Magnesium based implants for functional bone tissue regeneration -A review

Gvish Uppl ,Amit Thkur,? ,Amit Chuhn ,Sroj Bl

aUniversity Institute of Engineering and Technology,Panjab University,sector-14,Chandigarh,India b Sri Guru Gobind Singh College,Panjab University,sector-26,Chandigarh,India

Abstract Magnesium (Mg) has emerged as one of the third-generation biomaterials for regeneration and support of functional bone tissue.Mg is a better choice over permanent implants such as titanium,stainless steel,cobalt-chrome as magnesium is biodegradable and does not require a second surgery for its removal after bone tissue recovery.It also reduces the risk of stress shielding as its elastic modulus is closer to human bone in comparison to permanent implants and other biodegradable metallic implants based on Iron and Zinc.Most importantly,Mg is osteoconductive thus stimulates new bone formation and possess anti-bacterial properties hence reducing the risk of failure due to infection.Despite its advantages,a major concern with pure Mg is its rapid bio-corrosion in presence of body fluid due to which the mechanical integrity of the implant deteriorates before healing of the tissue is complete.Mechanical properties of Mg-based implants can be enhanced by mechanical processing,alloying,and topology optimization.To reduce the corrosion/degradation rate,Mg has been alloyed with metals,reinforced with ceramics,and surface coatings have been applied so that the degradation rate of Mg-based implant matches with that of healing rate of bone tissue.The present review discusses the effect of alloying elements and reinforcing ceramics on microstructure,mechanical,and corrosion properties of Mg-based orthopedic implants.In addition,the biocompatibility of Mg-based alloys,composites,and coatings applied on Mg implants has been highlighted.Further,different methods of fabricating porous implants have been highlighted as making the implant porous facilitates the growth of new bone tissue through the pores.

Keywords: Magnesium;Biodegradable orthopedic implant;Porous structure;Osteo-conductive;Antibacterial.

1.Introduction

1.1.Why magnesium (Mg) based orthopedic implants ?

High strength and fracture toughness make metallic biomaterials are an excellent choice for orthopedic applications over polymeric or ceramic biomaterials [1].One such metallic biomaterial is magnesium (Mg).Mg-based biomaterials have found applications as orthopedic implants [2].Mg is the fourth most abundant cation in the human body while within the cell,Mg is considered as the second most abundant cation[3].Mg2+ions play a crucial role in the functioning of the human body.Mg2+ions play a crucial role in the synthesis of proteins and nucleic acid,mitochondrial activity,ion channel modulation,stabilizing DNA and RNA,stabilization of plasma membrane and translational processes,and add support for osteogenic activity [4] [5].The amount of Mg contained in the human body per 70 kg body weight is about 21-35 g and about 180-350 mg of this element is required by the human body daily [6].About 50-60% of human bone tissue consists of Mg.Mg ions are considered crucial for the proliferation of bone cells and the formation of new bone[7] as approximately half of physiological Mg is found in bone [8].Mechanical properties of magnesium such as elastic or young’s modulus (E) and density are also very close to the human bone.Owing to these properties Mg is the perfect metal to be used as an orthopedic bio-implant.Table 1 shows the various properties of human bone along with the properties of different implant materials.Since the modulus of elasticity (E) of Mg is close to the bone it reduces the risk of stress shielding which is prominent in case of permanent implants(Titanium(Ti)alloys,stainless steel(SS),and cobaltchrome (CoCr) alloys).As clear from Table 1,a mismatch of elastic modulus between bone and implant which can result in complications like chronic inflammation early loosening of the implant,slow the rate of healing,etc.[9].Another advantage of biodegradable Mg-based implant over permanent implant is that there is no need for second surgery to remove the implant after the healing of bone tissue [10].Moreover,permanent implants are bio-inert,whereas Mg-based implants are bioactive.A bioactive material interacts with psychochemical environment and helps in assisting various functions of the body,whereas a bio-inert material does not participate in any of the metabolic activities going on inside the body.Permanent implants require a second surgery for their removal after their function is over as being in the body for long,they cause tissue inflammator problems.On the other hand,Mgbased implant disappears completely after the healing process is complete.

Mg and its alloys are a better option to reduce the chances of implant fracture during implantation as they are 3-16 times stronger than biopolymers and are more ductile in comparison to bio-ceramics [19].By-product after degradation/corrosion of Mg implant doesn’t harm the internal functioning of the body [15].Excess Mg accumulated in the human body is excreted via kidneys through urine [20].Even though being biodegradable,polymeric implants such as PLA,PGLA,PCL release acids during their intermediate stage of corrosion which is harmful to the human body [21].The mechanical strength of bio-polymer is lower than that of Mg,therefore not suitable for load-bearing applications.Comparing with bio-ceramics (HA and TCP),the compressive strength of TCP is poor compared to Mg while the elastic modulus of HA is higher than bone so Mg becomes an excellent choice as a degradable implant as the elastic modulus of Mg is in the range of cortical bone (Table 1).In comparison to other degradable metallic elements,pure zinc has an elastic modulus in the range 94-110 GPa [22] and whereas the value of elastic modulus of pure iron is 200 GPa [23].Taking advantage of the point that elastic modulus of Mg falls in the range of human bone,Mg is used as a temporary implant such as screws,rods,pins,etc.[24].It has been reported that Mg2+ions facilitate the formation of new bone and also reduce the time of fracture healing [25].Fabrication of complicated shaped Mg orthopedic implant is relatively easy as it has good machinability,cast-ability,weld-ability,and formability at elevated temperatures [26].Clinical experiments with Mg alloys show that they are biocompatible with no systemic inflammatory reaction [27].Three Mg-based orthopedic screws have received approval for medical use.The three Mg-based orthopedic screws are pure Mg,MAGNEZIX,and K-MET screws.Biodegradable pure Mg screws have reported a remarkable recovery when used for the fixation of vascularized bone grafts to treat osteonecrosis of the femoral head.The Harris hip score (HHS) which indicates the recovery of bone tissue was 95.7% for the Mg screw group and 84% for the control group after 12 months postoperative [28].Another study by Chenet al.supports similar results about improvement in HHS withthe use of pure Mg screws to treat trauma-induced femoral head necrosis [29].MAGNEZIX,a biodegradable compression screw (CS) was designed by Syntellix.This CS had the composition of Mg＞94% with small amounts of yttrium (Y)and rare earth elements[30].MAGNEZIX and Ti screws were used in clinical trials to treat hallux valgus.Both the screws were compared in terms of various clinical parameters postoperatively.It was concluded that MAGNEZIX screws were clinically equivalent to Ti screws as observed from the followup of patients [31,32,33,34,35].In a study,MAGNEZIX CS and Ti screws were used for treating osteochondral lesions of the talus (OLT) and their clinical and radiological outcomes were compared.MAGNEZIX screw seemed beneficial as no removal was required while one patient with Ti screw required implant removal complaining of pain and skin irritation [36].In another clinical trial MAGNEZIX and Ti screws were used for fixin medial malleolar fracture (MMF).Five patients in Ti screw group required removal of the screw due to complaint of pain in three patients and difficult in wearing shoes in other two while no implant removal case was reported in case of MAGNEZIX group [37].The use of Mgbased screws for mandibular condyle fractures (MCF) can be beneficial as compared to conventional Ti screws because a major complication that occurs during implant removal is facial nerve damage/injury.Moreover,during the remodeling process of recovering/healing mandibular condyle Ti screw could penetrate through the condyle surface and damage the temporomandibular joint (TMJ) [38].Another benefit of Mgbased implants over Ti-based implants is that Mg-based implants produce fewer artifacts while examining the implant by imaging techniques such as digital radiography (DX),multidetector computed tomography (MDCT),high resolution fla panel CT (FPCT) and magnetic resonance imaging (MRI)which aids in post-operative follow up [39].Another Mgbased screw is K-MET which is used for fixation of broken bones [40].K-MET screw is an alloy of Mg,Calcium (Ca)and Zinc (Zn) with composition as Mg5wt.%Ca1wt.%Zn.

Table 1Properties of Human Bone and various implant materials.

Another worthy influence of Mg that has been highlighted in literature is the remarkable osseointegration of Mg with enhanced bone-implant interface strength as compared to conventional titanium alloys [41].Mg-based biomaterials have not only been investigated as orthopedic implants but have also been explored to be used as cardiovascular stents,urological stents,and intestinal anastomosis rings.MgxZnySr alloys show acceptable cytocompatibility with human umbilical vein endothelial cells (HUVEC) and human urothelial cells(HUC),therefore these alloys could be employed in cardiovascular and urological stents [42] [43] [44].In-vitro tests were carried out on MgZnSr alloy to be used as intestinal anastomosis rings.MgZnSr alloy showed suitable biocompatibility and anti-bacterial properties [45].Mg6Zn alloy and Ti alloy were compared for their healing response in the case of the Sprague-Dawley rat’s intestinal tract.Compared to Ti alloy,Mg6Zn showed more formation of mature collagen at the anastomosis site.It was concluded that Mg6Zn alloy demonstrated a better healing response for intestinal anastomosis than Ti alloy [46].Apart from Mg,iron (Fe) and zinc (Zn)are the other two biodegradable implant materials.The drawback of pure Mg is that it corrodes rapidly in physiological fluids Pure Fe degrades very slowly but due to ferromagnetic nature,it hinders MRI.While Zn on the other hand has poor mechanical properties.It is important that the release of ions due to corrosion should be in the range that is acceptable to the body.The daily human body intake requirement of Fe is 40-45 mg [47],while that of Zn is 15 mg [22] and of Mg is 250-400 mg [22].

The review article is divided into two sections.Different methods to enhance the corrosion resistance of Mg-based implants and fabrication of porous implants to increase the biological performance by better osseointegration.Making an implant porous compromises their mechanical and corrosion properties but the implant needs to perform the well in-vivo application.The porous structure of the implant helps in the better growth of cells.Therefore,equal emphasis must be laid upon both.Mg despite being biocompatible with mechanical properties falling in the range close to human bone,face the challenge of being corroded rapidly in body fluid resulting in loss of mechanical integrity at a rate faster than that of the healing rate of bone tissue.

1.2.Rapid corrosion of Mg-based implants in body fluid -a big challenge

An implant,when placed inside the human body interacts with its physiological environment.Body fluid influence the properties of Mg implants,as chemical reactions occur between them.As discussed,Mg degrades rapidly in the presence of body fluid so it is important to understand the mechanism of corrosion of Mg.

Corrosion of Mg and its alloys follow an electrochemical process.In an aqueous environment i.e.in presence of body fluid following reaction occurs.

When the Mg implant comes in contact with the body fluid it gets oxidized thereby releasing Mg cations (Mg2+)and electrons (Eq.(1) and Fig.1(a)).The electrons are thus consumed for the reduction of water-producing hydroxyl ions(OHˉ) (Eq.(2)).At the same time bio-molecules present in body fluid (amino acids,lipids,and proteins) also get adsorbed on the implant surface thereby affecting its corrosion.The Mg(OH)2film is formed over the surface of Mg alloy,which is insoluble and protects it from further corrosion(Eq.(3) and Fig.1(b)).Also,hydrogen (H2) gas is released.As hydroxyl ions (OHˉ) are generated,the pH value of the medium near the Mg surface increases making the environment alkaline.The Mg(OH)2film is stable at a high pH value(＞11.5) but at a lower pH value (＜11.5) this layer tends to cease [48].Generally,after surgery secondary metabolic and resorptive process occurs at the implant-bone interface due to which local pH decreases.So,Mg(OH)2layer starts depleting [10].This layer reacts with the chloride ions present in the aqueous medium (body fluids) thus forming soluble MgCl2leading to pitting corrosion on the surface (Eq.(4)and Fig.1(c)) [48].

Fig.1.Bio corrosion mechanism of Biodegradable metal and body fluid (M represents Mg and n=2) [48].

The calcium and phosphate ions present in body fluid stimulate the formation of the biological apatite layer due to the alkaline environment generated (Fig.1(c)).Since the apatite layer formed is bioactive it will allow cell adhesion and proliferation (Fig.1(c)) [48].The pits formed become sites for crack initiation due to stress concentration [49].The cells thus form tissues over time.As the implantation time proceeds more degradation occurs and the corroded portion gets detached from the bulk implant (Fig.1(d)) [48,50].

