand, AZ31 screws showed almost no signs of degrada-tion after week 4 and small degradation after weeks 8 and12. The new bone formation was observed in both the cases.The study proved the potential of Mg for implant applicationand also suggested the use of alloying to control the degrada-tion profile and rate.72

The mechanical properties and corrosion resistance ofthe Mg can be effectively enhanced by alloying with Al but

the biocompatibility of these alloys is a concern. There aresome studies that have proved the short-term compatibilityof Mg–Al alloys, but there is no significant proof of long-termbiocompatibility. The limitations associated with the biocom-patibility of Mg–Al alloys makes them undesirable in thebiomedical field. These alloys should not be considered as anoption in the biomedical applications without conducting thelong-term in vivo compatibility study.

Mg–Ca-based alloysCa is one of the most important minerals in the compositionof bones.73 This fact makes Ca an attracting alloying elementin Mg for biomedical applications. The alloying of Mg withCa leads to the refinement of grain sizes,74–76 which resultsin enhanced mechanical properties of Mg. In literature,Mg–Ca alloys with an optimum Ca content of 0.6 wt % exhib-ited higher bending and compressive strengths than pureMg. Above that limit, both the bending and compressivestrengths decreased with increase in Ca content.75 TheMg–Ca alloy indicated an increase in tensile strength as com-pared to pure Mg at a low content of Ca (up to 4%).77 Alongwith improved mechanical properties, the alloying of Mgwith Ca also demonstrated an improvement in corrosionresistance as compared to pure Mg up to a specific Ca con-tent limit.75–78 The corrosion resistance of Mg–Ca alloysproved to be optimum at 0.6 wt % Ca, and any further

TABLE I. Effect of Different Alloying Elements on Properties of Magnesium

Element

Improved Mechanical Properties Improved Corrosion Properties

Property References

Performance Indicator of Improved Corrosion

Resistance References

Al Yield strength 91,264 Icorr57,58

Tensile strength 55,91,264 Ecorr57,58,60,61,91

Hardness 55

Elongation 91 Mass loss 57,58

Ca Yield strength 84 Icorr78,84

Tensile strength 77,84

Bending strength 75 Ecorr76,84

Compressive strength 75,84

Hardness 76,84 Mass loss 78

Elongation 84

Zn Yield strength 65,90,91 Icorr90–92

Tensile strength 65,89–91

Compressive strength 90 Ecorr65,90–92

Hardness 90 Mass loss 5,65,90,92

Elongation 65,90,91

Sr Yield strength 79,101 Icorr101,104–106

Tensile strength 79,101 Ecorr101,104–106

Mass loss 101,104–106

Zr Yield strength 91 Icorr mass loss 91

Tensile strength 91

Compressive strength 111 Ecorr91,112,265

Elongation 91

Cu Tensile strength 118 Corrosion resistance reduced 118

Compressive strength 118

Hardness 118

REEs Yield strength 91,119 Corrosion resistance reduced in binary Mg–REEs

alloys, improved corrosion resistance in other Mg

binary alloys

79,121,122,124

Tensile strength 91,119

Hardness 119


REVIEW ARTICLE

increase in Ca content reduced the polarization resistance ofMg–Ca alloys in SBF.75 Increasing the content of Ca abovethe optimum limit effectively increase the Mg2Ca secondaryphase. An increase in the volume fraction of the Mg2Ca sec-ondary phase increases the possibility of the galvaniccell and ultimately decreases in the corrosion resistance.79

The concentration of Ca in Mg should be less than 1 wt %(optimum is 0.6 wt %) to avoid the localized corrosion phe-nomenon in the Mg–Ca alloys.

Figure 3 represents the polarization resistance of thepure Mg and Mg–Ca alloys. The Mg–Ca alloy with 0.6 wt %Ca showed the highest polarization resistance.75 Another

important indicator of the degradation is the change in pHvalue of the electrolyte. Higher the degradation, higher willbe the alkalinity of solution and pH. The pH of SBFs is 7.4and any change from this neutral pH to higher pH indicatesdegradation of material understudy. The evaluation of Mg–Ca alloys based on pH change also proved that Ca content upto 0.5–0.6 wt % shows least degradation.76 Five Mg–Caalloys were immersed in Kokubo solution for 84 h at 37�Cand pH changed least in case of Mg–0.5Ca and highest incase of Mg–10Ca. The results showed a direct relationshipbetween pH change and Ca content as increasing the Ca con-tent increased the change in pH values.76

The Mg–Ca alloys are biocompatible by nature since bothCa and Mg intake is required for the body. It has beenproved that the degradation of Mg–Ca implants inside thebody does not produce any toxic or adverse effects.78 More-over, the in vivo study of Mg–0.8Ca for 2, 4, 6, and 8 weeksproved soft tissue biocompatibility of Mg–Ca alloys.4 Thecytocompatibility test of Mg–1Ca for L-929 cells showed notoxic effects on cells,80 making it a desirable option for bio-medical implants.

The mechanical and corrosion properties of Mg–Ca alloyscan be further modified by alloying them with small propor-tions of other elements, for example, Al, Zn, and Mn. The addi-tion of Al to Mg–Ca improved the age hardening response.81

Similarly, the Mg–0.3Ca–0.3Zn also showed a better age-hardening response as compared to Mg–0.3Ca.82 The additionof Zn to Mg–Ca–Sr improved the antitumor property of theMg–Ca-based alloys for the biomedical application.83 Mg–2Ca–0.5Mn–2Zn showed better corrosion resistance as compared

FIGURE 2. Tafel plots of pure Mg and 3 Mg–Al alloys after immersion of 16 and 24 days in SBF. (a) Pure Mg, (b) AZ91D, (c) AZ61, and (d) AZ31.57

FIGURE 3. Polarization resistance of pure Mg and various Mg–Ca

alloys.75


to Mg–2Ca and Mg–4Ca in Kokubo solution. Similarly, Mg–2Ca–0.5Mn–2Zn exhibited the smallest change in pH values inKokubo solution after immersion of 144 h. Additionally, theMg–Ca alloy doped with Zn and Ca exhibited better mechani-cal properties when compared to base Mg–Ca alloys.84

The in vivo studies of Mg–Ca-based alloys can be foundin the literature.85–87 Erdmann et al.85 evaluated themechanical and degradation performance of Mg–0.8Cascrews by implanting them in the tibia bones of rabbits. Tocompare the results, 316L stainless steel screws were usedas a control. The signs of tissue formations were observedafter 1 week of surgery in case of both Mg–0.8Ca and S316Lscrews. After 3 weeks, significant tissue formation resultedin both cases. A gas was accumulated in case of Mg–0.8Cascrews while no gas was detected in 316L screws as theywere nondegradable. In terms of degradation, the Mg–0.8Cascrews had minor mass losses after 2 weeks but significantmass loss and degradation of contours were observed after8 weeks. The pullout force for Mg–0.8Ca kept on decreasingover the implantation time, while the trend was opposite for316L screws where pull out force was increasing with theimplantation time. Although Mg–0.8Ca screws showed com-parable mechanical properties to 316L screws from weeks2 to 4, the degradation severely deteriorated mechanicalstrength from week 4 to up till week 8.85 The results vali-dated two basic concerns of Mg implants. First one is theloss of mechanical integrity and second is the evolution ofgas. The similar issue of gas evolution was observed whenMg–1.0Ca–0.5Sr pins were implanted in the tibia bones ofrats.86 Although the Mg–1.0Ca–0.5Sr pins showed good tis-sue formation trends, the evolution of gas was a serious con-cern, Secondly, the pins degraded completely in 6 weeksposing serious concerns about the reducing mechanicalstrengths.86

The biocompatibility of Mg–Ca systems, along withimproved mechanical and corrosion properties, makes it adesirable option for biomedical applications. However, theissue of rapid degradation that leads to the problems ofhydrogen evolution and mechanical strength should beaddressed first.

Mg–Zn-based alloysZn is one of the most abundant elements in the body andnutritionally an essential one.65 The alloying of Mg with Znleads to both the solid-state strengthening65,74 and precipita-tion strengthening.74 The increase of Zn content in Mgresults in the improvement of the yield strength up to a cer-tain limit, that is, 6 wt % Zn.1,88 Similarly, the ultimate ten-sile strength88,89 and elongation88 increase up to 4 wt % ofZn, and then with the further increase in Zn content, theproperties start to decline. In terms of corrosion resistance,the alloying of Mg with Zn showed an improvement as com-pared to pure Mg.65,90–92 The electrochemical investigationof three Mg–Zn alloys (Mg–1Zn, Mg–5Zn, and Mg–7Zn) inSBF proved better corrosion resistance of these alloys ascompared to pure Mg with Mg–5Zn exhibiting the best cor-rosion resistance among the three. Comparable to electro-chemical results, the Mg–5Zn showed the lowest change in

pH and all three Mg–Zn alloys had less change in pH valuesof SBF as compared to pure Mg when immersed for 70 h.90

The electrochemical study of different binary alloys of Mgproved the enhanced corrosion resistance of Mg–Zn alloy inSBF when compared to pure Mg. Moreover, the change in pHof DMEM for Mg–1Zn (8.47 � 0.12) sample was lower thanthat of pure Mg (9.18 � 0.14).91 The corrosion resistance ofMg–xZn alloys depends on the volume fraction of MgxZnysecond phase, and increasing the Zn content results in theincrease of the volume fraction of MgxZny, and subsequentreduction of the corrosion resistance.88,93 Therefore, the con-tent of Zn should not exceed the optimum limit, which hasbeen proved to be 2–3 wt % Zn.88,93

The addition of Zn in the Mg can increase the tolerancelimit of the impurities like Fe, Ni, and Cu.62 The increase inthe tolerance limit of these impurities effectively reduces thepossibility of the galvanic corrosion. Similarly, Zn can alsoincrease the tolerance limit of the above elements in Mg–Alsystems reducing the chances of galvanic corrosion. How-ever, an addition of Zn to Mg–Al alloys makes them suscepti-ble to another form of localized corrosion, that is, filiformcorrosion.62

The biocompatibility of Mg–Zn alloys has been estab-lished by a number of studies.5,6,94–96 The implantation ofMg–Zn (5.62 wt % Zn) rods in the femur of 12 rabbits wasused to test the biocompatibility of these rods with the func-tioning of heart, liver, kidney, and spleen.6 The resultsshowed no adverse effects on the function of any of theseorgans.6 The biocompatibility evaluation of Mg–Zn–Zr alloyZK30 tested against hydroxyapatite (HA) as a control wit-nessed higher cell plorification in case of ZK30.94 Four Mg–Zn alloys ZSr41 (4 wt % Zn, and varying Sr wt % of 0.15,0.5, 1, and 1.5%) were investigated for the biocompatibilityby using human embryonic stem cells and the results provedan enhanced cell viability of Mg–Zn alloys as compared topure Mg with the best results showed by ZSr41 at 0.15 wt %Sr.5 The biocompatibility of Mg–Zn–Ca alloys evaluated bythe cytotoxicity assessment of adipose-derived mesenchymalstem cells in Dulbecco’s Modified Eagle Medium (DMEM)showed improved cell viability after 24 h in the case of Mg–Zn–Ca alloy’s extracts as compared to cell viability inDMEM.96

A number of additives can be used in Mg–Zn alloys tofurther improve its properties. The addition of Ca (0, 0.5,1.5, and 2.0 wt %) to Mg–4Zn showed an improvement inultimate tensile strength and elongation values up to 0.2 wt% of Ca and the properties started to decline at 0.5 wt %Ca.88 Moreover, Mg–4Zn with 0.2 wt % Ca exhibited the low-est H2 evolution and highest corrosion resistance in Hank’ssolution.88 Similar to Ca, Mg–Zn alloys showed improvementin corrosion resistance up to a certain wt % of strontium(Sr). The mass loss of ZSr41 with 0.15 wt % Sr was the leastwhen compared to ZSr41 with 0.5, 1.0, and 1.5 wt % Sr.5 Inanother study, the ZSr41 with 1 wt % Sr showed higher cor-rosion resistance compared to pure Mg in SBF.97 Themechanical properties of ZSr41 such as hardness, ultimatetensile strength, and elongation showed variation with agingtime. These values were maximum for the aging time of 8 h


REVIEW ARTICLE

at 175�C, after that the values of these properties started todecline.97

Several in vivo studies on Mg–Zn-based alloys can befound in the literature.65,98–100 Bian et al.99 evaluated thein vivo performance of Mg–Zn-based alloy with the additionof an REE, Gd. The Mg–1.8Zn–0.2Gd pins were inserted atthe knee joints of the 18 female rats. The pins kept theirshape intact in first 2 months but were completely degradedafter 6 months with some areas showing residue. Localizedcorrosion of pins instead of more uniform corrosion wasobserved during in vivo evaluations. However, the alloyexhibited good bone formation trends.99 The localized corro-sion poses a serious threat to the strength of the implant atweak sites. A long-term study100 was carried out to evaluatethe performance of Mg–Zn-based alloy. The Mg–0.96Zn–0.21Zr–0.3REEs pins implanted in the tibia bones of the rab-bits showed good degradation resistance up to 12 weeksand degradation of almost 58 and 85% after 36 and53 weeks, respectively. Moreover, the implant showed posi-tive results in term of tissue formation. However, there weresome concerns regarding the density and porosity of bones.The authors concluded that the above mentioned Mg–0.96Zn–0.21Zr–0.3REEs alloy with all its advantages shouldnot be considered for further biomedical use due to its path-ological effects on bones.100

The results proved the significance of Mg–Zn alloys interms of mechanical properties, corrosion resistance, andbiocompatibility. The biocompatible nature of Zn with excel-lent strengthening properties makes it one of the preferredalloying elements for Mg in biomedical applications. Theoptimum wt % of Zn in the binary Mg–Al alloys has provedto be 4%. The Mg–Zn alloys should be doped with other ele-ments, without exceeding the 4 wt % Zn limits, to get thedesired properties. Moreover, long-term in vivo experimentsshould be performed to get reliable data about biocompati-bility and degradation pattern.

Mg–Sr-based alloysStrontium (Sr) presents in group 2 of periodic table alongwith Mg and Ca, possesses the bone formation property sim-ilar to Mg and Ca.101,102 Sr is a vital trace element of thehuman body and 99 wt % of its body content is inbones.101,103 Sr enhances the bone formation by acceleratingthe bone tissue growth and also slows down bone resorp-tion.103,104 The alloying of Mg with Sr leads to the refine-ment of grain sizes of Mg.79,102 The newly formedmicrostructure is based on α-Mg and Mg17Sr2 phases. Theincreased content of Sr in Mg also leads to the formations ofthe second phase precipitates along with refined grainsize.102 The tensile and yield strengths of Mg–Sr alloysincrease up to 2 wt % of Sr, with any further increase willresult in a decrease in strength.79,101 The increase in theyield strength up till this limit is due to the presence of sec-ond phase precipitates.105 On the other hand, the elongationof Mg decreases after alloying with Sr. Any increase in thecontent of Mg–Sr alloys decreases the elongation of thealloy.101,105 The corrosion resistance of Mg improves whenalloyed with Sr.101,104–106 The corrosion resistance increases

up to 1.5–2 wt % Sr and any further increase of Sr contentresults in the reduction of corrosion resistance.101,106 Theincrease of Sr content above that limit can lead to theincrease of Mg17Sr2 second phase and ultimately the forma-tion of the galvanic cell. The formation of the galvanic cellreduces the corrosion resistance by increasing the possibilityof galvanic corrosion. The in vitro study by Gu et al.101

showed that the corrosion resistance of Mg–Sr alloysincreased up to 2 wt % Sr in Hank’s solution at 37�C. Thecorrosion rates calculated from weight loss measurements,hydrogen evolution, and potentiodynamic scans showed sim-ilar trends. The corrosion rates dropped up to 2 wt % Sr andthen started to increase with an increase in Sr content. It isworth mentioning that although the increase of Sr contentabove 2 wt % showed an increase in corrosion rate, it wasstill lower than pure Mg up to 3 wt % Sr. The corrosion rateexceeded the pure Mg only at 4 wt % Sr. Similar trends wereobserved with pH change readings, as Mg–Sr alloy with 2 wt% Sr showed the least change in pH values for 500 h. More-over, the in vivo experiments carried out in male mice con-firmed the formation of bone along with the degradation ofthe Mg–Sr implant.101 Similar trends in terms of degradationwere achieved by Bornapour et al.106 The experiments car-ried out in SBF with pH 7.4 at 37�C showed that the corro-sion rates of Mg–Sr alloys were less than pure Mg up to a Srcontent of 1.5 wt % with the lowest corrosion rate at 0.5 wt% Sr.106

Sr is not only biocompatible in nature but it is also bioac-tive with great potential for bone formation.103,104 Theseproperties make Sr a suitable material to alloy with Mg forimplant applications. Moreover, the biocompatibility of Mg–Sr alloys has been established by a number of stud-ies.101,103,104 Zhao et al.104 proved the nontoxic nature ofMg–Sr alloys by the cytotoxicity assessment. The cytotoxicityof MG 63 cells was evaluated in the extracts of DMEM with asupplement of 10% fetal bovine serum. Mg–0.5Sr wasselected (based on corrosion results) for cytotoxicity testand it showed no negative or harmful effect on cells.104 Simi-larly, in another study,106 cytotoxicity of HUVECs (Humanumbilical vein endothelial cells) was evaluated in F-12Kmedium with few supplements. The cell viability in theextracts of Mg–0.5Sr was lower than the control (WE43) atinitial points (days 1 and 4) but higher than the control onday 7. The results pointed toward the fact that Mg–Sr alloyhas better cell viability than WE43 when compared for longtime periods.106

The biodegradation and biocompatibility of the Mg–Sr-based alloys have also been evaluated by in vivo experi-ments.101,107,108 Gu et al.101 investigated the in vivo responseof Mg–2Sr alloy rods by implanting them in the femur ofmice. A mouse with a drilled hole in the femur without anyimplantation was used as a control. The results (3D micro-CT and 2D tomography) showed the degradation of rods andgeneration of new bone after 4 weeks of implantation. Therod maintained its shape at the center despite having localcorrosion at the surface. After 4 weeks of the implantation,the newly formed bone exhibited smooth surface and goodcontiguity with the implant surface. The thickness and


density of newly formed bone were significantly higher thanthe bone formed in the control group. Moreover, the evolu-tion of the gas was observed after the surgery. However, gasdisappeared after 1 week of the surgery. The study of serumlevels after 4 weeks showed lower than the allowable con-centration of Sr serum in the body indicating biosafety ofMg–Sr alloys.101 A similar study108 investigated the responseof Mg–1.5Sr implants with and without heat treatments inthe rabbits. The results after 8 weeks of implantation provedthe potential of Mg–Sr alloy as an implant material.108 How-ever, both above-mentioned studies are limited to4–8 weeks of in vivo evaluation. Furthermore, long-termin vivo evaluation of Mg–Sr systems is required to investi-gate any long-term biocompatibility issues or pathologicaleffects.

Along with the desired mechanical and corrosion proper-ties, Mg–Sr alloys exhibited good biocompatibility and bioac-tivity. These properties make Mg–Sr alloy a potentialcandidate for the orthopedics applications. Although bioac-tivity is an advantage of Mg–Sr, the optimum corrosion andmechanical properties for orthopedic applications is still achallenge. The ternary Mg–Sr-based alloy systems with astrong anticorrosive element may be an answer to this prob-lem. Along with the improved properties, there is a need forlong-term in vivo experiments to prove the long-term bio-compatibility of these alloys.

Mg–Zr-based alloysZirconium (Zr) is most commonly used as a grain refiner inMg.109,110 The addition of Zr in Mg leads to highly refinedequiaxed grains.79 Although Zr has a comparatively low solu-bility limit of 3.8 wt % in Mg,1 its excellent grain refinementproperty alongside good corrosion resistance79 makes it suit-able for alloying with Mg. Zr exhibits very good corrosionresistance in corrosive agents like acids, and alkalis.79 Thecorrosion resistance of Mg–Zr alloys is due to the formationof Zr–Mg double oxyhydroxide when immersed in a boratebuffer solution.79 The study of different binary alloys showedthat Mg–1Zr had highest values of yield strength, ultimate ten-sile strength, and elongation. Moreover, the hydrogen evolu-tion values of Mg–1Zr in both SBF and Hanks solution for300 h were significantly lower than the pure Mg. Similarly, onimmersing in DMEM, Mg–1Zr (8.58 � 0.05) showed less pHchange as compared to pure Mg (9.18 � 0.14).91 For Zr-basedMg alloys, the most commonly investigated system is Mg–Zr–Sr.111,112 The addition of both Zr and Sr significantly improvesthe properties of Mg.111 Zr addition leads to grain refinement,enhanced ductility, and corrosion resistance. Similarly, theaddition of Sr leads to improved compressive strength,in vitro biocompatibility, and in vivo bone formation.111 Theaddition of up to 3.8 wt % Zr in Mg will decrease the grainsize, but any addition above that limit will create unalloyed Zrparticles. On the other hand, the addition of Sr will have noeffect on grain size but will broaden the grain boundarieswith an increase in Sr content.111

The results of a recent study showed improved ultimatecompressive strengths of Mg–xZr (x = 2–5%) when com-pared to pure Mg.111 In the case of Mg–xZr–ySr (x = 2–5%,

y = 2–5%), the ultimate compressive strengths decreasedwith an increase in Sr content. However, the results showedimproved compressive strength in the case of Mg–1Zr–ySr(y = 2–5%) with an increase in Sr content.111 The corrosionresistance of Mg–Zr–Sr alloys reduced with a decrease ingrain size.111 The apparent reason for reduced corrosionresistance is the increase in broadness and roughness ofgrain boundaries with an increase in Sr content.111 More-over, the cylindrical samples were implanted in the leftfemur of the rabbits for the in vivo evaluation of the sam-ples. The results showed that Mg–1Zr–2Sr and Mg–2Zr–5Srhad the higher values of bone mineral density and bone min-eral content after implantation of 3 months when comparedto Mg–5Zr. The results also showed that in the case of Mg–5Zr and Mg–2Zr–5Sr, the new bone lack contiguity with theimplant surface. On the other hand, the newly formed bonein case of Mg–1Zr–5Sr was well integrated with the implantsurface. The Mg–1Zr–2Sr alloy showed great potential forbiomedical applications due to its suitable degradation rateand mechanical properties as well as desirable biocompati-bility both in vitro and in vivo.111

The properties of the Mg–Zr–Sr systems can be enhancedby controlling the content of impurities, like REEs.112 Anexample of such system is Mg–1Zr–2Sr–xHo, which showedimprovement in degradation time, biocompatibility, andmechanical properties by the addition of optimum content ofholmium (Ho).112 The grain size of Mg–1Zr–2Sr–1Ho waslarger than that of Mg–1Zr–2Sr, but increasing the Ho con-tent form 1 to 5 wt % showed a drastic decrease in the grainsize.112 The addition of Ho from 1 to 5 wt % showedimprovement in corrosion resistance with Mg–1Zr–2Sr–3Hobeing the most corrosion resistant. Figure 4 presents thecorrosion behavior of Mg–1Zr–2Sr–xHo alloys in terms ofpolarization, H2 evolution, mass loss, and calculated corro-sion rates. It can be seen in Figure 4 that Mg–1Zr–2Sr–3Hoexhibited the highest corrosion resistance and least H2 evo-lution. The cytotoxicity of Mg–1Zr–2Sr–xHo (x < 5 wt %)alloys was lower than that of Mg, Mg–1Zr–2Sr, and controlwhen assessed using SaSO2 osteoblast cells. The addition ofHo up to 5 wt % enhanced the growth of cells.112

Zr is one of the best options for the grain refinement ofMg alloys. The mechanical properties, corrosion resistance,and biocompatibility of resulting Mg–Zr alloys can be con-trolled by manipulating the composition with elements likeSr, Sn, Ca, and Ho. Mg–1Zr–2Sr–3Ho is the best option avail-able in literature so far.

Mg–Cu-based alloysAlthough Cu has been used to increase the bacterial resis-tance in different alloys including Stainless steel113–115 andCobalt base alloy,116 its alloying with Mg has not beenwidely investigated. The Cu is antibacterial in nature117 andan important element in bone resorption and immune sys-tem.118 A recent study explored the effect of alloying Mgwith Cu.118 The results showed that three Mg alloys Mg–0.03Cu, Mg–0.19Cu, and Mg–0.57Cu had better Vickers hard-ness and ultimate tensile strength values than pure Mg.118

The corrosion rates of Mg–Cu alloys increased heavily with


REVIEW ARTICLE

an increase in Cu concentration and the corrosion rate ofMg–0.57Cu was much higher than pure Mg when tested inHank’s solution after 3 and 7 days of immersion. The corro-sion results were also verified by pH measurement whichshows similar trends. Like immersion results, the least pHchange was observed in case of pure Mg and any addition ofCu content changed the pH at higher levels than pure Mgwhen measured for 7 days. The reason for the increased cor-rosion rate is the presence of a dominant Mg2Cu phase inthe Mg matrix that acts as a cathode and increases the gal-vanic corrosion activity.118 The Cu is known as an effectivecathode in making the galvanic cell.63 The addition of the Cuabove the tolerance limit adversely affects the corrosionresistance of the Mg by causing localized corrosion.62 Thecell viability of the Mg alloys and pure Mg was similar after24 h of incubation for MC3T3-E1 cells. However, for HUVECscells, the viability of Mg–0.57Cu was lower than pure Mgwhile the others exhibited the same viability as pureMg. Mg–0.03Cu and Mg–0.19Cu showed superior cell adhe-sion than pure Mg, and Mg–0.57Cu. It can be observed inFigure 5 that the cells adhesion after 12 h of incubation wasthe highest in case of Mg–0.03Cu for both types of cells. Thecell migration into the wounds decreases with an increase inCu concentration. Mg–0.03 Cu exhibited the best cell migra-tion among the three alloys.118

Although the degradation rate of Mg alloys increaseswith the increasing concentration of Cu, it showed excellentantibacterial nature without any adverse effect on the bio-compatibility. Mg–0.03 Cu and Mg–0.19Cu proved to be bet-ter than Mg–0.57Cu in terms of compatibility.118 Therefore,

by adjusting the concentration of Cu, optimum values fordegradation and compatibility can be achieved.

Mg–REEs-based alloysREEs are extensively studied alloying elements in Mg alloysto improve creep resistance, corrosion resistance, andmechanical strength.1,8 The REEs are divided into two mainclasses based on their solid solubility in Mg: high solid solu-bility and limited solid solubility.1

The alloying of Mg with REEs leads to the improvementof mechanical properties.119,120 A comparison of five Mg-based alloys with Al, Sn, Ca, Gd, and La showed that the twoREEs, Gd and La, refined grain size better than the otherthree elements when extruded at 450�C. These two Mg–REEs alloys showed higher ductility values than others.120 Ina recent study, the Mg–8Gd–0.6Zr witnessed a slightimprovement in mechanical properties with the addition of1 wt % La or Ce. The addition of 1 wt % Ce to Mg–8Gd–0.6Zr showed an increase of 10% in maximum yield strengthand ultimate tensile strength as compared to the basealloy.119 In terms of corrosion resistance, Mg alloying withREEs showed reduced corrosion resistance in the case ofbinary systems and Mg–REEs-based ternary alloys.121–124

However, the addition of REEs to other Mg binary allowsshowed improved corrosion by retarding the galvanic corro-sion. The comparative corrosion study of Mg–REEs showedthat the corrosion rate of Mg–RE (Re = Y, Nd, Gd, Dy) washigher than the pure Mg in NaCl. Mg–Dy has the least corro-sion rate among Mg–REEs. The change is pH values were inaccordance with corrosion rates. Mg sample showed the

FIGURE 4. Corrosion behavior of Mg–Zr–Sr–Ho alloys. (a) Polarization curves, (b) H2 evolution, (c) mass loss rates, and (d) calculated corrosion rates

based on *H2 evolution, **weight loss, ***polarization. Reprinted from Ref. 112. Copyright The author(s) 2016.


lowest change in pH among all samples. Among the four Mg–REEs, Mg–Dy exhibited the lowest change in pH again verify-ing the corrosion results.121 The corrosion rate of high purityMg in 3.5 wt % NaCl saturated with Mg(OH)2 was found lowas compared to Mg–xRE alloys (RE = Gd, Nd, Y, Ce, La—x = 0.6–5) at 25�C.122 The corrosion study of two Mg–10Gd–xZn alloys showed that the corrosion rate of the alloysincreased by increasing the content of Zn from 2 to 6% in0.5 wt % NaCl solution.123 The corrosion rate of pure Mgwas also found to be better when compared with Mg–xY(x = 2, 4, 8, 12, 15, 18) in 0.1 M NaCl solution.124

The addition of some of the REEs can significantlyimprove the corrosion resistance of Mg alloys such as Mg–Aland Mg–Zn alloys.79,125,126 The addition of Ce in high concen-tration to the Mg–Al alloys creates the Al11Ce3 phase. Thisphase acts as a cathode for galvanic cell and retards the cor-rosion of Mg matrix.79 Similarly, the addition of erbium(Er) to the Mg–Al leads to the formation of two Mg–Al–Erphases (Mg95Al3Er2 and Mg95Al2Er3) that surrounds the Mgmatrix and hence increase the corrosion resistance.79

Another important REE is Gd which has similar positiveeffects on the corrosion resistance of Mg alloys if the contentof Gd is kept under 10 wt %. The low volume fraction ofMg5Gd leads to a decrease of galvanic corrosion and resultsin more uniform corrosion.79 The presence of Nd in the Mg–Al–Mn alloy improved the corrosion resistance of the alloydue to the formation of intermetallic precipitates. These pre-cipitates acted as less noble cathodes in the galvanic cell andsuppressed the localized galvanic corrosion.79,127 Similarly,the corrosion resistance of Mg alloy AZ91 was improvedwhen Zn was replaced by 1 wt % Ce and Er. The corrosionrate reduced in both the cases and showed the best resultswith Ce.125 The corrosion resistance of Mg–6Al–0.4Mn wasimproved when it was alloyed with 4 wt % Mischmetal. TheMg–6Al–4RE–0.4Mn showed better corrosion resistance in3.5 wt % NaCl as compared to Mg–6Al–0.4Mn.126 Most ofthe REEs results in the increased corrosion rates of Mg–REEsalloys. However, when these REEs are used in other binaryMg alloys such as Mg–Al and Mg–Zn, they tend to createintermetallic phases between alloying elements, thus creat-ing a galvanic cell with these phases as cathodes. Thesecathodic phases decelerate the corrosion rate of the main Mgmatric. Moreover, these REEs in Mg alloys tend to reduce

the galvanic corrosion of Mg phases and ultimately lead touniform corrosion of Mg alloys.

The in vivo performance of Mg–REEs alloys has beenwidely studied.33,128–131 In a recent study,128 the MgYREZrscrews and Ti–6Al–4V screws were compared in vivo byimplanting these screws in the tibiae of rabbits for 4, 12, and24 weeks. The growth of bone tissues in case of Mg alloy waslow as compared to the Ti alloy screws. Moreover, the evolu-tion of the gas was observed with Mg alloy screws. The evolu-tion of gas was maximum until week 4 but after that, it kepton decreasing until week 24. The volume of the Mg screwsreduced by 25% in the first 4 weeks. After that, no significantdifference was observed in the volume of Mg screws, whenmeasured after 12 and 24 weeks. Small traces of alloying ele-ments were found in the body but that was under the allow-able limit. The study concluded that the in term of bonegeneration, the Mg alloy screws performed similar to Ti–6AL–4V screws, but the prediction of biodegradation was termedas a challenge.128 The degradation rate in vivo depends on alot of factors including the location of the implant in the body,flow properties and fluid composition at that point, loads sur-rounding the implant, and so forth. Therefore, the degradationrate and degradation pattern of an alloy at a selected pointcannot be generalized. The in vivo studies should focus on theapplication-based analysis of implants to get an accurate pic-ture of biodegradation and biocompatibility.

Another critical factor in the Mg–REEs-based alloy sys-tems is the presence of long-period stacking order (LPSO)that plays a key role in the improvement of mechanical andcorrosion properties of such systems.132 The presence ofLPSO improves the mechanical properties of the alloy by lim-iting the movement of the dislocations.133 The corrosionproperties of the Mg–REEs alloys improve due to the follow-ing reasons: (1) the ability of the oxidation film to remediaterapidly, (2) the presence of improved passive layer, and(3) the continuous distribution of LPSO structure.134

Some of the commonly investigated alloys with LPSOstructures include Mg–Gd–Zn–Zr,134–137 Mg–Dy–Zn,132,135

Mg–Zn–Y,136,137 Mg–Er–Gd–Zn,138 and Mg–Ho–Zn.139 Zhanget al.140 showed the higher corrosion resistance and signifi-cantly lower hydrogen evolution of Mg–Gd–Zn–Zr alloys withLPSO phase as compared to the Mg–Gd–Zn–Zr alloys withoutLPSO phase in both Hank’s solution and SBF.134,140–142

FIGURE 5. Adhesion of (a) MC3T3-E1 cells and (b) HUVECs in different extracts after 12 h of incubation. Reprinted from Ref. 118. Copyright Springer

Nature 2016.


REVIEW ARTICLE

Moreover, the alloys with LPSO exhibited uniform corrosionwhile the alloys without LPSO exhibited localized corro-sion.134 In another study conducted by Peng et al.,132 thecorrosion performance of Mg–Dy–Zn alloys with differentLPSO phases was evaluated. The results showed that theMg–Dy–Zn alloy with 14-H type LPSO phase has better cor-rosion performance as compared to Mg–Dy–Zn alloy with18-R type LPSO phase. The oxidation film breaks with corro-sion in both the cases, but in 14-H type LPSO the remedia-tion ability of the film is much better than the 18-R typeLPSO.132 In another study on Mg–Zn–Y, the results exhibitedthat the increase in the volume fraction of LPSO phases dete-riorated the corrosion resistance.136,137 This is due to theformation of microgalvanic cells between the Mg matrix andLPSO phase which accelerates the localized corrosion.136,137

Zhang et al. compared the mechanical properties of pure Mg,Mg–6Ho with no stacking faults, and Mg–6Ho–1Zn withstacking faults and showed that Mg–6Ho–1Zn exhibitedhigher strengths (both yield and ultimate) in tension andcompression than pure Mg and Mg–5Ho. However, Mg–6Hoshowed the highest ductility among three cases. In terms ofcorrosion properties in SBF at 37�C, Mg–6Ho–1Zn withstacking faults showed better resistance than Mg–6Ho. How-ever, both the alloys performed poorly in terms of corrosionwhen compared to pure Mg.139 Similarly, the evaluation ofmechanical properties of pure Mg, Mg–8Er with no stackingfaults and Mg–4Er–4Gd–1Zn with stacking faults showedthat Mg–4Er–4Gd–1Zn has highest yield and ultimatestrengths among three owing to the presence of stackingfaults.138 The excellent mechanical strengths and desirablecorrosion properties of the Mg–REEs with LPSO structuresmake them potential candidates for biomedical applications.