Due to corrosion,the mechanical integrity of the implant deteriorates.Implant corrodes away before the complete healing of bone has taken place.Also,it’s evident from the reactions that hydrogen gas evolved during corrosion results in tissue separation [51].Keeping these key problems in mind,it is very important to control the rapid degradation rate of Mg-based implants as it causes hydrogen bubble build-up and a rise in local pH value.These changes lead to cell death and tissue inflammation.

2.Methods to enhance the corrosion resistance of Mg-based orthopedic implants

2.1.Mechanical working and heat treatment

Generally,pure Mg is brittle which may affect the behavior of implant under loading.This can be significantly improved by grain refinement by the selection of an appropriate method of production and the addition of certain alloying elements.Increased formability for magnesium is achievable at room temperature when the grain size is refine to micro and nano-level [52].The reason for this brittle behavior lies in transformation of pyramidal II＜c+a ＞slip dislocations into low-energy 〈c+a〉 and 〈c〉 dislocations.These lowenergy dislocations are immobile because they act as ‘forest’dislocation which restrains the dislocation movement leading to brittleness [52].The stabilization of pyramidal II 〈c+a〉 slip dislocation could be achieved by the addition of rare earth elements like yttrium,zinc,strontium,holmium,etc.More discussion on the effect of alloying elements has been done in Section 2.2.

Although the modulus of Mg and its alloys are close to that of human bone still the mechanical properties need to be improved so that they withstand the load after surgery and maintain their mechanical integrity during corrosion till the new bone tissue is formed.Methods that can improve the mechanical properties of Mg alloys are heat treatment,deformation processing such as rolling,extrusion,and severe plastic deformation techniques such as equal channel angular pressing (ECAP).These processes not only enhance the mechanical properties of Mg-based alloys but also enhance their corrosion resistance [53].As-cast Mg1Ca alloy was subjected to two different deformation processes,extrusion at 210 °C and rolling at 400 °C.From the tensile test results,it was concluded that hot rolled and hot extruded Mg1Ca possessed higher tensile yield strength (TYS),ultimate tensile strength (UTS),and elongation than as-cast alloy due to grain refinement Moreover,the hot rolled and extruded alloys had a lower corrosion rate in simulated body fluid (SBF) compared to as-cast alloy [54].ECAP processed Mg2.9Gd1.5Nd0.3Zn0.3Zr alloy possessed higher hardness(85.03 HV) and TYS (217.3 MPa) compared to as-cast alloy (68.05 HV and 85.8 MPa) as shown in Fig.2(a).The grain size was reduced from 10 to 70 μm for as-cast alloy to 2.5 μm for ECAP alloy.The corrosion rate of this alloy decreased compared to as-cast alloy as observed from immersion and electrochemical tests in SBF [55].

Fig.2.(a) Engineering stress-strain curve of as-cast and ECAPed alloys subjected to tensile loading [55] (b) Engineering stress-strain curve (compressive) of as-cast Mg4Ag alloy subjected to T4 treatment and ECAP [56].

ECAP was used to fabricate Mg97Zn1Y2 alloy by compacting the powder obtained from mechanical milling of the chips of as-cast ingots.Four passes of ECAP were applied at 300 °C.The Mg12Zn1Y1 phase present in the cast sample was converted into the Mg24Y5 phase during milling.After four passes of ECAP,it was found that the alloy had the lowest porosity and fines crystallite size of 41 nm as compared to as-cast alloy whose crystallite size was 59 nm.Also,the hardness and CYS of ECAP processed alloy were higher than as-cast alloy but with reduced ductility.The increase in CYS and hardness was due to solid solution and dispersion strengthening [57].As-cast Mg4Ag alloy was homogenized at 440 °C for 16 hrs followed by water quenching(T4 treatment) and ECAP was carried out at 370 °C for the firs pass and 330 °C for the second pass to evaluate the effect of these treatments on compressive strength.The dendritic Mg54Ag17phase dissolved completely after T4 treatment with no change in the grain size.Lattice defect and recrystallization during ECAP resulted in the reduction of grain size.T4 treatment improved the mechanical properties of ascast samples slightly while samples subjected to ECAP resulted in enhanced CYS and ultimate compressive strength(UCS) (Fig.2(b)).The hydrogen evolution and degradation rate of T4 and ECAPed samples in SBF was lower than ascast samples [56].Mechanical properties of MgSr alloys were evaluated after subjecting them to two heat treatments,homogenization at 450 °C for 12 h and aging at 160 °C for 30-300 h.The alloy samples after each treatment were water quenched.The hardness,tensile and compressive strength of as-cast alloy decreased after homogenization as the fraction of the second phase reduced whereas aged alloy samples possessed the highest hardness,tensile and compressive strength as nano-second phase re-precipitated during aging with no effect on grain size.The corrosion rate of homogenized and aged alloys was much lower as compared to as-cast alloys as observed from the immersion test [58].Effect of extrusion on microstructure and mechanical properties of as-cast Mg4Zn1La alloy was studied by Zenginet al..Dynamic recrystallization was reported during extrusion leading to grain size reduction and second phase particles alignment in the direction of extrusion.As compared to as-cast alloy,there was a grain refinement of 81% in extruded alloy while the elongation,yield strength,and tensile strength of extruded alloy increased by 179%,90%,and 40% respectively [59].Mg1.5Sr alloy was heat-treated to investigate its effect on invitro and in-vivo degradation.The authors showed that heat treatment improved corrosion resistance in-vitro and in-vivo as it reduced the amount of Mg17Sr2phase.Fig.3(a) shows the corrosion rate of untreated and heat-treated Mg1.5Sr alloy in hanks solution for 14 days and Fig.3(b) shows potentiodynamic polarization curves of untreated and heat-treated Mg1.5Sr alloy in hanks solution.Also,bone formation was observed in the case of heat treated implant than un-treated implant while both implants had better bone formation than TCP implant [60].

Fig.3.(a) Immersion and (b) Electrochemical test of untreated and heat-treated Mg1.5Sr alloy [60].

Mg3Zn alloy was subjected alloy to five cycles of rolling at 310 °C to understand its effect on the microstructure and mechanical properties of the alloy.Authors reported that grain size reduced with increasing cycles of rolling increased hardness,tensile yield strength,and ultimate tensile strength [61].Different amounts of rolling deformation (25 45,65,85%)were induced in Mg6Sn3Al1Zn alloy at 400 °C and their effect on microstructure and mechanical properties of the alloy were studied.The volume fraction of recrystallized grains increased linearly with an increase in deformation%.The grain structure was also refine due to recrystallization,which resulted in improved mechanical properties like hardness and tensile strength.It was concluded that rolling deformation of 65% resulted in the best mechanical properties,any higher deformation did not improve properties further [62].

The mechanical properties of porous implants are also influence by surface topology.Topology optimization (TO) is a computational process in which under a given set of loading and boundary conditions,system performance is maximized using finite element modeling [63].TO of implant helps in reducing equivalent stiffness,thereby reducing stress shielding effect which otherwise is a major concern.The optimization is done by adding or removing the volume using the Inverse homogenization theory [64-65].Load transfer is enhanced for topology optimized implants compared to that of non-optimized [66].However,the reduction in surface area of the porous implants may lead to increase stress but as the bone regenerates the surface area in contact will increase,resulting in the reduction of stress.Common discrete element methodologies for TO are solid isotropic material with penalization (SIMP) [67],homogenization method (HM) [68],level-set method(LSM)[69],evolutionary structural optimization (ESO) [70-71] and ground structure approach (GSA).

2.2.Alloying

Alloying with suitable elements and inappropriate concentration has enabled researchers to control the corrosion rate of Mg.Properties of any material such as strength,ductility,corrosion resistance are affected by its composition.For Mg alloys,improvements in strength and corrosion resistance are linked with variations of grain size.Elements such as Aluminium (Al),Zinc,Calcium (Ca),Strontium (Sr),Yttrium (Y),Cerium (Ce),etc.,are used as alloying elements in Mg-based biomaterials.Elements like Zn,Ca,Sr are present in the human body and assist in its metabolic functioning.Moreover,these are considered nutrient elements for the body.While selecting any alloying elements,any harmful effects on the human body should be kept in mind.Al is found to be harmful to osteoblasts and neurons [72].Also,high concentrations of Al in the body lead to phosphate depletion which causes Alzheimer’s disease [73].It has been reported that rare earth elements Pr,Y,Ce can cause hepatotoxicity [74].Y is known to cause liver toxicity [75] whereas Zirconium (Zr) can cause lung and breast cancer [76].Though other elements such as silicon (Si),manganese (Mn),and lithium (Li) have also been used as alloying elements in the Mg matrix and have shown excellent biocompatibility.

2.2.1.Binary Mg-based alloys

2.2.1.1.Mg-Zn alloys.Zn is considered to be a crucial nutrient element in the human body and approximately 85% of this element occurs inside muscles and bones [77].On daily basis,about 15 mg of Zn is required in the diet.The deficiency of Zn in animal models decreased the rate of DNA replication which in turn restrained protein synthesis and retarded bone maturation [78].Zn plays a crucial role in osteoblast mineralization and thereby bone metabolism.It is an inorganic element important for the growth of tissues around bio-implant [79].The addition of Zn as an alloying element increases the corrosion potential of Mg thus improving its corrosion resistance.The hemolysis rate of MgZn alloy was reported to be 3.4%which was less than the safety limit of 5%indicating no harm to red blood cells and supported MC3T3-E1 cell adhesion which proved its in-vitro biocompatibility[80].Further in-vivo biocompatibility of Mg6Zn alloy was evaluated by implanting it in the femoral shaft of rabbits.Histological examination of the rabbit’s heart,kidney,and spleen revealed that the implant showed no adverse effects on these vital organs.Also,the alloy supported the formation of new bone which was evident from hematoxylin and eosin(HE)stained sections surrounding the implant 6 and 18 weeks post-implantation (Fig.4) [81].

The corrosion by-products of porous Mg6Zn scaffolds immersed in hank’s solution for 72 hrs were analyzed by Yanet al.XRD analysis showed that by-products consisted of hydroxyapatite (HA),a component of human bone,which accelerates the healing process.Porous Mg6Zn scaffold possessed higher compressive and fl xural strength compared to porous Mg [82].The presence of HA has also been confirmed in corrosion products of Mg6Zn alloy immersed in SBF [80,81].In-vivo study of MgZn implant was done to study biocompatibility and biodegradability of the implant in New Zealand rabbits.Authors reported subcutaneous gas formation in the initial weeks followed by complete removal of gas within 6 weeks post-surgery.The implant completely disappeared after 24 weeks with the majority being reabsorbed within 14 weeks post-surgery.The implant was biocompatible as the pathological evaluation revealed no abnormalities in the heart,liver,kidney,and spleen extracted post 6 weeks surgery [75].In-vitro and in-vivo trials were conducted on MgZn and PLLA biodegradable materials to evaluate their biocompatibility as orthopedic implants.In-vitro trial with MC3T3-E1 cells,MgZn alloy displayed better cell attachment,mineralization ability,and higher values of osteogenic specific mRNA expressions of COL1α1 and OC than PLLA.In-vivo MgZn degraded faster than PLLA and also showed more bone formation without the presence of any fibrou tissue [83].The amount of the second phases present in the Mg matrix are closely related to the corrosion resistance of Mg alloys.The second phase present in Mg alloys increases the corrosion rate due to the formation of galvanic couples betweenαMg and the second phase.Microstructure,mechanical,and corrosion behavior of MgxZn (x=0,1,5,7 wt.%)alloys were investigated to understand the effect of the addition of different concentrations of Zn in the Mg matrix.Grain size reduced with the addition of Zn.The amount of the second phase (MgZn) in the Mg-Zn alloys increased with an increase in Zn content as revealed by the diffractogram(Fig.5(a)).

Fig.4.MgZn implant supporting new bone formation in animal model observed at (a) 6 weeks (b) 18 weeks post-implantation observed through HE staining(black arrows depicting newly formed trabecular and osteoblasts) (500 μm scale bar) [81].

Fig.5.(a) Diffractogram of pure Mg and MgxZn alloys (x=1,5,7 wt.%) [84],(b) Potentiodynamic Polarization curves of MgxZn (x=2,3,4,5 wt.%)alloys obtained by electrochemical test in 3.5%NaCl [85].

Mg5Zn exhibited the highest tensile strength while Mg7Zn showed the highest compressive strength among MgxZn alloys.The improvement in mechanical and corrosion properties was due to grain refinement and the present second phase (MgZn) [84].Another study on corrosion behavior of MgxZn (x=2,3,4,5 wt.%) alloys was done to investigate the effect of the second phase (MgxZny) on degradation properties.The anodic and cathodic current densities at a point below break down potential decreased in the order Mg5Zn＞Mg4Zn＞Mg3Zn＞Mg2Zn.Mg2Zn alloy with 3.5%NaCl (Fig.5(b)).The immersion test also gave similar results of the corrosion rate of these alloys [85].Porous MgxZn(x=4,6 wt.%) scaffolds were fabricated by Seyedraouf and Mirdamadi,and tested for compressive strength.Authors reported that compressive strength and modulus of Mg6Zn scaffolds were higher than Mg4Zn scaffolds [86].

Compressive strength of porous MgxZn (x=1,3,5 wt.%)alloys was found to increase with an increase in Zn content[87]

2.2.1.2.Mg-Ca alloys.Ca is one of the main elements that occur inside the human bone.Ca has a density of 1.55 g/cm3in the range of human bone and is close to that of Mg (1.74 g/cm3).The microstructure of Mg-Ca alloys consists ofαMg matrix and intermetallic phase Mg2Ca [54].The amount of phase Mg2Ca increases with an increase in Ca content in MgCa alloys [88].The microstructure of MgxCa (x=0,0.7,1,2,3,4 wt.%) alloys was examined to evaluate the effect of Ca addition.Addition of Ca refine the grain structure with a decrease in average grain size [89].In-vitro test of MgxCa (x=0.4,0.8,1.34,5,10,16.2,28 wt.%) alloys was conducted to understand the effect of Ca content on the corrosion behavior of these alloys.From electrochemical tests in hanks balanced salt solution (HBSS),minimum essential medium (MEM),MEM+Fetal Bovine Serum (FBS) author concluded that increasing Ca content increased the degradation rate due to the formation of Mg2Ca phase which has a high rate of reaction.The degradation rate decreased with electrolyte mimicking the physiological environment [90].It was observed that the addition of Ca content beyond 1 wt.%in MgCa alloys deteriorates their corrosion resistance [91].

Fig.6..(a) Bar graph indicating tensile characteristics (YS,UTS,and elongation) of as-cast,rolled,and extruded MgxCa (x=1,2,3 wt.%) alloys [54] (b)Cytotoxicity test using L929 cells.Results of MTT assay presented as% of viability cells in Mg/1Ca,Mg/5Ca,Mg/10Ca composite extraction medium after 1,2,4 days culture [93].

The hardness of MgCa alloys was found to increase with an increase in Ca content due to intermetallic phase Mg2Ca[92][89],.MgxCa (x=1,2,3 wt.%) alloys were fabricated and evaluated for their mechanical and corrosion properties.It was found that mechanical properties (TYS,UTS,and elongation) and corrosion resistance decreased with increasing Ca content.Processing techniques such as hot rolling and extrusion were carried out on Mg1Ca which refine its grain structure and enhanced its mechanical strength (Fig.6(a)) and corrosion resistance.The corrosion rate also decreased from 12.56 mm/yr for as-cast Mg1Ca to 1.63 mm/yr for as-rolled Mg1Ca and 1.74 mm/yr.for as-extruded Mg1Ca.Further extruded Mg1Ca and commercially pure Ti pins were implanted by the authors in rabbits for comparing their bone healing response through radiographic evaluation.Test results indicated new bone formation in the case of Mg1Ca pin due to high osteoblastic activity while Ti pins did not support any bone formation [54].The effect of the addition of varying Ca content in MgxCa (x=0.5,1.25,2.5,5,and 10 wt.%) alloys on the pH of corrosion medium (SBF) was investigated by immersing the alloys in SBF.The pH of SBF increased with increasing Ca concentration in MgCa alloys.[92].A high pH value is known to reduce the viability of cells.Cytotoxicity of Mg/xCa composites (x=1,5,10 wt.%) was tested with L929 cells.Mg/10Ca composite was found to be cytotoxic to the cells which could be due to its rapid degradation rate resulting in a rise in pH value of culture medium (Fig.6(b))[93].

MgxCa (x=0.6,1.2,1.6,2 wt.%) alloys were fabricated to study the effect of varying Ca content on bending and compressive strength by Wanet al.Authors observed that MgCa alloys had higher bending and compressive strength compared to pure Mg,Mg0.6Ca exhibiting highest bending and compressive strength amongst all [88].Mg0.8Ca alloy was implanted into the proximal femur,intramedullary tibial shaft,and thigh muscle of giant rabbits to evaluate the biological response of the alloy.Evaluation of blood samples taken 6 weeks post-surgery revealed that the levels of aspartate aminotransferase (AST),alanine aminotransferase (ALT)in blood were in the normal range indicating no damage to the liver and normal levels of urea and creatinine indicated no damage to kidneys [94].Mg0.8Ca alloy was investigated as an orthopedic implant.The corrosion by-products on the surface of Mg0.8Ca implant after immersion in HBSS for 60 days consisted of carbonated HA similar to the human bonelike apatite [95].The in-vivo degradation and bone formation rate of MgxCa (x=0.5,5 wt.%) implants 2-and 4-weeks post-implantation was assessed.The hydrogen gas pockets formed were less in the case of Mg0.5Ca.The evaluation from Micro-CT data of both implants after the 2nd and 4th week of implantation shows that BV/TV% (bone vol./tissue vol.) for Mg0.5Ca was 1.57% and 34.13% while for Mg5Ca was 0.79% and 4.88% (Fig.7(A),(B)).Also,Mg5Ca degraded completely after 2 weeks while the degradation rate of Mg0.5Ca was 63% and 49% after 2 and 4 weeks respectively(Fig.7(C)) [96].

Fig.7.Analysis of (A) Micro-CT images of Rabbit Femur implanted with MgxCa (x=0.5,5 wt.%) alloys after (a,c) 2 and (b,d) 4 weeks post implantation to evaluate (B) BV/TV (%) and (C) Degradation vol.% [96].

Fig.8.(a)Average corrosion rate of Mg-Sr alloys with varying Sr content after the immersion test in SBF measured in terms of mass loss and H2 evolution[103].(b) Bar graph showing the value of the colony-forming unit (unit/ml) of S.aureus in culture containing different samples [104].

2.2.1.3.Mg-Sr alloys.Another crucial alloying element is Sr.In the human body,almost 99% of the Sr is found in bones[97].It is claimed that Sr triggers bone formation and is known to increase bone mass and reduce the chances of fracture in osteoporotic patients[98].Daily intake of Sr should be between 320 and 400 mg [77].Sr has been found useful for treating osteoporosis due to its anabolic and anti-resorptive properties [99].This element belongs to Group II in the periodic table whose behavior is similar to Ca and can be used as a substitute for Ca in most biomedical compounds such as HA.Sr acts as grain refine when added to Mg matrix as it forms compounds that segregate to grain boundaries.The microstructure of MgSr alloys consists ofαMg and Mg17Sr2intermetallic phase and the amount of this phase increases with an increase in Sr content in the alloy as revealed by XRD analysis [100].It is very important to know the optimum amount of Sr that should be alloyed with Mg to develop an implant with appropriate mechanical and degradation properties.MgxSr (x=0.5,1,1.5 wt.%) alloys were analysed for corrosion behavior.As the Sr concentration increased the corrosion rate increased which was measured through the amount of H2evolved in HBSS [101].In another study electrochemical and immersion tests of MgxSr (x=0.5,1,1.5,2 wt.%)alloys were carried out in HBSS.The authors concluded that the corrosion rate increases with an increase in the Mg17Sr2phase as galvanic couples are formed between Mg17Sr2which acts as cathode andαMg acting as anode [102].The effect of varying Sr on corrosion behavior of MgxSr (x=0.3-2.5 wt.%) alloys was examined.It was found that Sr addition of more than 1 wt.% had a detrimental effect on corrosion properties of MgSr alloys as indicated by mass loss and H2vol.evolution in HBSS (Fig.8(a)) [103].

The corrosion products consisted of HA,and Sr substituted HA in the case of pure Mg and Mg0.5Sr respectively.Sr substituted HA has better bioactivity than HA.In vivo,Mg0.5Sr stent did not show thrombosis 3 weeks post-surgery while thrombosis occurred in the case of WE43 (Mg 93.51wt.%,Y 3.84wt.%,Nd 2 wt.%,Gd 0.5 wt.%,Pr 0.15 wt.%) stent as the elements Mg and Sr were familiar to the body while contents of WE43 were considered foreign substrates by the body [103].One crucial factor that an implant should possess is anti-bacterial property.In a study,it was found that MgxSr(x=0.25,1,1.5,2 wt.%) alloys and pure Mg displayed anti-bacterial nature in comparison to commercially available implants like CaSO4,tri-calcium phosphate (TCP),HA,and 316 L stainless steel (Fig.8(b)).While bending strength decreased with Sr addition but TYS and UTS increased with Sr addition up to 1.5 wt.%.The compressive strength of MgSr alloys was higher than pure Mg except for Mg0.25Sr [104].In-vitro and in-vivo tests were conducted to compare pure Mg and Mg1Sr alloy.Mg1Sr had better in-vitro biocompatibility as it had a lower hemolysis percentage and higher cell viability with a lower degradation rate as compared to pure Mg[105].

The in-vivo osteogenic response was analyzed by examining harvested bone-implant sections (Fig.9(a)).New bone formation percentage (NBF) and bone-to-implant contact percentage (BIC) for Mg1Sr was 48.3% and 15.9% while for pure Mg the values were 1.6% and 0 (Fig.9(b)) [105].The effect of Sr addition on mechanical properties of MgxSr (x=0.5,1,2 wt.%) alloys was evaluated by Shanget al.It was observed that TYS and UTS increased up to 1 wt.%of Sr and then decreased for higher wt.% of Sr.This decrease was attributed to the high percentage of Mg17Sr2phase which acted as crack initiator under tension [58].MgxSr (x=0.2,0.5,1,2 wt.%) alloys were investigated for their degradation and cytocompatibility with bone marrow mesenchymal cells.Mg1Sr showed better corrosion resistance in cell culture media.The cell adhesion density increased with the increase in Sr content but no clear relation between Sr content and BMSC adhesion was reported [106].

2.2.2.Ternary Mg-based alloys

The above section of binary alloys covered the effect of alloying elements Zn,Ca,Sr on mechanical properties,corrosion behavior,and biocompatibility of Mg on new bone formation.The following section discusses ternary alloys and their effect on the behavior of Mg-based implants.

Fig.9.(a) Image showing the histological evaluation of the osteogenic activity of pure Mg and Mg1Sr implant in a rabbit model (calcifie tissue represented by dark pink stain regions,light purple depict bone marrow cells,dark purple indicate nuclei and white arrows show new bone formed) and (b) Results of NBF (%) and BIC (%) of pure Mg and Mg1Sr alloy after 16 weeks of implantation [105].

2.2.2.1.Mg-Zn-Ca alloys.The phases formed in MgZnCa alloys depend on the Zn/Ca atomic ratio,if the ratio is greater than 1.2 thenαMg+Ca2Mg6Zn3phases will be formed and if the ratio is less than 1.2 thenαMg+Mg2Ca+Ca2Mg6Zn3phases will be formed [107][108],.The effect of Zn content on microstructure,mechanical,and corrosion properties of MgxZn0.2Ca alloys (x=0.6,2,2.5 wt.%) was investigated by Liet al.The grain size reduced with the addition of Zn.The UTS and elongation increased till 2 wt.%addition of Zn and decreased with further addition of Zn.Ca2Mg6Zn3phase acted as a strengthening phase whereas the Mg2Ca phase was brittle.The corrosion rate increased with increasing Zn content [107].In a similar study effect of Ca addition on microstructure,mechanical,and corrosion properties of Mg2ZnxCa (x=0,0.2,0.4,0.8 wt.%) alloys were investigated.Addition of Ca refine the grain structure of alloy with Mg2Zn0.2Ca exhibiting the highest UTS and elongation.However,corrosion resistance was reduced with Ca addition due to an increase in the Mg2Ca phase [108].The effect of the phases formed in MgxZnyCa (x=1,2,2.5 wt.%,y=0.6,0.6,1.5 wt.%) alloys on their corrosion behavior was evaluated by conducting electrochemical and immersion tests on these alloys in SBF.Mg2Zn0.6Ca had the lowest current density and highest polarization resistance so displayed excellent anti-corrosion behavior with the formation of HA over its surface compared to others.The corrosion potential of phases formed in MgZnCa alloys followed the order Ca2Mg6Zn3＞αMg＞Mg2Ca [109].The grain size and second phase influence the corrosion behavior of the material.Mg3Zn0.3Ca alloy was heat-treated at different temperatures and for a different period to figure out the effect of grain size and second phase on the corrosion behavior of the alloy.It was reported by the authors that grain size increased with an increase in temperature and time of heat treatment while the volume fraction of second phases presents decreased (Fig.10).The alloy heat-treated at 420 °C for 24 hrs displayed a minimum corrosion rate [110].