Although the Mg–REEs are not recommended in terms ofbiocompatibility,8,79 there are few studies that proved Mg–REEs as biocompatible in nature.143,144 The biocompatibilityevaluation of three Mg–RE alloys, Mg–La, Mg–Nd, and Mg–Cewitnessed no harmful effects on tissues. The cell viability ofMg–La and Mg–Nd after 1, 3, and 5 days was almost similar,but Mg–Ce showed lower viability values for the sameamount of time as compared to the other two alloys.143 Thefour novel Mg alloys with REEs showed almost the same bio-compatibility as pure Mg, which proves REEs as suitablealloying material for degradable Mg alloys for stentapplications.144

The use of REEs as alloying element in Mg-based alloyssignificantly increases mechanical properties. However, thereare some issues associated with Mg–REEs regarding corro-sion resistance and biocompatibility but if the content ofREEs is manipulated with other alloying elements the resultscould be better as in alloys with LPSO phases. The optimumconcentrations of these alloying elements need to be investi-gated to get the best compromise between mechanical prop-erties, corrosion resistance, and biocompatibility.

BMG of Mg alloysBMGs are formed by the rapid cooling of liquefied alloysbelow their glass-transition temperatures (TG).

74 The pur-pose of rapid cooling is to avoid the growth of crystalline

structures. The limited solubility of alloying elements in Mgis one of the reasons that high corrosion rate of Mg alloyscannot be addressed properly. The addition of alloying ele-ment more than the solubility limit results in the precipita-tion of additional content on cooling that subsequently leadsto galvanic corrosion at the interface of grain and grainboundaries. This precipitation of additional content due tosolubility limit can be avoided by the formation of rapidcooling of molten alloys.74 The ability to control the contentof different elements above the solubility limit makes BMGspotential candidates for biodegradable implant materials.

There are a number of studies that have proved superiorcorrosion properties (noble current potential, low currentdensity, low H2 evolution) of Mg–BMGs along suitablemechanical strengths and biocompatibility.11,145–147 Guet al.145 have studied Mg–25Zn–5Ca and Mg–30Zn–4CaBMGs and reported three times higher fracture strengthscompared to pure Mg. Moreover, both the systems exhibitednoble open-circuit potentials and better cell viability whencompared with as-rolled pure Mg. Similarly, both the BMGsystems showed less change in pH values of SBF as com-pared to pure Mg when measured for 30 days.145 Mg–35Zn–5Ca showed lower current density and noble potentialvalues as compared to as cast pure Mg in DMEM.146 TheMg–BMGs of Mg–xZn–yCa with higher content of Zn(x = 20–35) showed much lower H2 gas evolution as com-pared to Mg alloys with low Zn concentration.11 The in vivointeraction of Mg–xZn–5Ca BMG pins in the femur of the ratdemonstrated that the increase in Zn content results in areduction of H2 evolution. The H2 evolution reduced to anextremely low level at x > 29 and became negligible atx = 35.147 The effect of alloying elements on properties ofMg-based BMGs showed that an increase in the Zn contentimproves corrosion resistance. The study showed a reducedcorrosion rate when the content of Zn was increased from30 to 35 wt %. The content of Ca also caused a reduction inbiocorrosion. The Ca is a beneficial element in the biomedi-cal implants for bone formations. An important consider-ation in this regard is the Zn content that needs to beoptimized because higher Zn content will result in poor glassforming ability of Mg–xZn–yCa.148

The properties of commonly studied Mg–Zn-Ca can befurther enhanced by adding small concentrations of otherelements. The Mg–27Zn–4Ca BMG, when doped with Mn(0, 0.5, and 1 wt % Mn), resulted in the higher corrosionpotential and wider passivation regions as compared to pureMg and Mg alloy ZK60 when tested in SBF at 37�C. The threecompositions of Mn-doped Mg–Zn–Ca BMG showed muchlower H2 evolution than pure Mg in SBF at 37�C for 40 h.Like electrochemical and hydrogen evolution trends, thethree compositions of Mg–Zn–Ca BMGs doped with Mnshowed significantly lower pH changes in SBF as comparedto pure Mg when measured for 72 h at 37�C. All three BMGsexhibited better biocompatibility as compared to pure Mg.149

The study of the effect of Y on properties of Mg–27Zn–4Ca–xY (x = 0, 1, 2%) showed a significant improvement in themechanical properties of BMG. The corrosion potentials ofthe Mg–27Zn–4Ca–xY BMGs were higher than the pure


Mg. Also, H2 evolution was lower than Mg in SBF at 37�C.The cell viability of these new BMGs was higher than pureMg.150 Similarly, the addition of silver (Ag) in Mg–Zn–CaBMG showed an improvement in corrosion resistance, reduc-tion in H2 evolution in PBS and better biocompatibility whencompared to Ag-free BMGs.151

Most of the Mg-based BMGs contain transition metalsand REs because of their effective glass forming nature butmost of these elements are toxic in nature and not suitablefor biomedical applications.152 Therefore, Mg–Zn–Ca BMGsare most commonly investigated for biomedical applicationsbecause of their biocompatible nature.152

The liberty to control the concentration of a desired ele-ment above the solubility limit opened a new way of realiz-ing the dream of biodegradable implants. Furtherinvestigation is required to achieve optimum fabricationmethods and composition that give the best combination ofmechanical properties, glass forming abilities, corrosionresistance, and biocompatibility.

COATING TECHNIQUES APPLIED FOR MG ALLOYS

The biodegradation of Mg makes it a potential implant mate-rial; however, its rapid degradation is a limitation in its com-mercialization. The surface treatment has proved to be acost-effective method to control the rapid biodegradation ofMg and its alloys. The aim of surface treatment is to increasethe corrosion resistance of Mg and its alloys by coating witha biocompatible protective layer. Many coating techniqueshave been used and are also currently in practice to attainthe desired corrosion properties of Mg and its alloys. Thissection will focus on coating techniques, which have beenwidely investigated for Mg and its alloys. Table II presentsthe summary of common coatings on Mg and its alloysthrough different techniques.

Chemical conversionThe chemical conversion method is used to form metal oxidefilms on the surface of metals to enhance its passivity.153 Inthis method, the base metal is immersed in the treating solu-tion. The species from the treating solution reacts with thebase metal and form a protective layer. This newly formedprotective layer enhances the corrosion resistance of themetal surface.154 There are a number of factors that canaffect the quality of the coating, including alloy type, thecomposition of treating solution, pretreatment, posttreat-ment and experimental parameters.154 The conversion coat-ing technique has been widely used to achieve differentphosphate coatings including calcium phosphate (Ca-P),155–162 strontium phosphate,163 Zn–Ca–P164 coatings, andmagnesium phosphate165,166 on Mg and its alloys. Thismethod has also been used to coat Mg and its alloys withcerium coatings167 and magnesium fluoride (MgF2)coatings.21,78,168–179

This method is widely used to coat Ca-P layers on thesurface of Mg and its alloys.155–162 Ca-P coatings are of sig-nificant importance in orthopedic implant applicationsbecause of the formation of the HA layer, which is commonly

known as the bone mineral.180 One method to coat Ca-P onthe surface of Mg alloys is the immersion of these alloys inSBF.155,156,180 The precipitation of calcium ions with phos-phates on the surface of Mg alloys produces the Ca-P coat-ings.155 Another way to deposit a Ca-P layer is to dip Mgalloy samples in a calcium phosphate bath for a specifiedtime ranging from a few minutes to several hours.157 Severalstudies have been conducted to investigate the properties ofCa-P and HA coated Mg alloy surfaces. Hiromoto and Yama-moto159 studied the HA-coated Mg samples treated for 8, 16,and 24 h in Ca-EDTA (Ethylenediaminetetraacetic acid),KH2PO4, and sodium hydroxide (NaOH) solution with pH of7.3 at 95�C. The coated samples exhibited lower corrosioncurrent densities and approximately the same potentials(24 h coated samples showed higher potential) when com-pared to polished pure Mg sample.159 In another study, theHA-Mg(OH)2 coating was formed on pure Mg in two stepstermed as primary (HA) and secondary (Mg(OH)2) coating.In primary coating, the samples were immersed in the coat-ing bath for 2 min at 65�C and for secondary coating, the pri-mary coated samples were immersed in 10 g/L NaOHsolution for 60 min at 80�C. The corrosion current densitiesdecreased with both the coatings, and the secondary coatingshifted the corrosion potential toward nobility, indicatingmuch better corrosion resistance as compared to pureMg. The immersion tests showed that the coated samplescould slow the early corrosion attacks, although it could notcompletely eliminate corrosion.160 Yang et al.161 studied thedegradation and biocompatibility of Ca-P coated AZ31 alloyand bare AZ31 alloy rods in the thighbone of a rabbit. Boththe coated and bare alloy showed an increase in serum Mglevels of rabbits after 1 day, and 1, 4, 6, and 8 weeks ofimplantation. The serum Mg levels for both coated andnaked samples were almost at the same level for 0 h, 1 day,and 1 week. For the fourth, sixth, and eighth weeks, theserum Mg level of Ca-P coated AZ31 alloy implants was lessthan bare samples. The coated samples showed more uni-form corrosion behavior and less volume loss than the baresample.161 Similarly, Xu et al.158 studied the biocompatibilityof Mg alloy samples after immersion in the phosphating bathfor 6 min to coat a Ca-P layer on the surface. The cell viabil-ity of the Ca-P coated alloy was higher than that of the barealloy after 1, 3, and 5 days. For in vivo studies, the bare andCa-P coated samples were planted in the femur of rabbitsand the results showed that the newly formed tissuesaround Ca-P coated implant were more uniform and com-pact as compared to bare alloy implants.158

Apart from phosphate coatings, this technique is widelyused to coat MgF2 layers on the Mg alloys. The generallyused method is the boiling of samples in NaOH solutions fol-lowed by immersion in the hydrofluoric (HF) acid for a defi-nite period of time.21,78,172 The study of MgF2 coated AZ31alloy showed that the corrosion resistance of the alloyincreased with coating time. The increase was significantwhen coating time increased from 3 to 72 h. After 72 h ofcoating, the corrosion resistance reached an almost constantlevel and no significant change was observed in corrosionproperties after coating time increased further to 168 h.


REVIEW ARTICLE

TABLEII.CoatingTechniquesusedin

SurfaceTreatm

entofMganditsAlloys

Coating

Technique

Coatings

Substrates

Perform

anceIndicatorof

ImprovedCorrosionResistance

Biocompatibility/

BioactivityIndicator

References

Chemical

conversion

Calcium

phosphate/HA

Pure

Mg,AZ31,AZ61,Mg–Mn–

Znalloy,WE43

I corr,massloss,Mgionrelease,

volumeloss

Cellproliferation,in

vivo

mice,in

vivorabbits,

Mgserum

level

155–162

MgF2

Pure

Mg,ANd42,AZ31,

LAE442,LANd442,LNd42,

Mg/Al 2O

3composite,Mg–Ca

alloys,Nd2,ZK30,ZK60

Chargetransferresistance,

corrosionrate,E c

orr,H2

evolution,I corr,massloss,pH

change,polarization

resistance,volumeloss

Cellproliferation,cell

viability,

hemocompatibility,

invivorabbits

21,78,168–179,181

Zinccalcium

phosphate

AZ31

E corr,H2evolution,I corr

n/a

164

Strontium

phosphate

Pure

Mg

E corr,H2evolution,I corr


viability

163

Magnesium

phosphate

AZ31


corrosionrate,E c

orr,I corr

n/a

165,166

Cerium

AZ31

E corr,I corr

n/a

167

MAO

Calcium

phosphate/HA

AZ31,AZ91,Mg–Zn–Caalloy

E corr,H2evolution,I corr,

polarizationresistance

n/a

190–195

Other

AZ31,AZ91,Mg–Caalloy,Mg–

Li–Caalloy

Corrosionrate,E c

orr,H2

evolution,I corr,massloss,

polarizationresistance,pH

change


viability,

hemocompatibility

196–201

Anodization

Oxidelayer

AM60,AZ31,AZ91,ZK60


corrosionrate,E c

orr,I corr,Mg

ionrelease,pH

change


viability

3,59,205–209,211

Fluoridelayer

AZ31,WE43

Corrosionrate,E c

orr,I corr,mass

loss

n/a

204,210

Ion im

plantation

Zn,Ce,ZrN

,Hf,Nd,Fe,Ca,

Cr–O,Zr–O,N,Taion

implantations

Pure

Mg,AZ31,AZ91,Mg–Ca

alloys,WE43,ZK60


corrosionrate,E c

orr,I corr,Mg

ionrelease,pH

change,



viability

214–227

ECD

Calcium

phosphate/HA

AZ31,AZ91,Mg–Sralloy,Mg–

Zn–Caalloy,Mg–2Zn,Mg–

3Zn


corrosionrate,E c

orr,H2

evolution,I corr,massloss,pH

change


viability

18,229–234

Other

AZ31,Mg–1Li–1Ca,TZ51


E corr,H2evolution,I corr,Mg

ionrelease,pH

change

Cellviability,

hemocompatibility

236–240

PVD

HA

AZ31,AZ91,Mg–1Ca


E corr,I corr,massloss,


n/a

242–248

Others

AZ31,ZK60

E corr,I corr,polarization

resistance


viability

241,249


Moreover, the pH change measurements showed significantlysmaller pH change of SBF in the case of the coated sample(72 h) as compared to pure Mg for 30 days. In case of normalsaline, the pH change values were significantly different ini-tially, but over the period of 30 days, there was no differencein the pH values for both coated and uncoated Mg samples.173

The cytotoxicity study of MgF2 coated Mg–Zn–Zr screws byevaluating L-929 cells in RPMI-1640 medium with 10% fetalbovine serum showed the greater cell viability of MgF2extracts as compared to control and bare Mg sample extractsas shown in Figure 6e.175 Figure 6 also represents the mor-phology and elemental mapping of MgF2 coated screws.175

The good biocompatibility and bioactivity of MgF2 coated Mgalloys have been proved in the literature.78,174–178,181

The MgF2 coatings by conversion are often combinedwith other coatings such as polydopamine, Ca-P, andHA181–183 to get the desired corrosion properties. Renet al.182 coated Mg alloy AZ31 with Ca-P/MgF2 by firstimmersing the samples in HF acid for 24 h to coat MgF2 andthen using the combination of sol–gel and dip coatingmethods to coat a Ca-P glass layer. The composite coatedsamples exhibited better corrosion resistance compared toindividually coated MgF2 and Ca-P samples, and bare. Simi-larly, the pH change values were lowest in case of the com-posite coating when measured for 18 days. The pH changevalue of composite Ca-P/MgF2 was significantly lower thanpure Mg.182 A novel coating of HA doped with fluorine onthe surface of pure Mg was studied by Jiang et al.183 Thesamples were first immersed in HF acid for 2 h at 37�C toget a MgF2 layer and then immersed in a phosphating bathto get a fluorine doped HA coating. The novel coatingsshowed better corrosion resistance than uncoated samplesand only MgF2 coated samples.183

Chemical conversion is a relatively simple method and hasbeen successful in coating phosphate and fluoride coatings.The widely used Ca-P coatings by the conversion method havebeen successful in achieving the target of better corrosion

rates with a desirable property of HA formation. Similarly, theMgF2 coating by the conversion technique improved the cor-rosion resistance of Mg alloys without adversely affecting thebiocompatibility.

Microarc oxidationMicroarc oxidation (MAO), also known as plasma electrolyticoxidization, is one of the most common techniques used topassivate Mg surfaces.184–188 MAO setup consists of an elec-trolytic bath, the surface to be coated acts as a working elec-trode (anode), a counter electrode (cathode), and a voltagesource.180,189 The applied voltage should be greater than thedielectric breakdown voltage of the growing oxide layer usu-ally ranging from 300 to 500 V.180 In this process, plasma isgenerated and forms an oxide layer on the surface of thesubstrate. This process involves melting, the flow of the melt,and the solidification of the oxide layer. The application ofhigh currents results in the formation of small pores, usuallyof few microns, on the surface.180 The coatings produced bythis method are considered to be very stable, and good cor-rosion resistant, with the only disadvantage, is the brittle-ness of coatings.180 Ca-P and HA are the most commoncoatings on the surface of Mg and its alloys coated by thismethod.190–195

In a recent study, HA was coated on Mg alloy AZ31 byMAO at different voltages (300, 350, 400, 450, and 500 V) inan electrolyte of NaOH and C3H7CaO6P for 10 min. The high-est corrosion resistance was exhibited by the surface coatedat 400 V. Moreover, the pH change measurements for 168 hshowed that the sample treated at 400 V had the leastchange in pH value of Hank’s solution. However, there wasno significant difference in the pH values of samples treatedat 400, 450, and 500 V.190 In the following study by Tanget al., AZ31 was coated in the same bath (NaOH andC3H7CaO6P) at 400 V with varying times of 5, 10, 20, 40,and 60 min. The AZ31 sample coated with HA for 400 V and40 min exhibited the highest corrosion resistance.191 The

FIGURE 6. (a) Surface and (b) detailed morphology of the MgF2 coated screws. (c) Cross-sectional morphologies of the MgF2 coated screws and

(d) corresponding elemental mapping (A = substrate, B = coating, C = resin) of the line scan carried out along the yellow line. (e) Cell viability of

MgF2 coated surface versus bare surface and control.175


REVIEW ARTICLE

properties of MAO-treated surfaces can be enhanced withpretreatments, posttreatments, or if used with a combinationof other coating methods.196–203 Corrosion resistance andbioactivity of Mg alloy AZ31 were improved by hydrother-mally treating the MAO-treated surface. In the treating setup,the working electrode was the AZ31 specimen and the coun-ter electrode was a platinum plate. The electrolytic bath con-sisted of 1.0 M NaOH, 0.1 M glycerol and 0.1 M sodiumphosphate. The MAO-treated surface showed lower currentdensities and higher corrosion potentials as compared to thebare sample. Afterward, these MAO-treated samples weretreated hydrothermally by immersing the samples in a treat-ing bath (Ca-EDTA solution) at 90�C for different time inter-vals. The corrosion resistance further increased with thehydrothermal treatment. The in vitro cytotoxicity analysisshowed the best results for hydrothermally treated MAOcoated samples.196 Similarly, in another study, MAO wasused in combination with the dip coating to improve the cor-rosion resistance of Mg–1.34Ca alloy. For MAO coating, theelectrolytic bath was sodium-aluminate-based aqueous solu-tion, and a stainless-steel plate was used as a counter elec-trode. The MAO coated samples were then dip-coated to geta protective layer of polylactic acid. The electrochemicaltests carried out in SBF exhibited improved corrosion resis-tance for dually coated samples as compared to noncoatedand MAO coated samples.197

The MAO coated Mg surfaces are often further treatedwith different techniques including dip coating, the sol–gelmethod, hydrothermal treatment, and so forth to improvethe surface properties.196–203 The results have shown thatMAO coatings exhibit better adhesion when coupled with theabove-mentioned surface treatment techniques.