MgxZn1Ca (x=1,2,3,4,5,6 wt.%) alloys were investigated for their cytotoxicity with L929 cells.Cytotoxicity test was performed on MgxZn1Ca (x=1,2,3 wt.%) alloys as their pH value was within an acceptable range.The results show that these alloys displayed no cytotoxicity.These alloys supported the formation of HA upon degradation in hank’s solution.The UTS and elongation increased up to 4 wt.% Zn addition and then decreased [6].Two alloys ZX50 and WZ21 were examined for their in vivo degradation rate and interaction with bone tissue.ZX50 corroded faster than WZ21 but both had a positive influence on osteosynthesis.It was found that with the complete degradation of ZX50 after 16 weeks,the amount of H2decreased.Despite large gas accumulation,ZX50 showed quick recovery without any permanent harm to the new bone (Fig.11(a-d)).After 36 weeks ZX50 was completely absorbed while WZ21 was still present but with reduced volume (Fig.11(g,h)) [41].

Fig.10.Graph depicting dependence of corrosion rate on grain size and volume fraction of the second phase of material [110].

Therefore,ZX50 could be a suitable option in case of pediatric trauma where a short life span of the implant is required.Another study considering pediatric orthopedics was done by using Mg0.45 wt.%Zn0.45 wt.%Ca (ZX00) alloy in a sheep model.This implant reduced risk of stress shielding and interference in the growth of bone physics which is otherwise observed in the case of permanent implants.The ZX00 had a higher YS and UTS and low degradation rate than Mg1wt.%Zn0.3 wt.%Ca (ZX10) alloy.ZX00 showed no negative impact on the growth of the physics of the sheep model [111].

Fig.11.Histological slides of pins of (a-d) ZX50 (e-h) WZ21 stained with Levai-Laezko after 4,12,24,and 36 weeks depicting the degradation of pins and new bone formation in a rat model [41].

Fig.12.Microstructure of Mg5ZnxSr alloys with different wt.% Sr (a) 0 (b)0.2 (c) 0.6 (d) 1 wt.% [114].

2.2.2.2.Mg-Zn-Sr alloys.Mg4Zn1Sr alloy was investigated for corrosion behavior of alloy in comparison to pure Mg.The alloy had a much lower degradation rate than pure Mg which was confirmed by immersion and electrochemical test.The Tafel curve indicated that the alloy had higher corrosion potential and lower corrosion current density compared to pure Mg [112].In-vitro cytocompatibility of Mg4ZnxSr(x=0.15,0.5,1,1.5 wt.%) alloys with human embryonic cells were evaluated by examining fluorescenc images of culture wells.All alloys had better cytocompatibility than pure Mg and Mg4Zn0.15Sr displayed the highest cytocompatibility among the alloy group [113].Ternary MgxZn0.5Sr (x=2,4,6 wt.%) and binary MgxSr (x=0.5,1,1.5 wt.%) alloys were compared in terms of their microstructure,mechanical and corrosion characteristics.The ternary alloys had lower grain size and possessed higher UTS,YS,elongation values compared to binary alloys.Moreover,the degradation rate of ternary alloys was less than binary alloys [101].The effect of Sr on microstructure,mechanical and bio-corrosion properties of Mg5ZnxSr (x=0,0.2,0.6,1 wt.%) alloys was examined.The average grain size of the four alloys was 129 μm,73 μm,63 μm,99 μm (Fig.12(a)-(d)).While the volume fraction for second phases were 1.2%,1.2%,2.5% and 2.3%[114].

As stated above grain size and amount of second phases are the two important factors that control the mechanical and corrosion behavior.Mg5Zn0.2Sr displayed the highest UTS while all the ternary Mg5ZnxSr alloys had a slower degradation rate than Mg5Zn binary alloy.Among all ternary alloys,Mg5Zn0.2Sr had the lowest corrosion rate [114].The work was further carried out to study the in-vitro cytocompatibility of Mg5Zn0.2Sr and Mg5Zn alloys.Mg5Zn0.2Sr possessed higher cell proliferation and better cell morphology compared with Mg5Zn [115].

Anterior cruciate ligament reconstruction (ACLR) surgery is performed to heal the most commonly occurring knee injury i.e.,ACL tear.Peri-tunnel bone loss is a trouble/failure of ACLR found in most clinical studies.MgxZnySr (x=4,6 wt.% andy=0.2,0.5,1,2 wt.%) alloys were fabricated to be used as interference screws for ACLR.It was found that Mg6Zn0.5Sr possessed the highest UTS and lowest corrosion rate compared to all the alloys as it had a low fraction of second phases [116].

The alloy possessed a higher maximum torque value(38.54±0.35 Nm) than pure Mg screw (25.25±0.06 Nm)which was in accordance with the numerical data.Biodegradable Mg6Zn0.5Sr and PLA screws were used in-vivo for ACLR.The results from high resolution peripheral quantitative computed tomography (HR-pqCT) conducted at 6-,12-and 16-weeks post-surgery indicated that Mg6Zn0.5Sr had better bone growth/formation than PLA screw and degradation of Mg6Zn0.5Sr screw did not have any negative effect on peri-tunnel bone mass (Fig.13) [116].

2.3.Mg-based composites

2.3.1.Mg-Hydroxyapatite composites

Human bone comprises plate-like hydroxyapatite(HA/Hap/HAP) nanocrystals having dimensions 30-200 nm in length and 2-7 nm thick.As inferred from bone anatomy,it is specified that nearly 65% of total bone mass consists of an inorganic solid,carbonate HA [117].This carbonate HA is involved in bone ingrowth and prevention of osteolysis by expanding along with intercellular networks of collagen I.HA is a calcium phosphate-based bio-ceramic that can be used to enhance the bioactivity of Mg scaffolds [118].It is also very important to understand that the correct amount of HA should be added which enhances mechanical properties,corrosion resistance,and bioactivity of the composite fabricated.MgxHAp (x=0,5,10 wt.%) composites were developed and analyzed for their microstructure and mechanical behavior.It was found that grain size reduced with the addition of HAp.The HAp phase present in the composites results in enhancement of the compressive strength of the fabricated composites while decreasing their tensile strength[119].In another study done it was observed that compressive strength of MgxHA (x=5,10,15 wt.%) composites increased with an increase in HA content [120].MgxHA(x=10,20,30 wt.%) composites were fabricated and evaluated for their UTS,corrosion properties,and cytotoxicity.With increasing HA content UTS and corrosion resistance decreased.Mg10HA showed no cytotoxicity while Mg20HA and Mg30HA displayed reduced cell viability [121].

Fig.13.HR-qpCT analysis of Mg6Zn0.5Sr and PLA screws on (a) bone volume (Bv),(b) trabecular number (Tb.N),(c) trabecular thickness (Tb.Th),(d)trabecular separation (Tb.Sp) of peri-tunnel bone in rabbit at 0,6,12 and 16 weeks post implantation and (e) degradation of two screws 16 weeks post implantation (scale bar 1 mm) [116].

It’s essential that the scaffold fabricated should possess modulus in the range of human bone.MgxHA (x=0,8,10,15 wt.%) composites were prepared and a nano-indentation test was conducted on these composite samples.The modulus of implants with HA reinforcement decreased as analyzed from the force-displacement curve (Fig.14) thus reducing the mismatch between the modulus values of bone and implant thus combating the problem of stress shielding.It was also found by the authors that the corrosion rate of MgHA implants decreased compared to pure Mg with Mg10HA exhibiting the best anti-corrosion behavior.In Mg/HA composites,HA acts as an appropriate second phase and promotes the deposition of Ca-P compounds over the surface of the composite.HA was found to aid in the growth of the apatite layer over the implant surface.This layer prevents further corrosion of the implant since it retards the contact between the implant and the physiological environment [122].

In-vitro tests on MgxHA (x=0,5,10,15 wt.%) composites were conducted to evaluate the effect of HA addition on corrosion behavior and bioactivity of the composites.The results of the electrochemical test indicated that HA addition was beneficial in improving the corrosion resistance of Mg.The fabricated composites had a lower rate of H2evolution and lesser change in pH of the medium than the pure Mg and with increasing HA content the H2evolution rate and variation in pH decreased.Considering the in-vitro cell tests i.e.,cytotoxicity and alkaline phosphate (ALP) activity assessment with MG63 cells,the composites displayed enhanced cell viability and ALP activity compared to pure Mg.Results of ALP activity concluded that HA addition had a positive influence on the bioactivity of Mg.The size of HA particles to be used for fabricating the Mg-HA composite too influence the properties of the composite [123].Mg-HA composites were prepared using different particle dimensions of HA (micron and nanoscale powders) powders and evaluated the effect of particle dimensions on the mechanical properties of Mg-HA composites.Composites fabricated using nano HA powders had higher hardness and ultimate shear stress than composites fabricated with micron-scale HA powders [124].

Fig.14.(a) Force-Distance (Nano-indentation) curves and (b) Bar graph depicting E values of MgxHA (x=0,8,10,15 wt.%) composites [122].

2.3.2.Mg-Bioactive glass composites

Bioactive Glass (BG) is considered a Class A bioactive material which implies,it can trigger bone growth as it can chemically bond with bone as well as soft tissues.The ability of BG to bond with bone contributes to the formation of HA over the surface[125].MgxBG(x=5,10,15 wt.%)composites were fabricated via microwave sintering and evaluated for mechanical and corrosion properties.Reinforcing BG in pure Mg improved its compressive and fl xural strength.H2evolution and pH of MgBG composites were less than pure Mg.Mg15BG composite showed the least H2evolution with a stable pH value.Moreover,MgBG composites displayed higher cell viability and ALP activity than pure Mg [126].Different amount of BG was reinforced in Mg to form MgxBG (x=0,5,10,15 wt.%) composites by spark plasma sintering.The addition of BG improved the hardness and elastic modulus of pure Mg.Hardness and elastic modulus of Mg10BG increased by 3 and 1.7 times respectively that of pure Mg.Mg10BG was found to have the least porosity [127].In another study,MgxBG (0,5,10,15 wt.%) composites were prepared via hot press sintering and investigated their corrosion behavior and cytocompatibility.The addition of BG reduced the corrosion rate of pure Mg as estimated from immersion and electrochemical test in phosphate buffer saline (PBS).Compared to pure Mg,corrosion resistance of Mg10BG increased 6.8 times while the volume of H2released from Mg10BG was reduced by 4.5 times.Cyto-compatibility of Mg10BG was best among other fabricated composites [128].

2.4.Coatings

Coatings play a vital role in inhibiting corrosion of Mg and its based implant alloys by protecting them from being in direct contact with the body fluids In the case of orthopedic implants with a coating containing bio-active components not only retards degradation of the implant but also induces a positive osteogenic response.

2.4.1.HA coatings

HA being bioactive ceramic and having chemical and crystallographic structures similar to bone can chemically bond with osseous tissue [121].The effect of HA coating on the degradation behavior of Mg was evaluated by immersing samples in PBS and revised SBF (r-SBF).It was found that HA coating was effective in reducing the degradation rate of Mg in both the electrolytes as it was observed that the Mg sample without coating represented severe corrosion attack in PBS compared to the sample with coating.The Mg sample without coating degraded completely in 21 days (Fig.15(a)) while the one with coating degraded in 31 days in r-SBF (Fig.15(b))[129].

HA-coated Mg implants were investigated for their in-vitro and in-vivo bio-corrosion,biocompatibility,and response in bone healing in comparison with non-coated Mg implants.The coated implants had a much lower corrosion rate compared to non-coated Mg implants both in-vitro and in-vivo.In-vitro biocompatibility test indicated that coated implants had more cells attached to their surface with well spreading cell morphology compared to non-coated implants as indicated by confocal laser scanning microscopy (CLSM) images(Fig.16(a,b)).Also coated implants had a higher BIC value than non-coated implants indicating a good response in bone healing [130].

Fig.15.Images showing degradation of (a) non coated Mg (b) HA-coated Mg in r-SBF [129].

Fig.16.CLSM images representing MG63 cells attached on (a) bare Mg (b) HA-coated Mg surface [130].

Through another in-vivo study,it was shown that HA coating reduced the resorption rate of Mg alloy screws compared to non-coated Mg alloy screws.HA coating is effective in reducing the corrosion of Mg alloys thus maintaining the pH of the medium in an acceptable range [131].HA was coated over porous Mg and investigated their effect on the pH of SBF.It was found that the rise in pH of SBF for HA-coated Mg was less compared to un-coated Mg [132].It is very important that coating done on a surface should adhere to it without the formation of cracks upon solidification The effect of surface roughness of Mg3Zn substrates was studied for coating HA on it.The coating showed better corrosion resistance and biocompatibility in-vitro in comparison with bare substrates.It was found that smooth substrate surface reduced the cracks formed in coating upon solidification and it improved the mechanical properties such as elastic modulus,hardness and fracture toughness of the coating.Moreover corrosion resistance of coating done on substrate with lower surface roughness was better than that of the coating done on substrate with higher surface roughness [133].The calcium ions present in HA coatings act as positive charge sites which favors the adsorption of fibronectin and vitronectin proteins,which are crucial for cell attachment and spreading [133].A coating containing HA particles was done on micro-arc oxidation(MAO)coated Mg alloy and inspected/evaluated for its blood compatibility by assessing in-vitro blood platelet adhesion and static water contact angle.A low static water contact angle value of the material surface indicates better hydrophilic nature.An implant material with a hydrophilic surface has a less chance of absorption of blood platelet,proteins,and blood clotting factors thus enhancing its blood compatibility.The coating containing HA particles had lower platelet density compared to MAO coating and the static water contact angle measurement for HA coating was 18.7°while for MAO was 40.6°indicating HA coating had better blood compatibility [134].In another study by Tiwariet al.,it was proved that HA coating improved the value of contact angle [135].Substituting various elements such as Mg,Sr,Si,Zn into the crystal lattice of HA is known to improve its bioactivity.In the human body,Si is known to be a crucial mineral.It has been specified that Si performs a vital role in boosting the immune system and aids in the healing operation of bone [136].The effect of Si substituted HA coating over Mg alloy on corrosion behavior was studied.To prepare the coating four distinct electrolytes with distinct molars of SiO32-(0,0.0025,0.005,0.0075) were investigated by the authors.They concluded that coating with 0.005 mol/l addition of Si in the electrolyte enhanced its corrosion resistance up to 7 times in comparison with simple HA coating [137].

2.4.2.Micro-arc oxidation (MAO)/ plasma electrolytic oxidation (PEO) coating

In this coating technique,a stable oxide layer is formed on the surface of a substrate which acts as an anode connected to a cathode (usually stainless steel) in an electrolyte solution.The oxide layer thus prevents early corrosion of implant and even plays role in enhancing bioactivity.The corrosion behavior of MAO-coated Mg-HA composites in SBF was analyzed through the electrochemical test.The results of the electrochemical test proved that MAO coating enhanced the corrosion resistance of Mg-HA composites [138].The coating process parameters such as current density,oxidation time,duty cycle,electrolyte composition,and applied voltage govern,morphology,the composition of the surface,and corrosion rate.The effect of current density (3,4,5 A/dm2) on surface morphology and corrosion resistance of MAO coating on AZ91D Mg alloy was studied.With an increase in current density surface roughness decreased while the number of pores on the coated surface increased.The sample coated with 5A/dm2displayed the best corrosion resistance while all the coated samples had better corrosion resistance than the bare alloy [139].Influence of current density (0.026,0.046,0.067 A/cm2) on surface roughness of MAO coated AZ31B Mg alloy was analyzed by Nithinet al.Surface roughness increased with an increase in current density while the porosity decreased.The coating done at 0.046 A/cm2had the maximum thickness and had the highest polarization resistance.The polarization resistance of all the coated samples was approximately 4-5 times greater than bare alloy [140].Effects of variable MAO coating time (1 to 22 min) on pure Mg were investigated for their corrosion behavior.The thickness of the coating for 17 min was reported to have maximum thickness and showed the best resistance to corrosion [141].The effect of two different MAO times (1 and 5 min) on the corrosion resistance of the MAO coating was evaluated.The thickness of coating done for 5 min was more than that done for 1 min while its reported porosity was less.From surface images of samples exposed to SBF for 7 days,it was clear that 5 min.the coating had better corrosion resistance[142].Another advantage of MAO coating is the increase in surface roughness which is considered to facilitate the proliferation of cells on it.In a study,it was found that MAO coating done on porous AZ31 Mg alloy increased the surface roughness of alloy nearly by 57% [143].MAO coating done on AZ31 Mg alloy was found to support the better proliferation of MG63 cells than the untreated substrate as the surface of MAO treated substrate was rough while that of an untreated substrate was smooth.The cell morphology on MAO treated substrate was round as compared to the untreated substrate,which was fla [144].The effect of applied voltage (300,360,400 V) on corrosion resistance and biocompatibility of MAO-coated Mg-1Ca alloy was investigated.The MAO coated samples had a much lower amount of hydrogen evolution and pH in Hank’s solution compared to the un-treated sample with the 360 V sample displaying the least corrosion rate.The cell viability of coated samples was better than untreated samples while 360 V and 400 V samples displaying appreciable ALP activity [145].Different duty cycles (10,20,30,40%) of the MAO process were applied for coating.It was found that an increase in duty cycle enhanced the adhesive bonding strength of coated surface as detected by the lap-shear test.This enhancement in bonding strength was due to an increase in porosity with an increase in duty cycle [146].MAO-coated AZ91D Mg alloy was evaluated for their in-vitro cytocompatibility.It was found that MAO coating was not toxic to MG63 cells,moreover,it enhanced the proliferation of cells as compared to bare AZ91D Mg alloy[147].In another study,it was proved that MAO coating displayed cytocompatibility with MG63 cells [144].The blood compatibility of the MAO-coated Mg1Zn1Ca alloy was evaluated by conducting a hemolytic assay.MAO coating showed improved blood compatibility as the hemolytic ratio of MAO coating was around 2.25% while that of the uncoated sample was around 24.58% [148].An optimal orthopedic implant should not degrade during the initial stages of implantation till fracture has consolidated.The implant should quickly degrade completely to avoid inflammator reactions after healing of fracture is complete.An in-vivo study was carried out on rats using MAO coated ZX50 pins and un-coated ZX50 pins to understand their degradation behavior and bone formation response.The coated pins showed a closer contact with new bone due to minimum callus formation as they slowed the degradation during firs 3 weeks of implantation and reduced the amount of H2gas evolved as compared to un-coated pins.After 3 weeks coated pins degraded much faster than un-coated pins till complete bone formation as clear from micro-CT study (Fig.17) [149].

Fig.17.3D reconstruction of implant site for bare and MAO coated ZX50 implants for different time durations (with evolved H2 gas bubbles in blue) [149].

From the bone histology,it was concluded that coated pins had more osteoblasts and enhanced new bone formation even in corrosion pits while un-coated pins had mesenchymal cells and fibroblast contained in fibrou tissue [149].Another invivo study was done by implanting MAO-coated Mg in the femoral condyle of white rabbits.The coated implants degraded 12 weeks post-implantation while un-coated degraded after 8 weeks.After 12 weeks the heart,kidney,spleen,and liver tissue showed no abnormalities proving biocompatibility of coating.The bone histology did 3 months post-implantation revealed that more bone had been formed near MAO coated implants as compared to uncoated implants which were also confirmed by 3D reconstruct models obtained from micro-CT.The tensile strength of coated sample after being immersed in SBF for 30 days was 165 MPa while that of the un-coated sample was 140 MPa [150].

2.4.3.Chemical conversion coating

Conversion coatings are formed by immersing the substrate in a chemical phosphate solution for a particular time.The solution reacts with the substrate surface to form a protective phosphate conversion coating which can even stimulate bioactivity.Different phosphate solutions have been used to form different coatings.AZ31 Mg alloy coated with Ca-P coating was studied for degradation behavior in SBF for 15 days.Mass loss for coated samples was reduced by 9% that of the uncoated samples [151].Further work was extended by implanting Ca-P coated and bare AZ31 Mg alloy in the thigh bone of rabbits to evaluate in-vivo degradation response.From micro-CT analysis,the authors concluded that coated implants had a slower degradation rate than un-coated implants.Histological evaluation of the bone-implant section revealed the formation of new compact and uniform osteoid tissue for coated implants compared to uncoated implants [152].The surface bioactivity of Ca-P coated Mg1.2Mn1Zn alloy was compared with that of un-coated alloy in-vitro and in-vivo.In-vitro coated samples showed better cell proliferation than un-coated samples.Both coated and uncoated samples were implanted in the femoral shaft of rabbits to investigate bone formation response.The mean optical density (MOD) values of BMP-2,TGF-β1,and PDGF expressions at the boneimplant interface were higher for coated implants compared to uncoated implants[153].Ca-P coating was done on MgxCa(x=1.34,5,10,16.2 wt.%) and Mg3ZnyCa (y=1.34,5 wt.%)alloys,and their mass loss and H2evolution rate were calculated upon exposure to MEM for 24 hrs.It was deduced that the effectiveness of coating was dependent on substrate composition.The coating could reduce corrosion rate in all samples except Mg3Zn which was due to its low anodic dissolution in phosphating bath [154].SrP conversion coating was done over pure Mg at different temperatures (40 °C,50 °C,60 °C,70 °C,80 °C) to inspect the corrosion behavior and biocompatibility of the coated samples compared to uncoated Mg.From electrochemical tests in MEM,it was deduced that coating reduced the degradation rate of Mg and H2evolution rate and coating is done at 80°C had the lowest corrosion and H2evolution rate.It was also observed that coating showed proliferation and differentiation of hMSC [155].Two-layer Mg-P coatings were prepared on AZ31 alloy by two steps processes.In the firs step bare alloys were treated in phosphoric acid while in the second step samples were immersed in ammonium hydrogen phosphate for 30 and 60 min.It was observed that coating formed at 60 min was dense with fewer cracks compared to 30 min.The corrosion resistance of coating formed in the second step was higher than the firs step while both coating layers had higher corrosion resistance than bare alloy [156].In another study,it was concluded that Mg-P coating done on AZ31 Mg alloy had a lower corrosion rate than bare alloy as revealed by electrochemical and immersion tests [157].The parameters of the conversion coating process such as pH of phosphating solution influence the rate of formation of phosphate crystals and rate of metal substrate dissolution during the process of phosphating while the temperature of phosphating solution affects electrochemical synthesis reactions and diffusion process [158].Ca-P conversion coatings were done over AZ60 Mg alloy at different pH values and temperatures of phosphating solution to understand their effect on corrosion behavior.Samples coated at pH 2.8 and temperature 37 °C had the lowest corrosion current density(Icorr) and the least amount of H2evolved [158].Mg-P coated pure Mg was evaluated for its in-vitro and in-vivo response.The corrosion rate of pure Mg reduced with Mg-P coating as measured in blood plasma.Coated samples showed better cell viability.The Mg-P coating reduced the amount of H2evolved and corrosion rate in-vivo as compared to bare implant [159].Mg-P was coated over AZ31 alloy with varying pH (4.5,5.5,6.5,7.5) of phosphating solution and their corrosion behavior was analyzed.Sample fabricated with pH 7.5 phosphating solution showed the least corrosion rate amongst other solutions [160].Two-layer Ca-P coatings were applied on pure Mg to investigate the effect of each process of surface roughness and corrosion resistance.The two coating processes consisted of the primary phosphating process(PRI)and secondary alkaline treatment (SEC).The corrosion resistance as evaluated from the electrochemical test followed the order un-coated＜SEC＜PRI [161].AZ31 Mg alloy was coated with different Zn-P,Ca-P,Mg-P type conversion coatings and evaluated their corrosion behavior and biocompatibility.From the immersion test in hank’s solution,the order of coating for corrosion resistance was Mg-P type＜Ca-P type＜Zn-P type.Compared to Mg-P,Zn-P type conversion coatings,and uncoated Mg alloy,Ca-P type coatings showed better cell viability with MC3T3-E1 cells [162].