AnodizationAnodization is also a very widely used technique for the sur-face treatment of Mg alloys and its effects on corrosion resis-tance, biocompatibility, and bioactivity have been studiedextensively.204–213 A general equipment for anodization con-sists of an electrolyte, surface to be treated as a workingelectrode (anode), and a counter electrode (cathode). Apotential difference is provided from a direct current(DC) source. A protective oxide layer is formed on the sur-face of the working electrode. The quality and chemistry ofthe protective layer depend on a number of factors includ-ing, the composition of the electrolytic solution, applied volt-age, anodization time, and substrate type.189 The thicknessof the film has a direct relation to the voltage applied. Thetypical film thickness ranges from 5 to 200 μm.180 The rangeof anodization voltage depends on the electrolyte composi-tion as the species from the electrolyte form oxide layerswhich can increase blockage to electrons.180 One of thebiggest issues associated with the anodization of Mg is theformation of cracks on the coated surface.204

In a recent work by Lopez and coworkers,205 the Mgalloy AZ91D was anodized in a molybdate solution. Thisanodization showed an improvement in the corrosion resis-tance of the alloy when tested electrochemically in Ringer’ssolution at 37�C.205 The effect of varying anodization

voltages on the formation of a protective film on the Mgalloy AZ31 was studied by Jiang et al.204 The voltages vary-ing between 100 and 140 V were applied for 30 s. The elec-trochemical tests at 37�C and immersion tests in Hank’sbalanced salt solution (HBBS) showed enhanced corrosionresistance of the anodized surface. The anodized surfacewith an applied voltage of 130 V showed the least weightloss % after 1 and 4 weeks.204 Figure 7a represents the %weight loss after immersion for 1 and 4 weeks in HBBS, andFigure 7b represents the potentiodynamic curves of Mg sam-ples, where A is an untreated AZ31 sample and B, C, D, E,and F are anodized AZ31 samples at applied voltages of100, 110, 120, 130, and 140 V, respectively.204 Pompa et al.3

performed anodization on three different Mg alloys AZ31,AZ91, and ZK60 in an electrolyte based on alcohol and anorganic acid at a voltage of 10 V for 30 min at 10�C. Theelectrochemical results proved the enhanced corrosion resis-tance of all the three alloys after anodization.3 Anodization isone of the frequently used coating techniques to improve thecorrosion resistance of Mg alloys, and its effectiveness hasbeen proved in the literature. The appearance of cracks onthe anodized Mg alloy surfaces is a problem and the appar-ent solution is the combination of anodization with someother coating or treatment techniques, for example, physicalvapor deposition (PVD).

Ion implantationIon implantation is a commonly used technique for the sur-face treatment of Mg and its alloys. A typical ion implanterconsists of an ion source, which generates the ions to bedeposited on the surface of the targeted materials by accel-erating the ions to higher energies. There are some limita-tions of this technique including high processing cost and itscomplications to treat the surfaces of intricate geometries.180

The effectiveness of this technique in enhancing the corro-sion resistance of Mg alloys has been widely studied.214–227

Wu et al.215 studied the cerium ion implantation on the sur-face of pure Mg and witnessed an improvement in the corro-sion resistance of treated samples in three biologicalmediums including artificial hand sweat, Ringer’s solution,and cDMEM solution. The best results were exhibited incDMEM.215 In another study, the Mg–1Ca alloy implantedwith Zn ions in a metal vacuum arc plasma source hadshown better corrosion resistance than Mg–1Ca. Moreover,the cell viability of treated and untreated samples was simi-lar after 1 and 3 days of incubation.214 Similarly, the corro-sion potentials of Hafnium (Hf ) ion-implanted WE43, andneodymium (Nd) ion-implanted WE43 were enhanced in theSBF when compared to bare WE43. The cell viability of Hfand Nd ion-implanted WE43 was better than that ofuncoated WE43.221,222

Dual ion implantation is also in practice to improve thesurface properties of Mg.216,220,225–227 Dual ion implantationof Zr and N was performed on the Mg alloy AZ91 to form aprotective film with a thickness of approximately 80 nm.Multipurpose plasma immersion ion implanter was used forZr plasma ion implantation, with N2 gas was purged in thevacuum chamber at the flow rate of 20 sccm. The corrosion


potential of treated samples was increased as compared tountreated samples both in 0.9% NaCl and DMEM. The corro-sion rate comparison of treated and untreated samples inDMEM showed reduced corrosion rates of treated surfacesafter 1–4 weeks. Moreover, the pH change profiles in DMEMfor 28 days showed that the change was smaller in case ofZrN implanted sample as compared to unimplanted AZ91 forfirst 7 days. From weeks 1 to 4, there was no significant dif-ference in pH values of DMEM for both implanted and unim-planted samples. The deposition and settlement of corrosionproducts on the surface of the unimplanted sample resistedfurther change in pH.220

Although the single ion-implanted surfaces showed nomi-nal improvements in the corrosion resistance, the dual ionimplantation technique may effectively improve the corro-sion properties of Mg and its alloys in applications involvingnoncomplex geometries.

Electrochemical depositionElectrochemical deposition (ECD) is a common method todevelop uniform coatings on Mg alloys. In an ECD setup, theMg sample to be coated acts as a cathode, and the anode iseither graphite, stainless steel, or platinum. A solutiondepending on the desired coating acts as an electrolyte.228

Electrodeposition method is the most commonly usedtechnique to deposit Ca-P-based coatings on Mgalloys.18,229–235 Wen et al.235 studied Ca-P coating on AZ31 byECD in a solution of Ca(NO3)2, NH4H2PO4, and NaNO3 at 85�Cfor 60 min in the presence of graphite anode. The results indi-cated an increase of 0.18 V in the corrosion potential and adecrease of 10−3 A in the corrosion current in SBF at 37�C.235

However, the Ca-P coatings on Mg alloys by this method havea general problem of low adhesion and porosity. To counterthe above-mentioned problem, the concept of ECD withpulsed power was introduced.18,231 The pulsed method caneither be pulse potential method or pulse current method.229

Monasterio et al.229 performed the scanning electron micro-scope and X-ray diffraction (XRD) analysis on coatings from

four ECD methods and proved that the coatings achieved byconstant current and pulsed current method were more uni-form and adhesive as compared to coatings by constantpotential and pulsed potential method.229 Figure 8 representsthe appearances and XRD of Ca-P coatings on AZ31 using dif-ferent conditions.229 Shangguan et al.230 applied the pulsedelectrochemical deposition (PED) on Mg–1.5Sr and observedan improvement in corrosion resistance and reduction in theH2 evolution of both the as-extruded and as-cast alloy inHBSS.230 This method has also been used to obtain salinecoatings236 and layer by layer TiO2 coatings237 on Mg alloys.Both these coatings improved corrosion resistance of Mgalloys and showed good cytocompatibility. Moreover, the elec-trodeposition is a very common method to improve hydro-phobicity of Mg and its alloys.238–240

The primary advantages of ECD include low cost, uni-form, and adhesive coatings, and its ability to control thethickness and morphology of coated films by changing theparameters such as electrolyte concentration and voltages.Additionally, the issue of low adhesion can be addressedeffectively by using PED.

Physical vapor depositionPVD is a widely used method to coat thin layers on the sur-face of Mg and its alloys. The process is carried out in a highvacuum, where a solid target is energized to create plasmato be deposited on the substrate. In this technique, the posi-tively charged substrate is placed in an inert environmentwith high vacuum. The target material is negatively chargedwhich is excited by the collision of argon atoms and produceplasma. The generated plasma produces a uniform coatingon the surface of the substrate. One of the advantages ofPVD is that it involves no chemical interaction between thetarget and the substrate. Other advantages include uniformand adherent layers and its convenience to control the filmthickness.241–243

There are two commonly used PVD techniques: the firstis e-beam evaporation and the second is sputtering. For Mg

FIGURE 7. (a) % weight loss after immersion in HBBS for 1 and 4 weeks, (b) Potentiodynamic curves of (A) untreated AZ31, (B–F) are anodized AZ31

samples at 100, 110, 120, 130, and 140 V, respectively.204


REVIEW ARTICLE

substrates, the preferred technique is sputtering. In litera-ture, HA sputtered Mg alloys have been mostlyinvestigated,242–248 though there are some studies in whichAl, TO, and TiO2 sputtered Mg surfaces have also been stud-ied.241,249 In a recent study, a film of the thickness of980 � 50 nm was deposited on Mg alloy AZ91 by sputteringCa-P using a radio frequency (RF) magnetron source at500 W for 8 h. The investigation of mechanical propertieswith nanoindenter witnessed a significant increase in thevalues of nanohardness and Young’s Modulus as comparedto bare AZ91 sample.244 Similarly, an HA film with a thick-ness of 500 � 20 nm was deposited on AZ91 with a powerof 500 W using an RF magnetron sputtering gun. The elec-trochemical tests of these HA coated specimens in SBF solu-tion at a temperature of 37�C showed an improvement incorrosion resistance as the weight loss of HA sputtered sam-ples was significantly less than the uncoated sample after7 days of immersion.242 Moreover, multiple layers can alsobe sputtered on Mg substrates. Kiahosseini et al.243 sput-tered HA/ZrN on AZ91 for various times in ion beam sput-tering apparatus. The ZrN layer was first deposited on theAZ91 substrate at an initial vacuum of 2.3 × 10−3 Pa and400�C for 2 h and then HA was deposited at an initial vac-uum of 2.3 × 10−3 Pa and temperature of 300�C for180, 240, 300, 360, and 420 min. The thickness of films var-ied from 2 to 4.7 μm. The results showed an improvement inthe corrosion potentials when compared to untreated AZ91samples.243 The sputtering is not limited to only HA sput-tered layer but it has been also used to coat other biocom-patible materials on Mg alloys.241,249 A tantalum oxide(TO) thin film with a thickness of 1.48 � 0.02 μm wasdeposited on ZK60 alloy in a vacuum of 6.5 × 10−4 Pa at100 W for 3 h using a Ta target in DC sputtering gun. Theresults showed a significant increase in corrosion resistancein SBF, less change in pH values and a very good MC3T3-E1cells viability in DMEM for TO coated samples.249 Sputteringis one of the emerging coating techniques because of itsadvantages like uniform coatings, convenience in controlling

parameters, no chemical reactions involved and above all theliberty to control the different compositions of the layers byusing multiple gun sputtering system.

COMPUTATIONAL MODELS

Computational tools, such as FEA and computational fluiddynamic (CFD), are very powerful tools to investigate theperformance of almost any part undergoing mechanical load-ings and interacting fluid flow environment. The majoradvantages of computational tools include low processingtime, low cost, a high degree of accuracy, verification ofexperimental results, and profile of properties at any pointof the part. The biomedical industry is also experiencing therevolution of computational tools by analyzing the mechani-cal properties of different implants. As Mg is a potentialmaterial for the stent applications, in this section, we aregoing to focus mainly on the application of FEA in the stentindustry. Many researchers used this powerful tool to pre-dict the performance of stents made of different materialsincluding polymeric stents,250–253 Nitinol stents,254–257 stain-less steel stents,258,259 and BMG-based stents260 to name afew. FEA is in practice to analyze Mg stents for theirmechanical performance and corrosion behavior. Mg stentdesigns can be optimized using FEA by analyzing the stress–strain profiles with different parameters.43,261,262 Figure 9represents the comparison of maximum principal strains oftwo stents, one is the Magic stent used in the literature andthe second is the stent with an optimized design.262 The sim-ulation of Mg stents consists of four steps involving two cyl-inders and a stent: (1) the outer crimping cylindercompresses the stent and reduces its diameter, (2) the stentrecoils with the release of outer crimping cylinder, (3) theinner expanding cylinders increases the diameter of the stentbecause of expansion, and (4) the stent recoils after therelease of expanding cylinder. The compression and expan-sion are the most critical steps in the simulation of stents.262

CDM presented by Gastaldi et al.22 is used to study theloss of mechanical strength because of any damage. CDM is a

FIGURE 8. Appearance of Ca-P coatings on AZ31 by (a) constant potential, (b) pulsed potential, (c, d) current pulsed at different conditions, and

(e) XRD of Ca-P coatings.229


widely used mechanism to study the profiles of stents withthe degradation. Recently, Mg stent design was optimized bycomparing three different stent designs using CDM.15 In thefollowing study by Wu et al., the experimental data obtainedby immersing the AZ31 stents in the D-Hank’s solution for14 days at 37�C validated the FEA model based on CDM.45

In CDM, “D” is defined as the total damage, where D = 0means that the stent is completely intact with no damage,and D = 1 means that the stent is completely damaged.There are two corrosion mechanisms governing CDM ofstents. The first is uniform corrosion damage (DU), and thesecond is stress corrosion damage (DSC). DU represents themass loss of stents in a corrosive environment without anyapplied stress. DSC represents the corrosion that occurs inthe areas of the stent with concentrated stresses.22

D¼DU +DSC ð1Þ

The uniform corrosion and stress corrosion rates aredefined by the following equations.

_DU ¼ δULe

KU ð2Þ

_DSC ¼ LeδSC

Sσ*eq1−D

!R

ð3Þ

When σ*eq ≥ σth ≥ 0. _DSC = 0 when σ*eq ≤ σth.

In the above equations, _DU is the time derivative of uni-form corrosion, δU is the critical thickness of corrosion film,

KU is a parameter associated with the kinetics of uniformcorrosion, Le is the characteristic length of finite elements,_DSC is the time derivative of stress corrosion damage, δSC is

the characteristic dimension of stress corrosion process, σ*eqis the equivalent stress of the stress corrosion process, σth isthe equivalent stress value below which corrosion will notoccur, S and R are the parameters associated with the kinet-ics of stress corrosion.22

The model presented by the Gastaldi et al.22 wasimproved by the introduction of another parameter λe in themodel for pitting corrosion.47 The modified uniform corro-sion damage equation became:

_DU ¼ δULe

λeKU ð4Þ

The presented model is most commonly used but hassome limitations like it is specific to the properties of stentand vessel design. To address these limitations, a new corro-sion model for degradable metal stents based on ArbitraryLagrangian–Eulerian was presented.261 The new modelshowed more flexibility in analyzing the effects of corrosionand mechanical properties for a range of designs, loadings,boundary conditions, and materials.261 CFD has also beenemployed to investigate the fluid dynamics and its effect onthe degradation of Mg stents in a vascular environment.263

The computational method is one of the most effective toolsto simulate the experimental procedure. This method can beused to validate the experimental results. The mechanicalaspects of the stent can be analyzed accurately using FEA, but

FIGURE 9. Maximum principal strain of (a) magic stent and (b) optimized stent. Maximum principal strain after stent recoil (c) magic stent and

(d) optimized stent.262


REVIEW ARTICLE

the corrosion model still needs some modification to make itmore flexible and universal.

FUTURE CONSIDERATION

The primary issue related to the Mg and its alloys is fast deg-radation. The fast degradation leads to the rapid loss ofmechanical strength in case of implants and excessive H2 evo-lution. Both these issues are undesirable in a biomedicalimplant. The focus of the current research is the improvementof biocorrosion of Mg and its alloys, and it revolves aroundtwo main areas, alloying and surface coatings. The highly sen-sitive nature of targeted application makes surface coating ashort-term solution. After the degradation of the protectivelayer, the Mg surface will be exposed to the physiological con-ditions, and the possibility of H2 evolution will be very high.Although the issue of losing the mechanical integrity of Mgimplants with degradation can be addressed by delaying thedegradation process with coatings, it is not possible tocompletely eliminate the evolution of excessive H2 gas. Thefocus of the future research should be alloying of Mg for along-term solution. There is a need of alloying Mg with bio-compatible elements that can effectively increase the corro-sion resistance. In conventional alloying, there is a constraintof solubility limits of different alloying elements in Mg makingit tough to get the desired properties. The Mg-based BMGsshould be given the prime focus given its nature of controllingthe compositions above the solubility limits. There is a needof making BMGs of Mg with the elements like Zn and Ca thatare essential for the body. The focus should be on the forma-tion of BMGs that are biocompatible in nature with theimproved corrosion resistance, and without any compromiseon the “biodegradation”; the property behind the enhancedinterest in Mg for biomedical applications.

Looking at the rapid progress in the field of computationalsimulations, there is a need for improvements in the computa-tional tools for the performance prediction of Mg and itsimplants. The changing mechanical strengths with degrada-tion of implant make it hard to predict the accurate perfor-mance of implant in the body. The focus should be on thedesign of a sophisticated model that accurately predictsthe change in mechanical properties with degradation of thematerial. The other factor that needs to be considered is theenvironment in which the implant is going to be deployed.The creation of an environment, that accurately imitates theimplant surroundings, needs a thorough investigation ofboundary conditions for the model. The creation of a modelthat considers the losing strength and degradation of Mgimplant in a corrosive environment seems a potential area forfuture research, as it has not been investigated much in thepast. Additionally, the focus should be on the design of Mgspecific computational models for biomedical applications.

CONCLUSION

The recent research in the field of biomaterials proved thepotential of Mg as an ideal candidate for future bioimplants.The properties of Mg like biodegradation, biocompatibility,and good mechanical strengths make it superior to other

implant materials. The only limitation in the use of Mg asimplant material is its fast degradation rate that severelydamages the ability of Mg implants to retain its structure.The high degradation rate results in the loss of mechanicalstrength of implants in physiological conditions and implantsare no longer able to mechanically support the implant site.Alloying, surface treatment, and modification in implantdesigns are three possible routes to achieve the desiredproperties of Mg implants.

The aims of alloying Mg are to improve corrosion resis-tance, bioactivity, and mechanical strengths withoutcompromising the biocompatibility of the alloys. The use ofAl and REEs as alloying elements in Mg has been widelystudied. Although these elements improve the mechanicalproperties and corrosion resistances both individually and ina combination of other alloying elements, the long-term bio-compatibility of such alloys is questionable. There is a needof alloying Mg with elements like Ca and Zn that arerequired constituents for the body. Mg–Zn and Mg–Ca alloysystems are under investigation and efforts are being madeto improve the properties of such alloys to make them suit-able alloy materials. The introduction of Mg–BMGs mayprove a critical step in the actualization of the dream of Mgas an implant material. The ability to control the concentra-tion of a desired alloying element above the solubility limitsopened innovative ways in the research of improved proper-ties of Mg alloys. There is still a long way to achieve the goalof ideal Mg alloy for the bioimplants but the latest trends inMg alloying showed a ray of light in realizing this dream.

The other way to get the desired properties of Mg andits alloys is its surface treatment. Although the surface treat-ment will not alter the mechanical strengths of Mg implants,it can hinder the deterioration of mechanical properties byfast degradation. There are a number of surface coatingmechanism available for Mg, but so far, no technique couldeliminate the issue of rapid degradation. The targeted appli-cation of degradable implants makes it hard for the surfacecoatings to completely address the issue of fast degradation.After the coating is degraded, the Mg substrate will againexpose to physiological conditions and the possibility ofexcessive H2 evolution remains there but only delayed incase of coatings. However, once the issue of hydrogen evolu-tion is addressed by alloying, the surface treatment could bea good option for controlling the degradation time of theimplants for the targeted applications.

The design modification of implant like that of stentdesign may play a vital role in maintaining the mechanicalstrength of implants and reduce degradation due to stresscorrosion under load conditions. Finite element methods areextensively being used to study the strengths of cardiovascu-lar Mg implants with degradation. This is a cost-effective andfast method to verify the in vivo and in vitro results.Although few FEA models are in practice to mimic the degra-dation of Mg stents there is a need of a comprehensive FEAmodel that can realistically imitate the degradation of Mgstents and loss of mechanical strengths in the arteries.

Future work should focus on the achievement of biocompat-ible and bioactive Mg alloy systems with improved mechanical


and corrosion properties, and design of FEA model that caneffectively imitate the degradation of the stent and resultingloss of strengths inside the arterial system.

REFERENCES1. Chen Y, Xu Z, Smith C, Sankar J. Recent advances on the develop-

ment of magnesium alloys for biodegradable implants. Acta Bio-

mater 2014;10:4561–4573.

2. Riaz U, Rakesh L, Shabib I, Haider W. Effect of dissolution of mag-

nesium alloy AZ31 on the rheological properties of phosphate

buffer saline. J Mech Behav Biomed Mater 2018;85:201–208.

3. Pompa L, Rahman ZU, Munoz E, Haider W. Surface characteriza-

tion and cytotoxicity response of biodegradable magnesium

alloys. Mater Sci Eng C 2015;49:761–768.

4. Erdmann N, Bondarenko A, Hewicker-Trautwein M, Angrisani N,

Reifenrath J, Lucas A, Meyer-Lindenberg A. Evaluation of the soft

tissue biocompatibility of MgCa0.8 and surgical steel 316L in vivo:

A comparative study in rabbits. Biomed Eng Online 2010;9:63.

5. Cipriano AF, Zhao T, Johnson I, Guan RG, Garcia S, Liu H. In vitro

degradation of four magnesium zinc strontium alloys and their

cytocompatibility with human embryonic stem cells. J Mater Sci

Mater Med 2013;24:989–1003.

6. He Y, Tao H, Zhang Y, Jiang Y, Zhang S, Zhao C, Li J, Zhang B,

Song Y, Zhang X. Biocompatibility of bio-Mg-Zn alloy within bone

with heart, liver, kidney and spleen. Sci Bull 2009;54(3):484–491.

7. Ma E, Xu J. The glass window of opportunities. Nat Mater 2009;8:

855–857.

8. Walker J, Shadanbaz S, Woodfield TBF, Staiger MP, Dias GJ. Mag-

nesium biomaterials for orthopedic application: A review from a

biological perspective. J Biomed Mater Res Part B Appl Biomater

2014;102:1316–1331.

9. Kirkland NT. Magnesium biomaterials: Past present and future.

Corros Eng Sci Technol 2012;47:322–328.

10. Zeng RC, Li XT, Li SQ, Zhang F, Han EH. In vitro degradation of

pure Mg in response to glucose. Sci Rep 2015;5:13026.

11. Zberg B, Uggowitzer PJ, Löffler JF. MgZnCa glasses without clini-

cally observable hydrogen evolution for biodegradable implants.

Nat Mater 2009;8:887–891.

12. Hort N, Huang Y, Fechner D, Störmer M, Blawert C, Witte F,

Vogt C, Drücker H, Willumeit R, Kainer KU, Feyerabend F. Magne-

sium alloys as implant materials – Principles of property design

for Mg–RE alloys. Acta Biomater 2010;6:1714–1725.

13. Yun Y, Dong Z, Lee N, Liu Y, Xue D, Guo X, Kuhlmann J,

Doepke A, Halsall HB, Heineman W, Sundaramurthy S, Schulz MJ,

Yin Z, Shanov V, Hurd D, Nagy P, Li W, Fox C. Revolutionizing bio-

degradable metals. Mater Today 2009;12:22–32.

14. Erne P, Schier M, Resink TJ. The road to bioabsorbable stents:

Reaching clinical reality. Cardiovasc Intervent Radiol 2006;29:

11–16.

15. Wu W, Gastaldi D, Yang K, Tan L, Petrini L, Migliavacca F. Finite

element analyses for design evaluation of biodegradable magne-

sium alloy stents in arterial vessels. Mater Sci Eng B 2011;176:

1733–1740.

16. Witte F. Reprint of: The history of biodegradable magnesium

implants: A review. Acta Biomater 2015;23:S28–S40.

17. Kannan MB, Raman RKS. In vitro degradation and mechanical

integrity of calcium-containing magnesium alloys in modified-

simulated body fluid. Biomaterials 2008;29:2306–2314.

18. Wang HX, Guan SK, Wang X, Ren CX, Wang LG. In vitro degrada-

tion and mechanical integrity of Mg–Zn–Ca alloy coated with Ca-

deficient hydroxyapatite by the pulse electrodeposition process.

Acta Biomater 2010;6:1743–1748.

19. Witte F, Kaese V, Haferkamp H, Switzer E, Meyer-Lindenberg A,

Wirth CJ, Windhagen H. In vivo corrosion of four magnesium

alloys and the associated bone response. Biomaterials 2005;26:

3557–3563.

20. Song G. Control of biodegradation of biocompatable magnesium

alloys. Corros Sci 2007;49:1696–1701.

21. Riaz U, ur RZ, Asgar H, Shah U, Shabib I, Haider W. An insight into

the effect of buffer layer on the electrochemical performance of

MgF2 coated magnesium alloy ZK60. Surf Coat Technol 2018;344:

514–521.

22. Gastaldi D, Sassi V, Petrini L, Vedani M, Trasatti S, Migliavacca F.

Continuum damage model for bioresorbable magnesium alloy

devices—Application to coronary stents. J Mech Behav Biomed

Mater 2011;4:352–365.

23. Barlis P, Tanigawa J, Di Mario C. Coronary bioabsorbable magne-

sium stent: 15-month intravascular ultrasound and optical coher-

ence tomography findings. Eur Heart J 2007;28:2319–2319.

24. Pinto Slottow TL, Pakala R, Waksman R. Serial imaging and histol-

ogy illustrating the degradation of a bioabsorbable magnesium

stent in a porcine coronary artery. Eur Heart J 2008;29:314–314.

25. Kumar K, Gill RS, BatraU. Challenges and opportunities for biodegrad-

ablemagnesiumalloy implants.Mater Technol 2018;33:153–172.

26. Heublein B, Rohde R, Kaese V, Niemeyer M, Hartung W,

Haverich A. Biocorrosion of magnesium alloys: A new principle in

cardiovascular implant technology. Heart 2003;89:651–656.

27. Di Mario C, Griffiths H, Goktekin O, Peeters N, Verbist J,

Bosiers M, Deloose K, Heublein B, Rohde R, Kasese V, Ilsley C,

Erbel R. Drug-eluting bioabsorbable magnesium stent. J Interv

Cardiol 2004;17:391–395.

28. Waksman R, Pakala R, Kuchulakanti PK, Baffour R, Hellinga D,

Seabron R, Tio FO, Wittchow E, Hartwig S, Harder C, Rohde R,

Heublein B, Andreae A, Waldmann K-H, Haverich A. Safety and

efficacy of bioabsorbable magnesium alloy stents in porcine coro-

nary arteries. Catheter Cardiovasc Interv 2006;68:607–617.

29. Zartner P, Cesnjevar R, Singer H, Weyand M. First successful implan-

tation of a biodegradable metal stent into the left pulmonary artery of

a pretermbaby. Catheter Cardiovasc Interv 2005;66:590–594.

30. Erbel R, Di Mario C, Bartunek J, Bonnier J, de Bruyne B, Eberli FR,

Erne P, Haude M, Heublein B, Horrigan M, Ilsley C, Böse D,

Koolen J, Lüscher TF, Weissman N, Waksman R. PROGRESS-AMS

(clinical performance and angiographic results of coronary stent-

ing with absorbable metal stents) investigators. Temporary scaf-

folding of coronary arteries with bioabsorbable magnesium

stents: A prospective, non-randomised multicentre trial. Lancet

2007;369:1869–1875.

31. Haude M, Erbel R, Erne P, Verheye S, Degen H, Böse D,

Vermeersch P, Wijnbergen I, Weissman N, Prati F, Waksman R,

Koolen J. Safety and performance of the drug-eluting absorbable

metal scaffold (DREAMS) in patients with de-novo coronary

lesions: 12 month results of the prospective, multicentre, first-in-

man BIOSOLVE-I trial. Lancet 2013;381:836–844.

32. Lim GB. Interventional cardiology: DREAMS of a bioabsorbable

stent coming true. Nat Rev Cardiol 2013;10:120.

33. Schaller B, Saulacic N, Beck S, Imwinkelried T, Goh BT,

Nakahara K, Hofstetter W, Iizuka T. In vivo degradation of a new

concept of magnesium-based rivet-screws in the minipig mandib-

ular bone. Mater Sci Eng C 2016;69:247–254.

34. Myrissa A, Braeuer S, Martinelli E, Willumeit-Römer R,

Goessler W, Weinberg AM. Gadolinium accumulation in organs of

Sprague–Dawley® rats after implantation of a biodegradable

magnesium-gadolinium alloy. Acta Biomater 2017;48:521–529.

35. Angrisani N, Reifenrath J, Zimmermann F, Eifler R, Meyer-

Lindenberg A, Vano-Herrera K, Vogt C. Biocompatibility and deg-

radation of LAE442-based magnesium alloys after implantation of

up to 3.5 years in a rabbit model. Acta Biomater 2016;44:355–365.

36. Schaller B, Saulacic N, Imwinkelried T, Beck S, Liu EWY, Gralla J,

Nakahara K, Hofstetter W, Iizuka T. In vivo degradation of magne-

sium plate/screw osteosynthesis implant systems: Soft and hard

tissue response in a calvarial model in miniature pigs. J Cranio

Maxillofac Surg 2016;44:309–317.

37. Song B, Li W, Chen Z, Fu G, Li C, Liu W, Li Y, Qin L, Ding Y. Biome-

chanical comparison of pure magnesium interference screw and

polylactic acid polymer interference screw in anterior cruciate liga-

ment reconstruction—A cadaveric experimental study. J Orthop

Transl 2017;8:32–39.

38. Draxler J, Martinelli E, Weinberg AM, Zitek A, Irrgeher J,

Meischel M, Stanzl-Tschegg SE, Mingler B, Prohaska T. The poten-

tial of isotopically enriched magnesium to study bone implant

degradation in vivo. Acta Biomater 2017;51:526–536.

39. Schaller B, Saulacic N, Beck S, Imwinkelried T, Liu EWY,

Nakahara K, Hofstetter W, Iizuka T. Osteosynthesis of partial rib


REVIEW ARTICLE

osteotomy in a miniature pig model using human standard-sized

magnesium plate/screw systems: Effect of cyclic deformation on

implant integrity and bone healing. J Craniomaxillofac Surg 2017;

45:862–871.

40. Meischel M, Hörmann D, Draxler J, Tschegg EK, Eichler J,

Prohaska T, Stanzl-Tschegg SE. Bone-implant degradation and

mechanical response of bone surrounding Mg-alloy implants.

J Mech Behav Biomed Mater 2017;71:307–313.

41. Zhu YQ, Yang K, Edmonds L, Wei L-M, Zheng R, Cheng R-Y,

Cui W-G, Cheng Y-S. Silicone-covered biodegradable magnesium-

stent insertion in the esophagus: A comparison with plastic stents.

Therap Adv Gastroenterol 2017;10:11–19.

42. Galvin E, Cummins C, Yoshihara S, Mac Donald BJ, Lally C. Plastic

strains during stent deployment have a critical influence on the

rate of corrosion in absorbable magnesium stents. Med Biol Eng

Comput 2016;55:1261–1275.

43. Li J, Zheng F, Qiu X, Wan P, Tan L, Yang K. Finite element ana-

lyses for optimization design of biodegradable magnesium alloy

stent. Mater Sci Eng C 2014;42:705–714.

44. Bartosch M, Schubert S, Berger F. Magnesium stents – fundamen-

tals, biological implications and applications beyond coronary

arteries. BioNanoMaterials 2015;16:3–17.

45. Wu W, Chen S, Gastaldi D, Petrini L, Yang K, Tan L, Migliavacca F.

Experimental data confirm numerical modeling of the degradation

process ofmagnesiumalloys stents. ActaBiomater 2013;9:8730–8739.

46. Petrini L, Wu W, Gastaldi D, Altomare L, Farè S, Migliavacca F,

Demir AG, Previtali B, Vedani M. Development of biodegradable

magnesium alloy stents with coating. Fract Struct Integr 2014;0:

364–375.

47. Grogan JA, O’Brien BJ, Leen SB, McHugh PE. A corrosion model

for bioabsorbable metallic stents. Acta Biomater 2011;7:3523–3533.

48. Boland EL, Shine R, Kelly N, Sweeney CA, McHugh PE. A review

of material degradation modelling for the analysis and design of

bioabsorbable stents. Ann Biomed Eng 2016;44:341–356.

49. Grogan JA, Leen SB, McHugh PE. Computational micromechanics

of bioabsorbable magnesium stents. J Mech Behav Biomed Mater

2014;34:93–105.

50. Wang X, Cui F, Li J, Zhao X. Mechanical analysis on a new type of

biodegradable magnesium-alloy stent. Sheng Wu Yi Xue Gong

Cheng Xue Za Zhi 2009;26:338–341.

51. Chen H, Liu X, Yuan G, Zhang L, Li Z, Luo Q, Lin F. Finite element

analysis for compression and expansion behavior of magnesium

stent. Zhongguo Yi Liao Qi Xie Za Zhi 2014;38:161–164.

52. Wenwen D, Yangshan S, Xuegang M, Feng X, Min Z, Dengyun W.

Microstructure and mechanical properties of Mg–Al based alloy

with calcium and rare earth additions. Mater Sci Eng A 2003;

356:1–7.

53. Brar HS, Platt MO, Sarntinoranont M, Martin PI, Manuel MV. Mag-

nesium as a biodegradable and bioabsorbable material for medi-

cal implants. JOM 2009;61:31–34.

54. Bamberger M, Dehm G. Trends in the development of new Mg

alloys. Annu Rev Mat Res 2008;38:505–533.

55. Witte F, Hort N, Vogt C, Cohen S, Kainer KU, Willumeit R,

Feyerabend F. Degradable biomaterials based on magnesium cor-

rosion. Curr Opin Solid State Mater Sci 2008;12:63–72.

56. Esmaily M, Blücher DB, Svensson JE, Halvarsson M,

Johansson LG. New insights into the corrosion of magnesium

alloys—The role of aluminum. Scr Mater 2016;115:91–95.