Fig.18.Concentration of alloying elements at different intervals in blood after implantation (a) uncoated (b) MgF2 coated LAE442 [165] .

2.4.4.Polymeric coatings

Among the polymeric coating,Polycaprolactone (PCL) is one of the FDA-approved polymers for biomedical application.To control rapid corrosion of Mg in physiological fluid porous Mg scaffolds were fabricated and coated with PCL(3%,6%) coating of different concentrations.The authors observed that uncoated Mg scaffold degraded completely after 72 hr.in physiological saline solution (PSS) while the weight loss with 3% and 6% PCL was 36% and 23% respectively with a minimum rise in pH value for 6% PCL coated sample[163].Four bio-degradable polymer coatings (PLLA,PLGA(90:10),PLGA (50:50),PCL) were used for coating Mg samples for reducing corrosion rate and a comparative study on their interaction with human endothelial cells was carried out.PLGA (50:50) coating showed better adhesion of human umbilical vein endothelial cells (HUVEC) with a spreading morphology compared with the other three coatings.This might be because it had higher Mg2+and lower Ca2+concentrations in the culture medium.However,it was less effective in reducing the degradation rate of Mg than the other three[21].

2.4.5.Fluoride coatings

Our human skeleton and teeth contain fluorin as their natural component[164]so it is appropriate to use it for biomedical applications.So,fluorid coatings can be considered as a suitable option for inhibiting corrosion of Mg-based implants along with improving bio-activity.Corrosion inhibition and in-vivo trials of extruded LAE422 (Mg 90%,Li 4%,Al 4%,rare earth elements 2%) alloy with a surface coating of magnesium fluoride (MgF2) were studied.The alloy was divided into two groups one with MgF2coating and one without coating [165].

It was found that after 12 weeks of post-operation uncoated samples degraded 10%of their initial volume while the coated samples lost 4% of their initial volume.In the course of the firs six weeks of post-surgery,MgF2coating reduced the release of neodymium(Nd),cerium(Ce),and lanthanum(La)in the blood.But following six weeks the content of alloying elements in the blood increased from coated samples increased while that from uncoated was diminishing (Fig.18) [165].Another study on fluoride treatment for improving the corrosion resistance of Mg using 0.1 M and 1 M potassium fluorid(KF) solution was done.Authors reported that 0.1 M fluoride solution treatment (FST) showed better corrosion resistance than 1 M FST [166].

Fig.19.Blood compatibility of porous,uncoated and MgF2 coated Mg5%Al2O3 composites as indicated by%hemolysis [167].

A factor for failure of cardiovascular stents is thrombus formation due to low blood compatibility.An implant surface when comes in contact with blood at the implantation site adsorbs proteins.Increased platelet activation is triggered by these adsorbed proteins which may lead to thrombosis.In a study,MgF2was coated over porous Mg/Al2O3composite and their corrosion behavior and hemocompatibility were evaluated.It was found that coating not only improved corrosion resistance of the porous composite but also showed good hemocompatibility as the coating showed less platelet adhesion which could reduce the chances of thrombus formation.(Fig.19).

The surface of coated samples exhibited fewer cracks than uncoated ones after immersion in SBF.The amount of hydrogen released and pH value was less as compared to uncoated samples [167].The effect of fluoridate hydroxyapatite coating on MgZn alloy by conducting in-vitro tests with human bone marrow stromal cells (hBMSC) was examined.The coating supported bioactive nature for proliferation and differentiation of hBMSC [168].Fluoridated hydroxyapatite(FHA)/MAO composite coating was done on AZ31B Mg alloy and their corrosion behavior was studied.FHA coating was done by a simple hydrothermal process over MAO coating.The composite coating affected the morphology and phase composition with a change in fluoride ion concentration.FHA showed higher corrosion resistance and better biocompatibility than Ca-P/MAO and uncoated samples.With the increase in Fˉ,the coating became more uniform and compact[169].

Fig.20.Backscattered electron microscope images of Histological section of the bone-implant interface (a) without ZA coating (b) with ZA coating depicting bone ingrowth in a dog model (implant appears white,bone gray,and void spaces as black) [170].

2.4.6.Zoledronic acid (ZA) coating

ZA,nitrogen-containing bisphosphonate has been in use to treat various bone-related diseases.The effect of administering ZA via post-surgery intravenous injection in dogs in which porous tantalum implants were placed in their ulnae was studied.The authors analyzed the images of histological sections of implant-bone interface obtained through a back scattered electron microscope and found that mean bone ingrowth into the implants with ZA administration was 12.2%while the bone ingrowth into implants without ZA administration was 6.6% (Fig.20) [170].

Also,the size of bone islands formed in the pores of ZAtreated implants was 69% larger than un-treated implants.Thus it was concluded ZA could enhance bone ingrowth in orthopedic implants[170].Giant Cell Tumor of Bone(GCTB)is one of the benign and metastatic tumors in the bone.It is associated with high osteolytic abnormalities,bone destruction and resulting in progressive swelling near the joints.ZA coating on Mg-Sr alloy was investigated to be used as a therapeutic agent to inhibit GCTB.Three concentrations of ZA(10-2M,10-3M,10-4M designated as ZA2,ZA3,and ZA4 respectively) were used over Ca-P coated Mg-Sr alloys.From in-vitro experiments it was observed that ZA coatings did not induce hemolysis of human red blood cells (Fig.21(a)),ZA2 and ZA3 coating had lower absorbance which indicated they reduced cell proliferation of GCTB cells (Fig.21(b)),ZA2 coating had 553% cytotoxicity which marked the destruction of GCTB integrity (Fig.21(c)),DNA content of GCTB cells in ZA2 and ZA3 treated group was less indicating genotoxic nature of ZA to tumor cells (Fig.21(d)),ZA2 and ZA3 coating had less number of live GCTB than other coatings(Fig.21(e-i)).Therefore ZA could be considered as a GCTB inhibiting molecule [171].

Fig.21.In-vitro tests to measure response of ZA coating on GCTB cells (a) Hemolysis test represented as hemolysis ratio (b) CCK-8 assay used to measure cell proliferation (c) LDH assay used to measure the cytotoxic response (d) Cyquant cell proliferation assay kit used to measure DNA content of tumor cells(e)-(i) Live/Dead staining kit used to conduct Cell Viability (Dead tumor cells appear in red fluorescenc and live tumor cells appear in green fluores ence)[171].

Fig.22.Anti-bacterial test.(a) Re-cultivated E.coli and S.aureus colonies of samples.(b) Colony-forming unit numbers of the different samples.(c) Images displaying live/dead bacteria after staining E.coli and S.aureus on samples after 24 h incubation (d) Dead bacteria percentage [177].

One of the reasons for the delay in osteoporotic fracture recovery is the flow in osteogenesis and enormous osteoclast genesis.In a study ZA coated over Mg alloy was investigated.It was found that ZA coating promoted osteogenic differentiation of rat bone marrow stem cells and inhibited the proliferation and differentiation of rat bone marrow monocytes.Moreover,ZA coated Mg alloy implant withdrawn after 12 weeks from rat model had sufficient bending strength compared to bare Mg alloy and coating also reduced rapid bio-degradation of Mg alloy [172].It was found that ZA coating could be effective in resolving osteoclasts-related diseases,particularly peri-prosthetic osteolysis as analyzed from in-vitro tests of ZA coated Mg-Sr alloy.From in-vitro tests,it was found that ZA coating promoted osteogenic differentiation of preosteoblasts MC3T3-E1,and the coating induced apoptosis of osteoclasts and inhibited their differentiation which could be beneficial to treat bone defects [173].The degradation behavior of ZA coating in combination with fluoride pre-coated AZ31 Mg alloy was studied for orthopedic application.From the mass loss experiment in phosphate buffer solution (PBS)authors observed corrosion rate of ZA-coated samples was less than that of uncoated samples and also the pH value of coated samples was less than that of uncoated samples.So ZA coating apart from being used as a drug for various bone and musculoskeletal-related disorders could also control the degradation rate of temporary Mg implants [174].

2.4.7.Bone morphogenetic proteins (BMP) coatings

Osteoinductive factors are isolated from the demineralized bone matrix,known as bone morphogenetic proteins (BMPs).Using these growth factor-based substitutes is another strategy for bone tissue engineering.BMPs display osteoinductive properties and are a part of the transforming growing factorβsuperfamily.Recombinant human bone morphogenetic proteins (rh-BMP),rh-BMP2,and rh-BMP7 are the two potent growth factors that have been approved by FDA and found to be useful in a clinical setting for new bone formation [175].AZ31B Mg alloy was coated with different concentrations of BMP2 (20,50,100 ng/ml),and their in-vitro cytotoxicity,in-vivo biodegradation,and bone formation were evaluated.BMP2 was coated on AZ31B alloy in combination with MAO and hydrothermal surface treatment.The in-vitro analysis concluded that BMP-2 coating increased cell proliferation as assessed by water-soluble tetrazolium salt (WST) assay and induced higher ALP activity.In-vivo the implant with 50 ng/ml BMP-2 coating showed the lowest biodegradation rate as analyzed from 3D reconstruction models of micro CT data and better bone formation as revealed by histological examination of the bone-implant interface [176].

2.4.8.Zinc (Zn) and strontium (Sr) coatings

As discussed,earlier implant failure due to bacterial infection is a major concern,as sometimes it cannot be treated with antibiotics even.It is important that the implant possess anti-bacterial properties.Surface treatment of ZK60 alloy was carried out by incorporating Zn/Sr ions by hydrothermal process and studied its effect on corrosion behavior,osteoinduction,and antibacterial properties.Four groups were formed,one was un-doped,second with the only Zn2+,third with the only Sr2+,fourth was with combined Zn2++Sr2+.It was found that corrosion resistance of doped alloy improved than un-doped one.E.coli and S.aureus were used for the antibacterial test.Colony numbers of E.coli and S.aureus reduced for Zn and Zn+Sr doped group as calculated by counting fla colony (Fig.22).While no obvious difference was seen for Sr and un-doped groups.From the tests,it was concluded that Zn ions play the role of the antibacterial agent while Sr ions increase osteogenic activity [177].

3.Methods to fabricate porous implants

The human bone has a porous structure,with pore sizes ranging from 1 to 100 μm.The porosity of cortical bone is 5-13%about while that of cancellous bone is 30-90%[11][12],.A porous structure with appropriate pore size and interconnectivity of pores of the implant is crucial for it to be osteoconductive and osteoinductive.Osteo-conduction refers to the potential of a scaffold to facilitate the adhesion of osteoblasts on the surface and in the interior of the implant.Porous Mgbased implants are preferably used in bone tissue engineering [19].To fi fractured bone,or regenerating new bone to replace damaged bone,the scaffold is made porous for the bone tissue to grow into and through the pores,therefore making an appropriate bond with the implant.During an invitro experiment if an implant shows attachment,migration,and proliferation of osteoblasts it is regarded as an osteoconductive scaffold.An implant is said to be osteoinductive if it enhances the differentiation of mesenchymal stem cells down an osteoblastic lineage,finally resulting in the formation of mineralized tissue.The Osteogenicity of an implant requires it to stimulate bone formation starting from the beginning,which would happen in the deprivation of host cell invasion.Among the above-stated properties,osteogenic activity is one of the most important that bone implant material should possess.Pore size,pore structure,pore volume,and amount of pore interconnectivity are the parameters on which osteogenic activity depends [178].

Fig.23.Porous scaffolds of (a) Mg (b) Mg/1PCL-BG (c) Mg/3PCL-BG to allow cell migration and stimulate bone growth [179].

To ensure nutrient,oxygen dispersal,and removal of waste,3D scaffolds are designed to be highly porous with interconnected pore networks (Fig.23) [178].Pore size and porosity of implant have the potential to alter mechanical properties and direct cellular responses.Pore size and pore connectivity are the parameters that control the progress of bone growth into porous material.Regardless of implant material used,pore sizes should ideally range from 100 to 400 μm for bone ingrowth into porous-coated prostheses.To support continuous ingrowth of bone tissue a minimum pore diameter of 100 μm is needed [125].The proper pore size of the scaffold stimulates angiogenesis and new bone generation at the surgical site.Implant with a microstructure of interconnected pores leads to lower density and remarkably changed mechanical and deformation properties.So,pore size should be selected appropriately.Different techniques for the fabrication of porous implants have been discussed below.