57. Wen Z, Wu C, Dai C, Yang F. Corrosion behaviors of Mg and its

alloys with different Al contents in a modified simulated body

fluid. J Alloys Compd 2009;488:392–399.

58. Pardo A, Merino MC, Coy AE, Arrabal R, Viejo F, Matykina E. Cor-

rosion behaviour of magnesium/aluminium alloys in 3.5wt.%

NaCl. Corros Sci 2008;50:823–834.

59. Rahman ZU, Pompa L, Haider W. Electrochemical characterization

and in-vitro bio-assessment of AZ31B and AZ91E alloys as biode-

gradable implant materials. J Mater Sci Mater Med 2015;26:217.

60. Singh IB, Singh M, Das S. A comparative corrosion behavior of

Mg, AZ31 and AZ91 alloys in 3.5% NaCl solution. J Magnes Alloy

2015;3:142–148.

61. Wang L, Shinohara T, Zhang BP. Corrosion behavior of Mg, AZ31,

and AZ91 alloys in dilute NaCl solutions. J Solid State Electrochem

2010;14:1897–1907.

62. Song GL, Atrens A. Corrosion mechanisms of magnesium alloys.

Adv Eng Mater 1999;1:11–33.

63. Ghali E, Dietzel W, Kainer K-U. General and localized corrosion of

magnesium alloys: A critical review. J Mater Eng Perform 2004;

13:7–23.

64. Liu M, Uggowitzer PJ, Nagasekhar AV, Schmutz P, Easton M,

Song G-L, Atrens A. Calculated phase diagrams and the corrosion

of die-cast Mg–Al alloys. Corros Sci 2009;51:602–619.

65. Zhang S, Zhang X, Zhao C, Li J, Song Y, Xie C, Tao H, Zhang Y,

He Y, Jiang Y, Bian Y. Research on an Mg–Zn alloy as a degrad-

able biomaterial. Acta Biomater 2010;6:626–640.

66. El-Rahman SSA. Neuropathology of aluminum toxicity in rats (glu-

tamate and GABA impairment). Pharmacol Res 2003;47:189–194.

67. Ku CH, Pioletti DP, Browne M, Gregson PJ. Effect of different Ti–

6Al–4V surface treatments on osteoblasts behaviour. Biomaterials

2002;23:1447–1454.

68. Miura C, Shimizu Y, Imai Y, Mukai T, Yamamoto A, Sano Y,

Ikeo N, Isozaki S, Takahashi T, Oikawa M, Kumamoto H, Tachi M.

In vivo corrosion behaviour of magnesium alloy in association

with surrounding tissue response in rats. Biomed Mater 2016;11:

025001 Available from: http://stacks.iop.org/1748-605X/11/i=2/

a=025001?key=crossref.e54bca89908e340e2d73a0da50455bb5.

69. Meshinchi Asl K, Tari A, Khomamizadeh F. The effect of different

content of Al, RE and Si element on the microstructure, mechani-

cal and creep properties of Mg–Al alloys. Mater Sci Eng A 2009;

523:1–6.

70. Tong LB, Zhang QX, Jiang ZH, Zhang JB, Meng J, Cheng LR,

Zhang HJ. Microstructures, mechanical properties and corrosion

resistances of extruded Mg–Zn–Ca–xCe/La alloys. J Mech Behav

Biomed Mater 2016;62:57–70.

71. Liu Y, Zheng S, Li N, Guo H, Zheng Y, Peng J. In vivo response of

AZ31 alloy as biliary stents: A 6 months evaluation in rabbits. Sci

Rep 2017;7:40184.

72. Henderson SE, Verdelis K, Maiti S, Pal S, Chung WL, Chou D-T,

Kumta PN, Almarza AJ. Magnesium alloys as a biomaterial for

degradable craniofacial screws. Acta Biomater 2014;10:2323–2332.

73. Agna JW, Knowles HC, Alverson G. The mineral content of Nor-

mal human bone 1. J Clin Invest 1958;37:1357–1361.

74. Gu XN, Li SS, Li XM, Fan YB. Magnesium based degradable bio-

materials: A review. Front Mater Sci 2014;8:200–218.

75. Wan Y, Xiong G, Luo H, He F, Huang Y, Zhou X. Preparation and

characterization of a new biomedical magnesium–calcium alloy.

Mater Des 2008;29:2034–2037.

76. Rad HRB, Idris MH, Kadir MRA, Farahany S. Microstructure analy-

sis and corrosion behavior of biodegradable Mg–Ca implant

alloys. Mater Des 2012;33:88–97.

77. Salahshoor M, Guo Y. Biodegradable orthopedic magnesium-

calcium (MgCa) alloys, processing, and corrosion performance.

Materials 2012;5:135–155.

78. Drynda A, Hassel T, Hoehn R, Perz A, Bach F-W, Peuster M. Devel-

opment and biocompatibility of a novel corrodible fluoride-coated

magnesium-calcium alloy with improved degradation kinetics and

adequate mechanical properties for cardiovascular applications.

J Biomed Mater Res Part A 2009;93:763–775.

79. Ding Y, Wen C, Hodgson P, Li Y, Hauschka PV, Hannigan RE,

Oho A, Kaneko T, Chiba M, Inaba Y, Kurokawa Y, Heublein B,

Andreae A, Waldmann K-H, Haverich A. Effects of alloying ele-

ments on the corrosion behavior and biocompatibility of biode-

gradable magnesium alloys: A review. J Mater Chem B 2014;2:

1912–1933.

80. Li Z, Gu X, Lou S, Zheng Y. The development of binary Mg–Ca

alloys for use as biodegradable materials within bone. Biomate-

rials 2008;29:1329–1344.

81. Jayaraj J, Mendis CL, Ohkubo T, Oh ishi K, Hono K. Enhanced pre-

cipitation hardening of Mg–Ca alloy by Al addition. Scr Mater

2010;63:831–834.

82. Oh J, Ohkubo T, Mukai T, Hono K. TEM and 3DAP characterization

of an age-hardened Mg–Ca–Zn alloy. Scr Mater 2005;53:675–679.

83. Wu Y, He G, Zhang Y, Liu Y, Li M, Wang X, Li N, Li K, Zheng G,

Zheng Y, Yin Q. Unique antitumor property of the Mg-Ca-Sr alloys

with addition of Zn. Sci Rep 2016;6:21736.

84. Bakhsheshi-Rad HR, Idris MH, Abdul-Kadir MR, Ourdjini A,

Medraj M, Daroonparvar M, Hamzah E. Mechanical and bio-


http://stacks.iop.org/1748-605X/11/i

corrosion properties of quaternary Mg–Ca–Mn–Zn alloys com-

pared with binary Mg–Ca alloys. Mater Des 2014;53:283–292.

85. Erdmann N, Angrisani N, Reifenrath J, Lucas A, Thorey F,

Bormann D, Meyer-Lindenberg A. Biomechanical testing and deg-

radation analysis of MgCa0.8 alloy screws: A comparative in vivo

study in rabbits. Acta Biomater 2011;7:1421–1428.

86. Berglund IS, Jacobs BY, Allen KD, Kim SE, Pozzi A, Allen JB,

Manuel MV. Peri-implant tissue response and biodegradation per-

formance of a Mg–1.0Ca–0.5Sr alloy in rat tibia. Mater Sci Eng C

2016;62:79–85.

87. Makkar P, Sarkar SK, Padalhin AR, Moon B-G, Lee YS, Lee BT. In

vitro and in vivo assessment of biomedical Mg-Ca alloys for bone

implant applications. J Appl Biomater Funct Mater 2018;16:

126–136. Available from: http://www.ncbi.nlm.nih.gov/

pubmed/29607729.

88. Zhang BP, Wang Y, Geng L. Research on Mg-Zn-Ca alloy as

degradable biomaterial. Biomater - Phys Chem 2011;183–204.

89. Zhang B, Hou Y, Wang X, Wang Y, Geng L. Mechanical properties,

degradation performance and cytotoxicity of Mg–Zn–Ca biomedi-

cal alloys with different compositions. Mater Sci Eng C 2011;31:

1667–1673.

90. Cai S, Lei T, Li N, Feng F. Effects of Zn on microstructure, mechani-

cal properties and corrosion behavior of Mg–Zn alloys. Mater Sci

Eng C 2012;32:2570–2577.

91. Gu X, Zheng Y, Cheng Y, Zhong S, Xi T. In vitro corrosion and bio-

compatibility of binary magnesium alloys. Biomaterials 2009;30:

484–498.

92. Zhang S, Li J, Song Y, Zhao C, Zhang X, Xie C, Zhang Y, Tao H,

He Y, Jiang Y, Bian Y. In vitro degradation, hemolysis and

MC3T3-E1 cell adhesion of biodegradable Mg–Zn alloy. Mater Sci

Eng C 2009;29:1907–1912.

93. Song Y, Han EH, Shan D, Yim CD, You BS. The effect of Zn con-

centration on the corrosion behavior of Mg–xZn alloys. Corros Sci

2012;65:322–330.

94. Huan ZG, Leeflang MA, Zhou J, Fratila-Apachitei LE, Duszczyk J. In

vitro degradation behavior and cytocompatibility of Mg-Zn-Zr

alloys. J Mater Sci Mater Med 2010;21:2623–2635.

95. Li J, Song Y, Zhang S, Zhao C, Zhang F, Zhang X, Cao L, Fan Q,

Tang T. In vitro responses of human bone marrow stromal cells to

a fluoridated hydroxyapatite coated biodegradable Mg-Zn alloy.

Biomaterials 2010;31:5782–5788.

96. Fazel Anvari-Yazdi A, Tahermanesh K, Hadavi SMM, Talaei-

Khozani T, Razmkhah M, Abed SM, Mohtasebi MS. Cytotoxicity

assessment of adipose-derived mesenchymal stem cells on syn-

thesized biodegradable Mg-Zn-Ca alloys. Mater Sci Eng C 2016;69:

584–597.

97. Guan R, Cipriano AF, Zhao Z, Lock J, Tie D, Zhao T, Cui T, Liu H.

Development and evaluation of a magnesium-zinc-strontium alloy

for biomedical applications? Alloy processing, microstructure,

mechanical properties, and biodegradation. Mater Sci Eng C 2013;

33:3661–3669.

98. Wang X-H, Ni J-S, Cao N-L, Yu S, Chen Y-G, Zhang S-X, Gu B-J,

Yan J. In vivo evaluation of Mg-6Zn and titanium alloys on colla-

gen metabolism in the healing of intestinal anastomosis. Sci Rep

2017;7:44919.

99. Bian D, Deng J, Li N, Chu X, Liu Y, Li W, Cai H, Xiu P, Zhang Y,

Guan Z, Zheng Y, Kou Y, Jiang B, Chen R. In vitro and in vivo stud-

ies on biomedical magnesium low-alloying with elements gadolin-

ium and zinc for orthopedic implant applications. ACS Appl Mater

Interfaces 2018;10:4394–4408.

100. Dziuba D, Meyer-Lindenberg A, Seitz JM, Waizy H, Angrisani N,

Reifenrath J. Long-term in vivo degradation behaviour and bio-

compatibility of the magnesium alloy ZEK100 for use as a biode-

gradable bone implant. Acta Biomater 2013;9:8548–8560.

101. Gu XN, Xie XH, Li N, Zheng YF, Qin L. In vitro and in vivo studies

on a Mg–Sr binary alloy system developed as a new kind of biode-

gradable metal. Acta Biomater 2012;8:2360–2374.

102. Brar HS, Wong J, Manuel MV. Investigation of the mechanical and

degradation properties of Mg–Sr and Mg–Zn–Sr alloys for use as

potential biodegradable implant materials. J Mech Behav Biomed

Mater 2012;7:87–95.

103. Jiang W, Cipriano AF, Tian Q, Zhang C, Lopez M, Sallee A, Lin A,

Cortez Alcaraz MC, Wu Y, Zheng Y, Liu H. In vitro evaluation of

MgSr and MgCaSr alloys via direct culture with bone marrow

derived mesenchymal stem cells. Acta Biomater 2018;72:407–423.

104. Zhao C, Pan F, Zhang L, Pan H, Song K, Tang A. Microstructure,

mechanical properties, bio-corrosion properties and cytotoxicity

of as-extruded Mg-Sr alloys. Mater Sci Eng C 2017;70:1081–1088.

105. Bornapour M, Celikin M, Cerruti M, Pekguleryuz M. Magnesium

implant alloy with low levels of strontium and calcium: The third ele-

ment effect and phase selection improve bio-corrosion resistance

andmechanical performance.Mater Sci Eng C 2014;35:267–282.

106. Bornapour M, Muja N, Shum-Tim D, Cerruti M, Pekguleryuz M.

Biocompatibility and biodegradability of Mg–Sr alloys: The forma-

tion of Sr-substituted hydroxyapatite. Acta Biomater 2013;9:

5319–5330.

107. Tie D, Guan R, Liu H, Cipriano A, Liu Y, Wang Q, Huang Y, Hort N.

An in vivo study on the metabolism and osteogenic activity of

bioabsorbable Mg–1Sr alloy. Acta Biomater 2016;29:455–467.

108. Dong J, Tan L, Yang J, Wang Y, Chen J, Wang W, Zhao D, Yang K.

In vitro and in vivo studies on degradation and bone response of

Mg-Sr alloy for treatment of bone defect. Mater Technol 2018;33:

387–397.

109. Qian M, StJohn DH, Frost MT. Characteristic zirconium-rich coring

structures in Mg–Zr alloys. Scr Mater 2002;46:649–654.

110. Qian M, Das A. Grain refinement of magnesium alloys by zirco-

nium: Formation of equiaxed grains. Scr Mater 2006;54:881–886.

111. Li Y, Wen C, Mushahary D, Sravanthi R, Harishankar N, Pande G,

Hodgson P. Mg–Zr–Sr alloys as biodegradable implant materials.

Acta Biomater 2012;8:3177–3188.

112. Ding Y, Lin J, Wen C, Zhang D, Li Y. Mechanical properties,

in vitro corrosion and biocompatibility of newly developed biode-

gradable Mg-Zr-Sr-Ho alloys for biomedical applications. Sci Rep

2016;6:31990.

113. Ren L, Yang K, Guo L, Chai H. Preliminary study of anti-infective

function of a copper-bearing stainless steel. Mater Sci Eng C 2012;

32:1204–1209.

114. Chai H, Guo L, Wang X, Fu Y, Guan J, Tan L, Ren L, Yang K. Anti-

bacterial effect of 317L stainless steel contained copper in preven-

tion of implant-related infection in vitro and in vivo. J Mater Sci

Mater Med 2011;22:2525–2535.

115. Hong IT, Koo CH. Antibacterial properties, corrosion resistance

and mechanical properties of Cu-modified SUS 304 stainless steel.

Mater Sci Eng A 2005;393:213–222.

116. Wang S, Yang C, Ren L, Shen M, Yang K. Study on antibacterial

performance of Cu-bearing cobalt-based alloy. Mater Lett 2014;

129:88–90.

117. Dan ZG, Ni HW, Xu BF, Xiong J, Xiong PY. Microstructure and

antibacterial properties of AISI 420 stainless steel implanted by

copper ions. Thin Solid Films 2005;492:93–100.

118. Liu C, Fu X, Pan H, Wan P, Wang L, Tan L, Wang K, Zhao Y,

Yang K, Chu PK. Biodegradable Mg-Cu alloys with enhanced oste-

ogenesis, angiogenesis, and long-lasting antibacterial effects. Sci

Rep 2016;6:27374.

119. Peng Q, Wang J, Wu Y, Meng J, Wang L. The effect of La or Ce on

ageing response and mechanical properties of cast Mg–Gd–Zr

alloys. Mater Charact 2008;59:435–439.

120. Stanford N, Barnett MR. The origin of “rare earth” texture devel-

opment in extruded Mg-based alloys and its effect on tensile duc-

tility. Mater Sci Eng A 2008;496:399–408.

121. Zhao X, Shi L, Xu J. A comparison of corrosion behavior in saline

environment: Rare earth metals (Y, Nd, Gd, Dy) for alloying of biode-

gradablemagnesiumalloys. JMater Sci Technol 2013;29:781–787.

122. Shi Z, Cao F, Song GL, Liu M, Atrens A. Corrosion behaviour in

salt spray and in 3.5% NaCl solution saturated with Mg(OH)2 of

as-cast and solution heat-treated binary Mg–RE alloys: RE=Ce, La,

Nd, Y, Gd. Corros Sci 2013;76:98–118.

123. Srinivasan A, Blawert C, Huang Y, Mendis CL, Kainer KU, Hort N.

Corrosion behavior of Mg-Gd-Zn based alloys in aqueous NaCl

solution. J Magnes Alloy 2014;2:245–256.

124. Sudholz AD, Gusieva K, Chen XB, Muddle BC, Gibson MA,

Birbilis N. Electrochemical behaviour and corrosion of Mg–Y

alloys. Corros Sci 2011;53:2277–2282.

125. Rosalbino F, Angelini E, De Negri S, Saccone A, Delfino S. Electro-

chemical behaviour assessment of novel Mg-rich Mg–Al–RE alloys

(RE=Ce, Er). Intermetallics 2006;14:1487–1492.


REVIEW ARTICLE

http://www.ncbi.nlm.nih.gov/pubmed/29607729

http://www.ncbi.nlm.nih.gov/pubmed/29607729

126. Liu N, Wang J, Wu Y, Wang L. Electrochemical corrosion behavior

of cast Mg–Al–RE–Mn alloys in NaCl solution. J Mater Sci 2008;43:

2550–2554.

127. Liu N, Wang J, Wang L, Wu Y, Wang L. Electrochemical corrosion

behavior of Mg–5Al–0.4Mn–xNd in NaCl solution. Corros Sci 2009;

51:1328–1333.

128. Diekmann J, Bauer S, Weizbauer A, Willbold E, Windhagen H,

Helmecke P, Lucas A, Reifenrath J, Nolte I, Ezechieli M. Examina-

tion of a biodegradable magnesium screw for the reconstruction

of the anterior cruciate ligament: A pilot in vivo study in rabbits.

Mater Sci Eng C 2016;59:1100–1109.

129. Chou D-T, Hong D, Saha P, Ferrero J, Lee B, Tan Z, Dong Z,

Kumta PN. In vitro and in vivo corrosion, cytocompatibility and

mechanical properties of biodegradable Mg–Y–Ca–Zr alloys as

implant materials. Acta Biomater 2013;9:8518–8533.

130. Hänzi AC, Gerber I, Schinhammer M, Löffler JF, Uggowitzer PJ.

On the in vitro and in vivo degradation performance and biologi-

cal response of new biodegradable Mg–Y–Zn alloys. Acta Bioma-

ter 2010;6:1824–1833.

131. Castellani C, Lindtner RA, Hausbrandt P, Tschegg E, Stanzl-

Tschegg SE, Zanoni G, Beck S, Weinberg A-M. Bone–implant

interface strength and osseointegration: Biodegradable magne-

sium alloy versus standard titanium control. Acta Biomater 2011;

7:432–440.

132. Peng Q, Guo J, Fu H, Cai X, Wang Y, Liu B, Xu Z. Degradation

behavior of Mg-based biomaterials containing different long-

period stacking ordered phases. Sci Rep 2015;4:3620.

133. Suzuki M, Kimura T, Koike J, Maruyama K. Strengthening effect of

Zn in heat resistant Mg–Y–Zn solid solution alloys. Scr Mater

2003;48:997–1002. Available from: https://www.sciencedirect.com/

science/article/pii/S1359646202005900.

134. Zhang X, Ba Z, Wang Q, Wu Y, Wang Z, Wang Q. Uniform corro-

sion behavior of GZ51K alloy with long period stacking ordered

structure for biomedical application. Corros Sci Pergamon 2014;

88:1–5.

135. BI G, JIANG J, ZHANG F, FANG D, LI Y, MA Y, HAO Y. Microstruc-

ture evolution and corrosion properties of Mg-Dy-Zn alloy during

cooling after solution treatment. J Rare Earths 2016;34:931–937.

136. Li CQ, Xu DK, Zeng ZR, Wang BJ, Sheng LY, Chen X-B, Han EH.

Effect of volume fraction of LPSO phases on corrosion and

mechanical properties of Mg-Zn-Y alloys. Mater Des 2017;121:

430–441. Available from: https://www.sciencedirect.com/science/

article/pii/S0264127517302137.

137. Zhao X, Shi L, Xu J. Mg–Zn–Y alloys with long-period stacking

ordered structure: in vitro assessments of biodegradation behav-

ior. Mater Sci Eng C 2013;33:3627–3637.

138. Xu C, Zhang J, Liu S, Jing Y, Jiao Y, Xu L, Zhang L, Jiang F,

Zhang M, Wu R. Microstructure, mechanical and damping proper-

ties of Mg–Er–Gd–Zn alloy reinforced with stacking faults. Mater

Des 2015;79:53–59.

139. Zhang L, Zhang J, Xu C, Jing Y, Zhuang J, Wu R, Zhang M. Forma-

tion of stacking faults for improving the performance of biode-

gradable Mg–ho–Zn alloy. Mater Lett 2014;133:158–162. Available

from: https://www.sciencedirect.com/science/article/pii/S0167577X

1401235X.

140. Zhang X, Wu Y, Xue Y, Wang Z, Yang L. Biocorrosion behavior

and cytotoxicity of a Mg–Gd–Zn–Zr alloy with long period stacking

ordered structure. Mater Lett 2012;86:42–45.

141. Zhang X, Ba Z, Wang Z, Xue Y. Microstructures and corrosion

behavior of biodegradable Mg–6Gd–xZn–0.4Zr alloys with and

without long period stacking ordered structure. Corros Sci 2016;

105:68–77.

142. Zhang X, Wang Q, Chen F, Wu Y, Wang Z, Wang Q. Relation

between LPSO structure and biocorrosion behavior of biodegrad-

able GZ51K alloy. Mater Lett 2015;138:212–215.

143. Willbold E, Gu X, Albert D, Kalla K, Bobe K, Brauneis M,

Janning C, Nellesen J, Czayka W, Tillmann W, Zheng Y, Witte F.

Effect of the addition of low rare earth elements (lanthanum, neo-

dymium, cerium) on the biodegradation and biocompatibility of

magnesium. Acta Biomater 2015;11:554–562.

144. Zhao N, Watson N, Xu Z, Chen Y, Waterman J, Sankar J, Zhu D. In

vitro biocompatibility and endothelialization of novel magnesium-

rare earth alloys for improved stent applications. PLoS One 2014;

9:e98674.

145. Gu X, Zheng Y, Zhong S, Xi T, Wang J, Wang W. Corrosion of,

and cellular responses to Mg–Zn–Ca bulk metallic glasses. Bioma-

terials 2010;31:1093–1103.

146. Datta MK, Chou D-T, Hong D, Saha P, Chung SJ, Lee B,

Sirinterlikci A, Ramanathan M, Roy A, Kumta PN. Structure and

thermal stability of biodegradable Mg–Zn–Ca based amorphous

alloys synthesized by mechanical alloying. Mater Sci Eng B 2011;

176:1637–1643.

147. Wessels V, Le Mené G, Fischerauer SF, Kraus T, Weinberg A-M,

Uggowitzer PJ, Löffler JF. In vivo performance and structural

relaxation of biodegradable bone implants made from MgZnCa

bulk metallic glasses. Adv Eng Mater 2012;14:B357–B364.

148. Ramya M, Sarwat SG, Udhayabanu V, Subramanian S, Raj B,

Ravi KR. Role of partially amorphous structure and alloying ele-

ments on the corrosion behavior of Mg–Zn–Ca bulk metallic glass

for biomedical applications. Mater Des 2015;86:829–835.

149. Wang J, Huang S, Li Y, Wei Y, Xi X, Cai K. Microstructure,

mechanical and bio-corrosion properties of Mn-doped Mg–Zn–Ca

bulk metallic glass composites. Mater Sci Eng C 2013;33:

3832–3838.

150. Wang J, Li Y, Huang S, Wei Y, Xi X, Cai K, Pan F. Effects of Y on

the microstructure, mechanical and bio-corrosion properties of

Mg–Zn–Ca bulk metallic glass. J Mater Sci Technol 2014;30:

1255–1261.

151. Li H, Pang S, Liu Y, Liaw PK, Zhang T. In vitro investigation of Mg–

Zn–Ca–Ag bulk metallic glasses for biomedical applications. J Non

Cryst Solids 2015;427:134–138.

152. Dambatta MS, Izman S, Yahaya B, Lim JY, Kurniawan D. Mg-

based bulk metallic glasses for biodegradable implant materials:

A review on glass forming ability, mechanical properties, and bio-

compatibility. J Non Cryst Solids 2015;426:110–115.

153. Wilcox GD, Walker DE. Chemical Conversion Coatings. Boston,

MA: Encycl Tribol; 2013.

154. Chen XB, Easton MA, Birbilis N, Yang HY, Abbott TB. Corrosion-

resistant coatings for magnesium (Mg) alloys. Corros Prev

Magnes Alloy 2013;282–312.

155. Rettig R, Virtanen S. Composition of corrosion layers on a magne-

sium rare-earth alloy in simulated body fluids. J Biomed Mater

Res Part A 2009;88A:359–369.

156. Lorenz C, Brunner JG, Kollmannsberger P, Jaafar L, Fabry B,

Virtanen S. Effect of surface pre-treatments on biocompatibility of

magnesium. Acta Biomater 2009;5:2783–2789.

157. Gray Munro JE, Strong M. The mechanism of deposition of cal-

cium phosphate coatings from solution onto magnesium alloy

AZ31. J Biomed Mater Res Part A 2009;90A:339–350.

158. Xu L, Pan F, Yu G, Yang L, Zhang E, Yang K. In vitro and in vivo

evaluation of the surface bioactivity of a calcium phosphate

coated magnesium alloy. Biomaterials 2009;30:1512–1523.

159. Hiromoto S, Yamamoto A. High corrosion resistance of magne-

sium coated with hydroxyapatite directly synthesized in an aque-

ous solution. Electrochim Acta 2009;54:7085–7093.

160. Chen X-B, Birbilis N, Abbott TB. A simple route towards a

hydroxyapatite–Mg(OH)2 conversion coating for magnesium. Cor-

ros Sci 2011;53:2263–2268.

161. Yang J, Cui F, Lee I-S, Zhang Y, Yin Q, Xia H, Yang S. In vivo bio-

compatibility and degradation behavior of Mg alloy coated by cal-

cium phosphate in a rabbit model. J Biomater Appl 2012;27:

153–164.

162. Hiromoto S, Inoue M, Taguchi T, Yamane M, Ohtsu N. In vitro and

in vivo biocompatibility and corrosion behaviour of a bioabsorb-

able magnesium alloy coated with octacalcium phosphate and

hydroxyapatite. Acta Biomater 2015;11:520–530.

163. Chen XB, Nisbet DR, Li RW, Smith PN, Abbott TB, Easton MA,

Zhang D-H, Birbilis N. Controlling initial biodegradation of magne-

sium by a biocompatible strontium phosphate conversion coating.

Acta Biomater 2014;10:1463–1474.

164. Zeng R, Hu Y, Zhang F, Huang Y, Wang Z, Li S, Han E. Corrosion

resistance of cerium-doped zinc calcium phosphate chemical con-

version coatings on AZ31 magnesium alloy. Trans Nonferrous

Met Soc Chin 2016;26:472–483.


https://www.sciencedirect.com/science/article/pii/S1359646202005900




https://www.sciencedirect.com/science/article/pii/S0167577X1401235X

https://www.sciencedirect.com/science/article/pii/S0167577X1401235X

165. Jayaraj J, Amruth Raj S, Srinivasan A, Ananthakumar S,

Pillai UTS, Dhaipule NGK, Mudali UK. Composite magnesium

phosphate coatings for improved corrosion resistance of magne-

sium AZ31 alloy. Corros Sci 2016;113:104–115.

166. Van Phuong N, Gupta M, Moon S. Enhanced corrosion perfor-

mance of magnesium phosphate conversion coating on AZ31

magnesium alloy. Trans Nonferrous Met Soc Chin 2017;27:

1087–1095.

167. Hariprasad S, Gowtham S, Arun S, Ashok M, Rameshbabu N. Fab-

rication of duplex coatings on biodegradable AZ31 magnesium

alloy by integrating cerium conversion (CC) and plasma electro-

lytic oxidation (PEO) processes. J Alloys Compd 2017;722:

698–715.

168. Wang Z, Guo Y. Corrosion resistance and adhesion of poly(L-lactic

acid)/MgF2 composite coating on AZ31 magnesium alloy for bio-

medical application. Russ J Non-Ferrous Met 2016;57:381–388.

169. Carboneras M, García Alonso MC, Escudero ML. Biodegradation

kinetics of modified magnesium-based materials in cell culture

medium. Corros Sci 2011;53:1433–1439.

170. Chiu KY, Wong MH, Cheng FT, Man HC. Characterization and cor-

rosion studies of fluoride conversion coating on degradable Mg

implants. Surf Coat Technol 2007;202:590–598.

171. Lin DJ, Hung FY, Yeh ML, Lee HP, Lui TS. Development of a novel

micro-textured surface using duplex surface modification for bio-

medical Mg alloy applications. Mater Lett 2017;206:9–12.

172. Seitz J-M, Collier K, Wulf E, Bormann D, Bach F-W. Comparison of

the corrosion behavior of coated and uncoated magnesium alloys

in an in vitro corrosion environment. Adv Eng Mater 2011;13:

B313–B323.

173. Yan T, Tan L, Xiong D, Liu X, Zhang B, Yang K. Fluoride treatment

and in vitro corrosion behavior of an AZ31B magnesium alloy.

Mater Sci Eng C 2010;30:740–748.

174. Jiang H, Wang J, Chen M, Liu D. Biological activity evaluation of

magnesium fluoride coated Mg-Zn-Zr alloy in vivo. Mater Sci Eng

C 2017;75:1068–1074.

175. Li Z, Shizhao S, Chen M, Fahlman BD, Liu D, Bi H. In vitro and

in vivo corrosion, mechanical properties and biocompatibility

evaluation of MgF 2 -coated Mg-Zn-Zr alloy as cancellous screws.

Mater Sci Eng C 2017;75:1268–1280.

176. Kang MH, Jang TS, Kim SW, Park HS, Song J, Kim HE, Jung KH,

Do JH. MgF2-coated porous magnesium/alumina scaffolds with

improved strength, corrosion resistance, and biological perfor-

mance for biomedical applications. Mater Sci Eng C 2016;62:

634–642.

177. Yan T, Tan L, Zhang B, Yang K. Fluoride conversion coating on

biodegradable AZ31B magnesium alloy. J Mater Sci Technol 2014;

30:666–674.

178. Witte F, Fischer J, Nellesen J, Vogt C, Vogt J, Donath T,

Beckmann F. In vivo corrosion and corrosion protection of magne-

sium alloy LAE442. Acta Biomater 2010;6:1792–1799.

179. Li Q, Zhu P, Chen S, Zhang B, Yang K. In vitro study on degrada-

tion of AZ31B magnesium alloy with fluoride conversion coating.

Mater Technol 2017;32:409–414.

180. Hornberger H, Virtanen S, Boccaccini AR. Biomedical coatings on

magnesium alloys—A review. Acta Biomater 2012;8:2442–2455.

181. Liu X, Zhen Z, Liu J, Xi T, Zheng Y, Guan S, Zheng Y, Cheng Y.

Multifunctional MgF2/polydopamine coating on Mg alloy for vas-

cular stent application. J Mater Sci Technol 2015;31:733–743.

182. Ren M, Cai S, Liu T, Huang K, Wang X, Zhao H, Niu S, Zhang R,

Wu X. Calcium phosphate glass/MgF2 double layered composite

coating for improving the corrosion resistance of magnesium

alloy. J Alloys Compd 2014;591:34–40.

183. Jiang ST, Zhang J, Shun SZ, Chen MF. The formation of FHA coat-

ing on biodegradable Mg-Zn-Zr alloy using a two-step chemical

treatment method. Appl Surf Sci 2016;388:424–430.

184. Zhang XP, Zhao ZP, Wu FM, Wang YL, Wu J. Corrosion and wear

resistance of AZ91D magnesium alloy with and without microarc

oxidation coating in Hank’s solution. J Mater Sci 2007;42:

8523–8528.

185. Xu X, Lu P, Guo M, Fang M. Cross-linked gelatin/nanoparticles

composite coating on micro-arc oxidation film for corrosion and

drug release. Appl Surf Sci 2010;256:2367–2371.

186. Wilke BM, Zhang L, Li W, Ning C, Chen C, Gu Y. Corrosion perfor-

mance of MAO coatings on AZ31 Mg alloy in simulated body fluid

vs. Earle’s balance salt solution. Appl Surf Sci 2016;363:328–337.

187. Yu H, Dong Q, Dou J, Pan Y, Chen C. Structure and in vitro bioac-

tivity of ceramic coatings on magnesium alloys by microarc oxida-

tion. Appl Surf Sci 2016;388:114–119.

188. Cui LY, Zeng RC, Guan SK, Qi WC, Zhang F, Li SQ, Han EH. Degra-

dation mechanism of micro-arc oxidation coatings on biodegrad-

able Mg-Ca alloys: The influence of porosity. J Alloys Compd

2017;695:2464–2476.

189. Uddin MS, Hall C, Murphy P. Surface treatments for controlling

corrosion rate of biodegradable Mg and Mg-based alloy implants.

Sci Technol Adv Mater 2015;16:053501.

190. Tang H, Han Y, Wu T, Tao W, Jian X, Wu Y, Xu F. Synthesis and

properties of hydroxyapatite-containing coating on AZ31 magne-

sium alloy by micro-arc oxidation. Appl Surf Sci 2017;400:

391–404.

191. Tang H, Wu T, Wang H, Jian X, Wu Y. Corrosion behavior of HA

containing ceramic coated magnesium alloy in Hank’s solution.

J Alloys Compd 2017;698:643–653.

192. Tang H, Gao Y. Preparation and characterization of hydroxyapatite

containing coating on AZ31 magnesium alloy by micro-arc oxida-

tion. J Alloys Compd 2016;688:699–708.

193. Ma X, Zhu S, Wang L, Ji C, Ren C, Guan S. Synthesis and proper-

ties of a bio-composite coating formed on magnesium alloy by

one-step method of micro-arc oxidation. J Alloys Compd 2014;

590:247–253.

194. Wu C, Wen Z, Dai C, Lu Y, Yang F. Fabrication of calcium phospha-

te/chitosan coatings on AZ91D magnesium alloy with a novel

method. Surf Coat Technol 2010;204:3336–3347.

195. Yao Z, Li L, Jiang Z. Adjustment of the ratio of Ca/P in the ceramic

coating on Mg alloy by plasma electrolytic oxidation. Appl Surf

Sci 2009;255:6724–6728.