3.1.Powder metallurgy (PM)

A promising technique to produce porous scaffolds is powder metallurgy in which spacer particles are used to generate pores.Governing the size and shape of the powder grains in the beginning material and by appropriately choosing processing parameters such as compaction pressure,sintering time,sintering temperature,etc.close control over pore size and distribution is possible [180].In the PM method,the starting raw powders are mixed with spacer agents and then compacted to form green compacts.The green compacts are then sintered to remove spacer particles and a porous scaffold is obtained (Fig.24) [181].

3.1.1.Effect of PM processing parameters on fabricated scaffolds

Processing parameters in PM such as compaction pressure,sintering temperature,and time affects the porosity along with mechanical and corrosion characteristics.The corrosion resistance of a material is significantly affected by changes in porosity.As porosity is increased,the corrosion rate also increases because the contact area between the corrosion medium and implant material increases.PM was used to fabricate Mg alloy and the response of different sintering temperatures (500 °C,550 °C,565 °C,580 °C) on mechanical properties of the alloy was studied.The authors reported that the grain size increased with an increase in sintering temperature.They identified 550 °C as the optimum sintering temperature,as the compressive strength at this temperature was maximum [86].Porous Mg alloy was prepared to study the effect of sintering temperature on the porosity and corrosion behavior of the alloy.It was found that porosity decreased while the corrosion resistance of porous alloy increased with the rise in sintering temperature [182].The effect of varying manufacturing conditions (compaction pressure,sintering temperature,and time) on the mechanical,electrochemical,and biocompatibility properties of AZ91D Mg alloy was investigated.Bending strength increased with an increase in compaction pressure whereas it decreased with an increase in sintering temperature and time.Also,samples with smooth surfaces exhibited lower corrosion resistance than those with porous and micro-textured surfaces as the proteins easily get adhered to the textured surface thereby inhibiting direct contact between corrosive medium and implant [183].Pure Mg was processed by powder metallurgy and investigated the difference between hot and cold-pressed samples by evaluating the effect of compaction pressure along with sintering temperature.The authors reported a decrease in the porosity of specimens on increasing compaction pressure and sintering temperature.In the case of hot-pressed samples,pores present in micro-structure were smaller as compared to coldpressed samples.Hot-pressed samples exhibited higher fl xural strength (Fig.25(a)) and hardness (Fig.25(b)) along with better corrosion resistance compared with cold-pressed samples [184].

Milling time too influence the mechanical and corrosion characteristics of fabricated alloys.Mg7Ca alloy was prepared by PM by varying milling time (3,5,8 hrs.) and sintering temperature(450°C,550°C)to understand their effect on the physical properties of the alloy.The value of hardness was enhanced with increasing milling time and sintering temperature due to the formation of the Mg2Ca phase but at an expense of increased corrosion rate [185].To figure out the response of sintering temperature on mechanical integrity of bio-degradable Mg6Zn1Ca alloy,sintering was carried out at two different temperatures.Mechanical properties were found to be affected adversely with the rise in sintering temperature as the amount of magnesium oxide (MgO) phase increased with the rise in sintering temperature [186].To analyze the impact of compaction pressure,sintering temperature,and time on porosity and micro-hardness,Mg alloy was synthesized by powder metallurgy.Authors found that on increasing compaction pressure,micro-hardness increased while the porosity decreased as the particles are attached more firmly whereas porosity increased on increasing sintering temperature and time.The particle size was reduced with increasing blending time with homogeneous distribution [187].Porous Mg alloy was manufactured for tissue engineering application via powder metallurgy and studied the effect of processing parameters on UCS and E.Increasing compaction pressure,sintering temperature and time had a positive influence on UCS and E of alloy [188].To investigate the impact of sintering temperature on microstructure and mechanical properties,Mg3Zn and Mg3Mn alloys were fabricated via powder metallurgy.Authors concluded that with the rise in sintering temperature porosity of the alloys firs decreased and then increased while the micro-hardness of Mg3Zn firs increased and decreased and that of Mg3Mn firs decreased and then increased [189].Mechanical alloying of magnesium hydride powder was carried out in planetary ball mill with varying vial filling factor and time of milling to understand its effect on crystallite size and desorption temperature of powder.The crystallite size and desorption temperature reduced with increasing vial filling factor and milling time indicating that increase in both the ball mill parameters was beneficial [190].

Fig.24.Powder Metallurgy process using spacer particles for fabrication of porous scaffolds [181].

Fig.25.Graphs depicting the effect of compaction pressure and sintering temperature on (a) fl xural strength (b) microhardness of pure Mg processed by PM[184].

3.1.2.Spacer agents used in PM for fabricating porous scaffolds

Various spacing elements/agents have been used to create a porous structure for bone tissue engineering via PM technique.Urea/carb-amide(CO(NH2)2)a human waste product occurring in urine has been used as a spacing agent[186,187,181,80,188] as it won’t harm the body even if its traces are left after sintering in the implant material.The decomposition temperature of urea is about 135°C.Ammonium bicarbonate (NH4HCO3) has also been investigated as a spacing element to fabricate porous implants [189],[190],[177],[191],[81].The decomposition temperature of NH4HCO3is between 36 and 60°C and is conveniently removed by the sintering process.Porous Mg scaffolds have been fabricated using sodium chloride(NaCl)as the spacer agent[127,162,192].Out of various spacer agents,NaCl can be leached out using NaOH solution while other spacer agents get evaporated when samples are sintered.Another spacer agent that has been used to manufacture porous scaffolds is naphthalene [193].It is efficiently removed during the sintering process as it sublimes at 80 °C.Polymethyl Methacrylate (PMMA) has also been used as a spacer agent to manufacture Mg-based implants with porous structures [194].The decomposition temperature of PMMA is 165 °C.From all the literature cited above,it could be concluded that using spacer agents to fabricated porous scaffolds decreases the mechanical properties and enhance corrosion rate as the pores formed to act as stress regions in the scaffolds.

3.2.Titanium wire space holder (TWSH) technique

The pores generated by spacer particles in PM are not uniformly distributed and also the shape and size of the pores also vary.To get a uniform network of interconnected pores of required shape and size TWSH technique is a good option.In this technique,Ti wires are used as a spacer agent.There are two methods by which the TWSH technique could be used.It can either be used with Mg melt or with Mg powder mixture.Ti wires of the required diameter are selected to prepare a 3D entangled structure which is used as a spacer.In the firs method,a porous Mg structure can be fabricated by using the TWSH technique in which Mg ingots are melted in a protective SF6+CO2atmosphere and Ti wires are immersed in this melt to produce Ti/Mg composite.Finally,the composite is etched off particularly in hydrofluori (HF) acid to eliminate Ti wires(Fig.26(a))[191].It has been found that after the etching process is over MgF2was present on the surface of the sample which could further reduce the corrosion rate of the Mg implant [192].In the second method of PM route,the starting powders are mixed and then compacted in die along with Ti wire mesh.After compaction,the samples are sintered,and finally to remove spacer samples are dipped in HF solution (Fig.26(b)) [188].With a suitable selection of Ti wire diameter and configuration of its assembly,the porosity% and pore shape/structure can be manipulated effectively.The pore size required to be generated can be manipulated by the diameter of the wire used.Porous Mg samples for orthopedic application were made using the TWSH technique.Ti wire of diameter 270 μm as used as spacer element.With increase in porosity,yield strength and modulus of elasticity decreased [191].To evaluate the effect of pore size of scaffold on bone regeneration,TWSH method was used to fabricate porous Mg implants.Samples with different pore sizes (250 μm and 400 μm) having the same porosity (55%)were fabricated.The values of compressive strength 250 μm and 400 μm scaffolds was 41.2 MPa,46.3 MPa while E of 250 μm and 400 μm scaffolds was 2.18 GPa,2.23 GPa respectively [192].

Implants with 400 μm pore dimensions degraded faster in-vitro.Considering the osteogenic effect,400 μm scaffolds induced more new bone generation compared to 200 μm scaffolds in-vivo as indicated by micro-CT analysis.(Fig.27)[192].

Fig.26.Process representing various stages of Titanium wire space holder technique with (a) Mg melt [191] (b) Mg powder mixture [188].

Fig.27.Micro CT 2D of newly formed bone after 8 and 16 weeks of surgery along with 3D reconstructive model representing newly formed bone [192].

It was reported that magnesium fluoride formed on the surface of the scaffold after etching was beneficial for the growth and mineralization of osteoblasts in vitro and stimulated bone formation in-vivo [192].Porous Mg10Zn alloy was fabricated through a PM route using Ti wires of diameter 300 μm as a space holder which generated samples with a porosity of 60% and UTS of 78 MPa [193].Porous Mg10Zn4Y alloy was manufactured by PM using Ti wires as a spacing agent.Compressive strength was found to decrease with increasing porosity [188].

3.3.Additive manufacturing (AM) techniques

AM techniques have emerged as an excellent choice of manufacturing porous implant structures with complex geometries which are otherwise difficult to fabricate through conventional manufacturing techniques.AM allows the fabrication of implants with internal and external features matching those of human bone [194].In AM the material is built layer-by-layer.AM techniques are basically of two types powder bed fused (PBF) and directed energy deposition (DED).

3.3.1.Powder bed fused (PBF) technique

The powder Bed Fused (PBF) technique is also known as selective laser melting (SLM).SLM technique utilizes high energy density laser for rapid melting and rapid solidification of powder to obtain a bulk material.In this process,the starting powders are deposited on the work bed through a feeder gun and then melted using a laser beam.Argon (Ar) gas is supplied so that oxidation of powder material does not occur.The powder then cools and solidifies on which the next layer of powder is feed.The entire process repeats to fabricate the required material.The process is controlled by instructions commanded through a computer system thus leading to 3D structures (Fig.28)

Fig.28.Laser additive manufacturing (LAM) process schematic [195].

Samples produced utilizing SLM have refine grain structure and higher micro-hardness compared to cast Mg samples[195,196].Porous WE43 Mg alloy was fabricated using SLM technique and studied its cytotoxicity and mechanical properties were studied upon immersion in body fluids It was concluded that samples showed zero level cytotoxicity and retained mechanical properties near trabecular bone even after degradation in body fluid for 4 weeks [197].Mg9%Al alloy was sintered using different combinations of laser power and scanning speed.At laser power,15 W and scanning speed 0.02 m/s alloy with minimum porosity and maximum density was reported.The temperature of the melt pool is directly proportional to laser power while inversely proportional to scanning speed [198].In another study,the effect of laser parameters on sintering of biodegradable WE43 Mg alloy was evaluated by examining the weld track deposits.Stable and continuous weld deposits were obtained within a suitable range of surface energy density (Es).The micro-hardness was not affected by the change in laser power and scanning speed[199].The effect of varying scanning velocity and hatch spacing on the microstructure of AZ91D Mg alloy were investigated in terms of volume energy density (Ev).The surface quality improved with increase in Evup-to a certain value,any further increase increased the surface roughness [200].In another work,Mg11Gd1.77Zn0.43Zr alloy was fabricated via SLM by varying scanning speed and hatch spacing.Grain size and the fraction ofβ-(Mg,Zn)3Gd phase reduced with an increase in scanning speed.As compared to cast alloy the tensile properties of alloy fabricated via SLM were superior.The alloy fabricated at scanning speed 300 mm/s with hatch spacing 100 μm had the highest UTS [196].Also,the above two studies proved that the relative density of material decreased with an increase in scanning speed and hatch spacing[200,196].Effect of varying laser power,pulse width,frequency of SLM on the microstructure of Mg-Ca alloys was evaluated in terms of Ev.XRD plot analysis revealed that peaks ofα-Mg shifted to low angles with an increase in Ev(Fig.29).With the rise in Ev,the temperature of the melt pool increased resulting in the evaporation of Mg.It also increased Ca content in theα-Mg matrix due to which the lattice parameters ofα-Mg became larger.The porosity of fabricated samples was reported to decrease with an increase in laser power,pulse width,and frequency of process [195].