196. Kim SY, Kim YK, Ryu MH, Bae TS, Lee MH. Corrosion resistance

and bioactivity enhancement of MAO coated Mg alloy depending

on the time of hydrothermal treatment in Ca-EDTA solution. Sci

Rep 2017;7:9061.

197. Bakhsheshi-Rad HR, Hamzah E, Ebrahimi-Kahrizsangi R,

Daroonparvar M, Medraj M. Fabrication and characterization of

hydrophobic microarc oxidation/poly-lactic acid duplex coating on

biodegradable Mg–Ca alloy for corrosion protection. Vacuum

2016;125:185–188.

198. Zeng RC, Cui LY, Jiang K, Liu R, Zhao BD, Zheng YF. In vitro corro-

sion and cytocompatibility of a microarc oxidation coating and

poly(L-lactic acid) composite coating on Mg−1Li−1Ca alloy for

orthopedic implants. ACS Appl Mater Interfaces 2016;8:

10014–10028.

199. Tian P, Liu X, Ding C. In vitro degradation behavior and cytocom-

patibility of biodegradable AZ31 alloy with PEO/HT composite

coating. Colloids Surf B Biointerfaces 2015;128:44–54.

200. Wei Z, Tian P, Liu X, Zhou B. Hemocompatibility and selective cell

fate of polydopamine-assisted heparinized PEO/PLLA composite

coating on biodegradable AZ31 alloy. Colloids Surf B Biointerfaces

2014;121:451–460.

201. Castellanos A, Altube A, Vega JM, García-Lecina E, Díez JA,

Grande HJ. Effect of different post-treatments on the corrosion

resistance and tribological properties of AZ91D magnesium alloy

coated PEO. Surf Coat Technol 2015;278:99–107.

202. Gao JH, Shi XY, Yang B, Hou SS, Meng EC, Guan FX, Guan SK.

Fabrication and characterization of bioactive composite coatings

on Mg–Zn–Ca alloy by MAO/sol–gel. J Mater Sci Mater Med 2011;

22:1681–1687.

203. Shi P, Ng WF, Wong MH, Cheng FT. Improvement of corrosion

resistance of pure magnesium in Hanks’ solution by microarc oxi-

dation with sol–gel TiO2 sealing. J Alloys Compd 2009;469:

286–292.

204. Jiang HB, Wu G, Lee S-B, Kim K-M. Achieving controllable degra-

dation of a biomedical magnesium alloy by anodizing in molten

ammonium bifluoride. Surf Coat Technol 2017;313:282–287.

205. Forero López AD, Lehr IL, Saidman SB. Anodisation of AZ91D

magnesium alloy in molybdate solution for corrosion protection.

J Alloys Compd 2017;702:338–345.


REVIEW ARTICLE

206. Anawati A, Asoh H, Ono S. Effects of alloying element Ca on the

corrosion behavior and bioactivity of anodic films formed on

AM60 Mg alloys. Materials 2016;10:11.

207. Mousa HM, Hussein KH, Woo HM, Park CH, Kim CS. One-step

anodization deposition of anticorrosive bioceramic compounds on

AZ31B magnesium alloy for biomedical application. Ceram Int

2015;41:10861–10870.

208. Mousa HM, Lee DH, Park CH, Kim CS. A novel simple strategy for

in situ deposition of apatite layer on AZ31B magnesium alloy for

bone tissue regeneration. Appl Surf Sci 2015;351:55–65.

209. Mousa HM, Hussein KH, Pant HR, Woo HM, Park CH, Kim CS. In

vitro degradation behavior and cytocompatibility of a bioceramic

anodization films on the biodegradable magnesium alloy. Colloids

Surf A Physicochem Eng Asp 2016;488:82–92.

210. Turhan MC, Lynch RP, Jha H, Schmuki P, Virtanen S. Anodic

growth of self-ordered magnesium oxy-fluoride nanoporous/tubu-

lar layers on Mg alloy (WE43). Electrochem Commun 2010;12:

796–799.

211. Barchiche CE, Rocca E, Juers C, Hazan J, Steinmetz J. Corrosion

resistance of plasma-anodized AZ91D magnesium alloy by electro-

chemical methods. Electrochim Acta 2007;53:417–425.

212. Shi Z, Song G, Atrens A. Corrosion resistance of anodised single-

phase Mg alloys. Surf Coat Technol 2006;201:492–503.

213. Blawert C, Dietzel W, Ghali E, Song G. Anodizing treatments for

magnesium alloys and their effect on corrosion resistance in vari-

ous environments. Adv Eng Mater 2006;8:511–533.

214. Liu J, Zheng Y, Bi Y, Li Y, Zheng Y. Improved cytocompatibility of

Mg-1Ca alloy modified by Zn ion implantation and deposition.

Mater Lett 2017;205:87–89.

215. Wu H, Wu G, Chu PK. Effects of cerium ion implantation on the

corrosion behavior of magnesium in different biological media.

Surf Coat Technol 2016;306:6–10.

216. Zhao Y, Wu G, Lu Q, Wu J, Xu R, Yeung KWK, Chu PK. Improved

surface corrosion resistance of WE43 magnesium alloy by dual

titanium and oxygen ion implantation. Thin Solid Films 2013;529:

407–411.

217. Wu G, Ding K, Zeng X, Wang X, Yao S. Improving corrosion resis-

tance of titanium-coated magnesium alloy by modifying surface

characteristics of magnesium alloy prior to titanium coating depo-

sition. Scr Mater 2009;61:269–272.

218. Wan YZ, Xiong GY, Luo HL, He F, Huang Y, Wang YL. Influence of

zinc ion implantation on surface nanomechanical performance

and corrosion resistance of biomedical magnesium–calcium

alloys. Appl Surf Sci 2008;254:5514–5516.

219. Wang X, Zeng X, Wu G, Yao S, Lai Y. Effects of tantalum ion

implantation on the corrosion behavior of AZ31 magnesium

alloys. J Alloys Compd 2007;437:87–92.

220. Cheng M, Qiao Y, Wang Q, Qin H, Zhang X, Liu X. Dual ions

implantation of zirconium and nitrogen into magnesium alloys for

enhanced corrosion resistance, antimicrobial activity and biocom-

patibility. Colloids Surf B Biointerfaces 2016;148:200–210.

221. Jin W, Wu G, Gao A, Feng H, Peng X, Chu PK. Hafnium-implanted

WE43 magnesium alloy for enhanced corrosion protection and

biocompatibility. Surf Coat Technol 2016;306:11–15.

222. Jin W, Wu G, Feng H, Wang W, Zhang X, Chu PK. Improvement of

corrosion resistance and biocompatibility of rare-earth WE43 mag-

nesium alloy by neodymium self-ion implantation. Corros Sci

2015;94:142–155.

223. Zheng Y, Li Y, Chen J, Zou Z. Surface characteristics and corrosion

resistance of biodegradable magnesium alloy ZK60 modified by

Fe ion implantation and deposition. Prog Nat Sci Mater Int 2014;

24:547–553.

224. Xu R, Yang X, Li P, Suen KW, Wu G, Chu PK. Eelectrochemical

properties and corrosion resistance of carbon-ion-implanted mag-

nesium. Corros Sci 2014;82:173–179.

225. Jamesh MI, Wu G, Zhao Y, Jin W, McKenzie DR, Bilek MMM,

Chu PK. Effects of zirconium and nitrogen plasma immersion ion

implantation on the electrochemical corrosion behavior of Mg–Y–

RE alloy in simulated body fluid and cell culture medium. Corros

Sci 2014;86:239–251.

226. Xu R, Yang X, Jiang J, Li P, Wu G, Chu PK. Effects of chromium

ion implantation voltage on the corrosion resistance and

cytocompatibility of dual chromium and oxygen plasma-ion-

implanted biodegradable magnesium. Surf Coat Technol 2013;

235:875–880.

227. Jamesh MI, Wu G, Zhao Y, McKenzie DR, Bilek MMM, Chu PK.

Effects of zirconium and oxygen plasma ion implantation on the

corrosion behavior of ZK60 Mg alloy in simulated body fluids. Cor-

ros Sci 2014;82:7–26.

228. Liu Q, Kang Z. One-step electrodeposition process to fabricate

superhydrophobic surface with improved anticorrosion property

on magnesium alloy. Mater Lett 2014;137:210–213.

229. Monasterio N, Ledesma JL, Aranguiz I, Garcia-Romero A, Zuza E.

Analysis of electrodeposition processes to obtain calcium phos-

phate layer on AZ31 alloy. Surf Coat Technol 2017;319:12–22.

230. Shangguan Y, Wan P, Tan L, Fan X, Qin L, Yang K. Investigation of

the inner corrosion layer formed in pulse electrodeposition coat-

ing on Mg-Sr alloy and corresponding degradation behavior.

J Colloid Interface Sci 2016;481:1–12.

231. Seyedraoufi ZS, Mirdamadi S. In vitro biodegradability and bio-

compatibility of porous Mg-Zn scaffolds coated with nano

hydroxyapatite via pulse electrodeposition. Trans Nonferrous Met

Soc Chin 2015;25:4018–4027.

232. Manoj Kumar R, Kuntal KK, Singh S, Gupta P, Bhushan B,

Gopinath P, Lahiri D. Electrophoretic deposition of hydroxyapatite

coating on Mg–3Zn alloy for orthopaedic application. Surf Coat

Technol 2016;287:82–92.

233. Jia L, Liang C, Huang N, Zhou Z, Duan F, Wang L. Morphology

and composition of coatings based on hydroxyapatite-chitosan-

RuCl3 system on AZ91D prepared by pulsed electrochemical

deposition. J Alloys Compd 2016;656:961–971.

234. Kannan MB, Orr L. In vitro mechanical integrity of hydroxyapatite

coated magnesium alloy. Biomed Mater 2011;6:045003.

235. Wen C, Guan S, Peng L, Ren C, Wang X, Hu Z. Characterization

and degradation behavior of AZ31 alloy surface modified by bone-

like hydroxyapatite for implant applications. Appl Surf Sci 2009;

255:6433–6438.

236. Gu XN, Guo HM, Wang F, Lu Y, Lin WT, Li J, Zheng YF, Fan YB.

Degradation, hemolysis, and cytotoxicity of silane coatings on bio-

degradable magnesium alloy. Mater Lett 2017;193:266–269.

237. Cui LY, Qin P-H, Huang X-L, Yin Z-Z, Zeng R-C, Li S-Q, Han E-H,

Wang Z-L. Electrodeposition of TiO2 layer-by-layer assembled

composite coating and silane treatment on Mg alloy for corrosion

resistance. Surf Coat Technol 2017;324:560–568.

238. Liu Y, Xue J, Luo D, Wang H, Gong X, Han Z, Ren L. One-step fab-

rication of biomimetic superhydrophobic surface by electrodeposi-

tion on magnesium alloy and its corrosion inhibition. J Colloid

Interface Sci 2017;491:313–320.

239. Han B, Yang Y, Fang L, Peng G, Yang C. Electrodeposition of super-

hydrophobic nickel film on magnesium alloy AZ31 and its corrosion

resistance. Int J Electrochem Sci 2016;111123:9206–9215.

240. Kang Z, Li W. Facile and fast fabrication of superhydrophobic sur-

face on magnesium alloy by one-step electrodeposition method.

J Ind Eng Chem 2017;50:50–56.

241. Zhu X, Zhu Z, Chen C, Bing N, Xu Z, Li Y, Chen Q, Zhu X, Zhu Z,

Chen C, Bing N, Xu Z, Li Y, Chen Q. Surface properties contrast

between Al films and TiO2 films coated on magnesium alloys by

magnetron sputtering. Mater Res 2017;20:481–486.

242. Mukhametkaliyev TM, Surmeneva MA, Vladescu A, Cotrut CM,

Braic M, Dinu M, Vranceanu MD, Pana I, Mueller M,

Surmenev RA. A biodegradable AZ91 magnesium alloy coated

with a thin nanostructured hydroxyapatite for improving the corro-

sion resistance. Mater Sci Eng C 2017;75:95–103.

243. Kiahosseini SR, Afshar A, Mojtahedzadeh Larijani M, Yousefpour M.

Structural and corrosion characterization of hydroxyapatite/zirco-

nium nitride-coated AZ91 magnesium alloy by ion beam sputtering.

Appl Surf Sci 2017;401:172–180.

244. Melnikov ES, Surmeneva MA, Tyurin AI, Pirozhkova TS,

Shuvarin IA, Prymak O, Epple M, Surmenev RA. Improvement of

the mechanical properties of AZ91D magnesium alloys by deposi-

tion of thin hydroxyapatite film. Nano Hybrids Compos 2017;13:

355–361.

245. Dinu M, Ivanova AA, Surmeneva MA, Braic M, Tyurin AI, Braic V,

Surmenev RA, Vladescu A. Tribological behaviour of RF-


magnetron sputter deposited hydroxyapatite coatings in physio-

logical solution. Ceram Int 2017;43:6858–6867.

246. Surmeneva MA, Mukhametkaliyev TM, Khakbaz H, Surmenev RA,

Bobby Kannan M. Ultrathin film coating of hydroxyapatite (HA) on

a magnesium–calcium alloy using RF magnetron sputtering for

bioimplant applications. Mater Lett 2015;152:280–282.

247. Surmeneva MA, Surmenev RA. Microstructure characterization

and corrosion behaviour of a nano-hydroxyapatite coating depos-

ited on AZ31 magnesium alloy using radio frequency magnetron

sputtering. Vacuum 2015;117:60–62.

248. Surmeneva MA, Tyurin AI, Mukhametkaliyev TM, Pirozhkova TS,

Shuvarin IA, Syrtanov MS, Surmenev RA. Enhancement of the

mechanical properties of AZ31 magnesium alloy via nanostructured

hydroxyapatite thin films fabricated via radio-frequency magnetron

sputtering. J Mech Behav Biomed Mater 2015;46:127–136.

249. Jin W,Wang G, Lin Z, Feng H, Li W, Peng X, Qasim AM, Chu PK. Cor-

rosion resistance and cytocompatibility of tantalum-surface-

functionalized biomedical ZK60Mgalloy. Corros Sci 2017;114:45–56.

250. Luo Q, Liu X, Li Z, Huang C, Zhang W, Meng J, Chang Z, Hua Z.

Degradation model of bioabsorbable cardiovascular stents. PLoS

One 2014;9:e110278.

251. Debusschere N, Segers P, Dubruel P, Verhegghe B, De Beule M. A

finite element strategy to investigate the free expansion behaviour

of a biodegradable polymeric stent. J Biomech 2015;48:

2012–2018.

252. Liu Y, Zhang P. Characterization of compression behaviors of fully

covered biodegradable polydioxanone biliary stent for human

body: A numerical approach by finite element model. J Mech

Behav Biomed Mater 2016;62:128–138.

253. Barros AA, Oliveira C, Lima E, Duarte ARC, Reis RL. Gelatin-based

biodegradable ureteral stents with enhanced mechanical proper-

ties. Appl Mater Today 2016;5:9–18.

254. Frischkorn J, Reese S. Solid-beam finite element analysis of nitinol

stents. Comput Methods Appl Mech Eng 2015;291:42–63.

255. Auricchio F, Conti M, Ferrara A, Morganti S, Reali A. Patient-

specific finite element analysis of carotid artery stenting: A focus

on vessel modeling. Int J Numer Method Biomed Eng 2013;29:

645–664.

256. Arokiaraj MC, De Santis G, De Beule M, Palacios IF, Nicola V,

Pak HA. Novel Tram stent method in the treatment of coronary

bifurcation lesions—Finite element study. PLoS One 2016;11:1–15.

257. Gökgöl C, Diehm N, Nezami FR, Büchler P. Nitinol stent oversizing

in patients undergoing popliteal artery revascularization: A finite

element study. Ann Biomed Eng 2015;43:2868–2880.

258. Auricchio F, Constantinescu A, Conti M, Scalet G. A computational

approach for the lifetime prediction of cardiovascular balloon-

expandable stents. Int J Fatigue 2015;75:69–79.

259. Eshghi N, Hojjati MH, Imani M, Goudarzi AM. Finite element analy-

sis of mechanical behaviors of coronary stent. Procedia Eng 2011;

10:3056–3061.

260. Huang L, Pu C, Fisher RK, Mountain DJH, Gao Y, Liaw PK,

Zhang W, He W. A Zr-based bulk metallic glass for future stent

applications: Materials properties, finite element modeling, and

in vitro human vascular cell response. Acta Biomater 2015;25:

356–368.

261. Grogan JA, Leen SB, McHugh PE. Optimizing the design of a

bioabsorbable metal stent using computer simulation methods.

Biomaterials 2013;34:8049–8060.

262. Wu W, Petrini L, Gastaldi D, Villa T, Vedani M, Lesma E,

Previtali B, Migliavacca F. Finite element shape optimization for

biodegradable magnesium alloy stents. Ann Biomed Eng 2010;38:

2829–2840.

263. Wang J, Giridharan V, Shanov V, Xu Z, Collins B, White L, Jang Y,

Sankar J, Huang N, Yun Y. Flow-induced corrosion behavior of

absorbable magnesium-based stents. Acta Biomater 2014;10:

5213–5223.

264. Wei-Chao Z, Shuang-Shou L, Bin T, Da-Ben Z. Microstructure and

properties of Mg-Al binary alloys. China Foundry 2006;3:270–274.

265. Zhang W, Li M, Chen Q, Hu W, Zhang W, Xin W. Effects of Sr and

Sn on microstructure and corrosion resistance of Mg–Zr–Ca mag-

nesium alloy for biomedical applications. Mater Des 2012;39:

379–383.


REVIEW ARTICLE

the current trends of mg alloys in biomedical applications ... · biomedical applications,...

Documents