Different values of laser power were employed to fabricate ZK61 Mg alloy via LAM.It was observed that porosity decreased with an increase in Ev.The surface quality improved with the increase in Evup to a certain value before resulting in a poor surface finis with a further increase in Ey[201].SLM was used to fabricate MAP21 alloy (composition comprising of elements Nd,Gd,Zn,Zr) with a porous structure.Samples with porosity in the range 0.6-2.4% were fabricated[202].The effect of varying laser power on the surface morphology of AZ61D alloy fabricated via SLM was studied.The alloy displayed smooth surface morphology when laser power was in the range 85-95 W.Relative density and grain size decreased with an increase in laser power while the amount of the second phase increased [203].

Though PBF is a useful technique for producing complex structures but to improve the efficiency of this method either laser power or scanning speed has to be increased.High values of scanning speed lead to improper fusion of powder particles.On the other hand,if the high laser power is employed then the material evaporates as the boiling point of Mg-based alloys are low due to which dense structures are not formed.So another AM technique i.e.,wire arc additive manufacturing has been used as an alternative to PBF [204].

3.3.2.Wire arc additive manufacturing (WAAM)

Wire arc additive manufacturing (WAAM) comes under the DED type of AM process.WAAM technique employs wire material as the feedstock.The feedstock is supplied at a constant speed and arc generated which acts as the source of heat.The wire is melted and deposited on the substrate and solidified A protective atmosphere of Ar gas is used to prevent oxidation.To avoid overheating of the welding machine and the substrate cooling chamber is also provided in the setup (Fig.30).

Fig.29.(a) XRD plot of Mg-Ca alloy at different volume energy density (Ev),(b) &(c) Enlarged view of section A and B [195].

Fig.30.Schematic diagram for WAAM [205].

Fig.31.Schematic of 3D Powder Printing [211].

Guoet al.manufactured AZ31 alloy using WAAM at different pulse frequencies (1,2,5,10,100,500 Hz) and investigated its effect on microstructure and tensile strength.At frequencies 5 Hz and 10 Hz least grain size and highest UTS[206].Porous AZ31B alloy was fabricated using WAAM by varying current and torch feed speed.At low current value,the torch feed speed decreased so less material was available for melting which led to improper weld bead shape or humping.When torch feed speed was increased,the microstructure got refine but humping occurred because of improper bead formation.The authors also observed that a higher feed speed resulted in fine microstructure [205].AZ80M alloy was fabricated via WAAM and its tensile strength was tested in the horizontal and vertical direction.The average YS and UTS of samples in the horizontal direction was higher than in the vertical direction [207].In a study by Holguinet al.different values of current and speed of feed,the torch was investigated and was used to fabricate AZ91D alloy by WAAM.The contact angle between the printed sample and printing substrate decreased with increasing current and decreasing speed [208].

3.3.3.Three-Dimensional (3D) jet printing

3.3.3.1.With binder agents.3D Jet printing techniques involve two steps.In the firs step,the powder materials are feed along with a binder agent to form 3D structures.In the second step binder is removed through the sintering process (Fig.31).3D powder printing was used by Meiningeret al.to manufacture Sr substituted magnesium phosphate cement with degassed ultrapure water as a binder.The samples were sintered after printing to obtain the final structure.This process generated samples with a porosity of 15-25%[209].Meiningeret al.also fabricated Sr substituted magnesium phosphate scaffolds using 3D powder printing followed by sintering.In this study,an aqueous solution (TCP:15 v/v% H3PO4,MgP:1.0 mol NH4H2PO4) was used as a binder.Samples with compressive strength of 16.1 MPa were reported by the authors [210].

Pure Mg scaffolds were fabricated using 3D printing with different vol.% of Mg powder in ink loaded in the nozzle for printing.The ink is comprised of Mg powder along with binder polystyrene,chloroform,and dibutyl phthalate.It was observed that with an increase in vol.% of Mg powder in ink a greater amount of pressure was required for the ink to flow through the nozzle [211].

3.3.3.2.Without binder agents.Binder-free 3D jet printing has also emerged owing to the troubles faced while using binder jet printing techniques.There are binder-particle interactions during the de-binding process i.e.,the sintering process which results in a change of chemical composition.The process becomes faster as no viscous binder agents are used.Capillary forces within the powder particles bind them together.A single-phase solvent is used to bind particles.Salehiet al.fabricated Mg5.9Zn0.13Zr samples via capillarity-induced binder-less 3D printing.Samples were sintered at different temperatures and for different periods.Increasing the sintering temperature improved mechanical properties but distorted the shapes of samples due to swelling.Whereas,increasing the sintering holding time improved mechanical properties.Samples with porosity 29% and average pore size 15 μm having modulus and compressive strength comparable to that of cortical bone were fabricated [212].Salehiet al.used 3D printing to fabricate Mg5.9Zn0.13Zr alloy samples without using binder agents.The as-printed samples were sintered to obtain final structures.The ultimate compressive strength (UCS) of samples fabricated was greater in comparison to bone and PCL [213].Microwave sintering of inkjet 3D printed Mg5.06Zn0.15Zr alloy samples was done by Salehiet al.Capillarity-directed binder-free 3D printing was used.Microwave sintering done at a higher temperature(＞510 °C) distorted the shape of the sample.Microwave sintering reduced sintering time by a factor of two to three times and enhanced densification of green compact due to interaction between interparticle bridges and applied electromagnetic field [214].

3.3.4.Paste extrusion deposition (PED)

Another AM technique is PED.In SLM,in 3D printing,the material has to be sintered at higher temperatures so that the final product is formed.The human bone comprises HA along with organic components(collagen).To mimic the bone composition organic materials can be added to the scaffolds.The organic components cannot withstand high sintering temperatures and are therefore denatured.

PED is carried out at controlled temperatures (near ambient temperature) which allows the addition of organic components.The material is fille in a syringe and extruded out from the nozzle.The then deposited material is allowed to harden.This technique can be useful to fabricate scaffolds containing drugs that can aid in the healing process of bone tissue [194][215],.Porous MgP scaffolds comprising of gelatine as the organic material was developed using PED.Different wt.%of gelatine was used and its effect on wettability of scaffold surface was investigated.The authors observed that with an increase in wt.% of gelatine the hydrophilic nature of scaffolds increased.Moreover,lysozyme protein was incorporated in the scaffold mixture to observe drug release effica y.The release rate of lysozyme increased with an increase in gelatine content [216].Drugs could be incorporated in scaffold structure as PED was carried out at controlled temperatures.MgPSr30PCL scaffolds were fabricated via extrusion-assisted 3D printing.The scaffolds possessed compressive strength of 4.3 MPa and displayed bone regeneration both in-vitro and in-vivo [217].

3.4.Electric discharge drilling (EDD)

EDD is a non-contact thermal erosion process used for making pores in electrical conducting metallic materials.

The schematic in Fig.32 represents the different parts of the common EDD setup.The required shape can be removed from the work material using an electrode of that shape.The most important requirement is that the work material should be conductive.The voltage variation between the work material and electrode causes an electric spark hence pores can be generated.Perforated AZ31 samples were prepared using EDD to evaluate the effect of porosity on apatite formation in body fluids Authors reported the formation of apatite on perforated AZ31 samples in 4 days post immersion in SBF while non-perforated samples had no apatite formation till 14 days post immersion (Fig.33) [219].

Fig.32.Process specification of Electric Discharge Drilling (EDD) [218].

Ahujaet al.used EDD to fabricate perforated ZM21 (Mg 97 wt.%,Zn 2 wt.%,Mn 1 wt.%).Effect of the process parameters of EDD (electrode material,discharge current,pulse on-time,pulse off-time) was investigated and optimum processing parameters which would produce samples with good hole circularity.It was reported that out of copper and brass,the brass electrode with discharge current 3 A and pulse on and off time 60 μs and 50 μs respectively produced holes with minimum overcut [218].

3.5.Melting techniques

In the case of producing porous scaffolds with melting techniques,there can be two options.The firs method is the melt foaming method.In the firs option,Mg and its alloys foams are produced using CaCO3as the blowing agent in the molten melt of the material.The blowing agent is added after melting the starting material and the CO generated due to its decomposition produces the required foam(Fig.34(a)).CaCO3was used as a blowing agent to fabricate AZ91 and AZ60 Mg alloys foam structures via the melt foaming method.From the interpretation of TGA-DTA curves of Mg and CaCO3,it was verified that CaCO3is suitable for forming Mg foams in the temperature range 690°C to 750°C[221].In similar work,AZ91 and AZ60 Mg alloy foams were manufactured via the melt foaming method using CaCO3as a blowing agent.It was found that bulk porosity and mean pore size of samples increased with increasing foaming temperature [220].CaCO3was used as a blowing agent to develop AZ31 Mg alloy foam.The yield strength (compressive) obtained from the tests of samples with porosities 60%,65%,70%,and 75% were 20.55,16.98,13.45,and 2.26 MPa respectively [222].In the second option,a NaCl preform is fabricated may be using rapid prototyping.The material (Mg or its alloys) is melted and then infiltrated/penetrate inside the NaCl preform.After solidification NaCl preform can be removed using distilled water and NaOH to generate a porous structure (Fig.34(b)).The second method is called the negative salt pattern or infiltration/replication technique.In both methods,melting is carried out in an inert atmosphere to prevent the ignition of Mg.Mg foams were prepared using a negative salt pattern with rapid prototyped NaCl preform.The surface roughness of samples increased with an increase in infiltration pressure [223].The infiltration method was used to fabricate five different Mg alloys with porous structures.Compressive strength decreased with increasing porosity of samples.Also,it was noticed that with an increase in infiltration pressure the porosity of samples decreased [224].Four porous Mg alloy structures were fabricated using the negative salt pattern technique.A solution comprising of equal concentrations of Cl and F ions was used to leach out NaCl from four Mg alloys to avoid corrosion attack while the leaching process.Samples of porosity of 75% with 500 μm average pore size were made [225].

Fig.33.(A) Non-perforated AZ31 alloy post immersion in SBF a) 4 days b) 7 days c) 14 days d) 21 days.(B) Perforated AZ31 alloy post immersion in SBF a) 4 days b) 7 days c) 14 days d) 21 days [219].

Fig.34.(a) Set-up to produce Mg and its alloys foam using blowing agent [220] (b) Negative salt pattern or infiltration technique [19].

4.Conclusion

In brief,this article discussed applications of Mg as orthopedic implants.Despite having the advantage of being osteoconductive and anti-bacterial,Mg faces a major challenge of rapid corrosion in body fluid due to which the mechanical integrity of the implant is affected and it degrades before the bone tissue has completely healed.To slow down the rapid corrosion rate of Mg various techniques such as mechanical working,heat treatment,alloying,and surface coating are done.Following conclusions can be drawn from this review:

1.Mechanical working and alloying reduce grain size thereby enhancing corrosion resistance of Mg.Heat treatment of Mg-based alloys reduces the number of second phases thus improving corrosion resistance Alloying of Mg with elements like Zn,Ca and Sr improves the microstructure,and enhance mechanical and corrosion properties of Mg alloys.

2.Mg-BG composites exhibit suitable mechanical properties,corrosion resistance and enhanced cytocompatibility for bone tissue repair as BG are bioactive.

The biocompatibility of reinforcement to be added in the Mg matrix is the most important factor.If the implant is not biocompatible it would lead to adverse tissue reactions.

The addition of different elements results in different phases.Thus,affecting mechanical and corrosion characteristics.

3.Coatings create a barrier so that the implant does not come in contact with the body fluid directly as soon as the implant is placed inside the body.Compared to an uncoated implant,a coated implant shows a much slower corrosion rate in body fluids To improve the performance of the coating,certain surface roughness is necessary so that bone cells can attach for bone regeneration.Drugs can be incorporated in the coatings for faster recovery of damaged bone tissue.

4.Method of fabrication and process parameters have a significant effect on microstructure,thus affecting the various properties of the implant.Porous implants enhance osseointegration but compromise corrosion resistance and mechanical strength.3D printing is an appropriate manufacturing technique to fabricate porous implants with which we can have control the complex internal structure of the implant to improve the mechanical strength of the implant.

Future scope for the development of biodegradable Mgbased implants should focus on the use of biocompatible alloying elements,an appropriate method of fabrication of porous scaffolds,and extensive in-vitro and in-vivo experimentation.Therefore,more research needs to be carried out related to the fabrication of implants via additive manufacturing,especially 3D printing.

Declaration of Competing Interest

None.

Funding

This research did not receive any specific grant from funding agencies in the public,commercial,or not-for-profi sectors.