Surface modification of biomedical Mg-Ca and Mg-Zn-Ca alloys using selective laser melting:Corrosion behaviour,microhardness and biocompatibility

Xiyu Yao ,Jinheng Tang ,Yinghao Zhou ,Anrej Atrens ,Matthew S.Dargush ,Bjoern Wiese ,Thomas Eel,Ming Yan,?

aSchool of Materials Science and Engineering,Harbin Institute of Technology,Harbin 150001,China

b Department of Materials Science and Engineering,Southern University of Science and Technology,Shenzhen 518055,China

c The University of Queensland,School of Mechanical and Mining Engineering,Centre for Advanced Materials Processing and Manufacturing(AMPAM),QLD 4072 Australia

d Institute of Materials Research,Helmholtz-Zentrum Geesthacht,Geesthacht 21502,Germany

Abstract Magnesium alloys such as Mg–Ca and Mg–Zn–Ca are good orthopaedic materials;however their tendency to corrode is high.Herein we utilize selective laser melting(SLM)to modify the surface of these Mg alloys to simultaneously improve the corrosion behaviour and microhardness.The corrosion rate decreased from 2.1±0.2mm/y to 1.0±0.1mm/y for the laser-processed Mg–0.6Ca,and from 1.6±0.1mm/y to 0.7±0.2mm/y for laser-processed Mg–0.5Zn–0.3Ca.The microhardness increased from 46±1 HV to 56±1 HV for Mg–0.6Ca,and from 47±3 HV to 55±3 HV for Mg–0.5Zn–0.3Ca.In addition,good biocompatibility remained in the laser processed Mg alloys.The improved properties are attributed to laser-induced grain refinement,confined impurity elements,residual stress,and modified surface chemistry.The results demonstrated the potential of SLM as a surface engineering approach for developing advanced biomedical Mg alloys.

Keywords: Mg alloys;Selective laser melting;Surface modification;Corrosion behaviour;Microhardness.

1.Introduction

Magnesium alloys in general exhibit favourable biocompatibility,low density,and a low elastic modulus that is similar to that of the human bone.They are highly promising candidates for degradable orthopaedic implants [1-5].However,their high sensitivity to impurities and microstructures is a major factor limiting their wider clinical applications [6,7].Achieving lower corrosion rates is a critically important research target for developing advanced biomedical Mg alloys.Furthermore,the comparably low wear strength of most Mg alloys necessitates research attention.

The corrosion behaviour of Mg and Mg alloys is known to be affected by the microstructure [8],grain size [9],secondary phase,[10,11]and surface contamination [12].Traditional methods to decrease corrosion rates are either physical and/or chemical approaches.One approach is to create a surface coating to protect Mg alloys from corrosion [13].

The mechanical properties of Mg alloys can be increased by alloying them with Ca and Zn,both of which are essential,non-toxic elements to the human body [14].Any alloying element,however,may increase the corrosion tendency by producing second phases,which accelerate the corrosion of Mg alloys by forming micro-galvanic couples between the second-phase particles and the Mg matrix [15,16].Hence,the Ca concentration in the binary Mg–Ca alloy should be below 1wt.% [14].Similarly,the Zn content of the ternary Mg–Zn–Ca alloy should be less than 5wt.% [17].

As an advanced tool,laser has been utilized to fabricate and/or modify Mg and Mg alloys to improve their mechanical and corrosion resistance properties.For instance,laser shock peening has been used to generate compressive residual stresses on the surfaces of pure Mg and Mg–Zn alloys to improve their corrosion resistance;the corresponding compressive residual stresses reached 30–50MPa and a depth of ～1mm [18].However,crack formation during processing should be avoided[19].Moreover,the surfaces of commercial Mg alloys AZ31B [20]and AZ91D [21]have been modified via laser treatment,resulting in grain refinement and intermetallic phases,which further reduced the corrosion rate.The aluminium in the compositions rendered these Mg alloys unsuitable as biodegradable implants [22].Laser-based additive manufacturing for Mg alloys has been investigated as well,although Mg powders are extremely flammable [23,24].

SLM is a mainstream laser-based additive manufacturing(AM)method.Whilst it is most often used as a bulk material processing technique [25,26],it can be used for advanced surface processing.Compared with the previous studies [18-22],SLM shows certain advantages when used as a surface modification approach:(a)SLM can conveniently raster a small(～60μm)focal point over a targeted surface.The small spot radius results in a small molten pool of approximately 100μm in width and a fairly limited heat affected zone(HAZ);(b)SLM controls the scanning path of the laser,thereby avoiding serious deformations due to the thermal stress accumulated during laser processing;and(c)SLM involves rapid melting and solidification.The high cooling rate may reach 106–107K/s [27]and typically decreases the grain size [28],which is expected to increase the hardness and strength,while simultaneously decreasing the corrosion rate [29-31].SLM is highly automated,convenient to use,and environmentally friendly.It produces no waste chemicals,such as dichromic acid(Cr6+),which is often used for chemical surface modification but is harmful to environment and can be hazardous to the human health [32].

In this study,SLM was used for the surface modification of two typical biomedical Mg alloys:Mg–0.6Ca(wt.%)and Mg–0.5Zn–0.3Ca(wt.%).The purpose was to develop biomedical Mg alloys with a low corrosion rate and high hardness/wear resistance.It will be shown that the research targets were achieved and that SLM performed well for both the alloys,demonstrating its potential in the development of advanced biomedical Mg alloys through surface engineering.

2.Methods

2.1.Materials and SLM processing

A HANS M160 SLM system(Han’s Laser Smart Equipment Group)was used to process the Mg alloys.The light source was a continuous wavelength ytterbium fibre laserunder the TEM00working mode(IPG Laser Model YLR-300-AC).The laser beam diameter was ～60 μm with a wavelength of 1070nm.The laser scan speed and laser power were important SLM parameters,and they were varied for optimization,as will be discussed later.The laser beam position was controlled using a SCANLAB scanning system(intelliSCANR○III 20,SCANLAB),as shown in Fig.1(a).In this study,the hatch distance was 30μm.An‘island’scanning strategy was adopted,which can reduce the residual stress and prevent large macroscopic deformations[33].Every single‘island’was a 3mm×3mm square area,on which laser scanned linearly.The laser scanning direction in the adjacent square was perpendicular to the previous path.The whole scanning range was 15mm×15mm,which encompassed the surface of sample,with diameter of 10mm.The laser scanning path and order are shown in Fig.1(b).SLM was performed in a nitrogen atmosphere,which maintained the oxygen content to below 50ppm.

The as-extruded Mg alloys(?10 mm;prepared by Helmholtz-Zentrum,Geesthacht)were cut using electron discharge machining(EDM)into 2.5mm thick discs,as shown in Fig.1(c).The compositions of the alloys are listed in Table 1.Before laser processing,the bottom and top sides of the specimens were ground using 2000# sandpaper.

Table 1 Composition of the as-extruded Mg alloys,wt.%.

2.2.Microstructural analysis

A digital optical microscope(OM;Keyence VHX-6000)was used to measure the surface roughness.Roughness measurements were conducted on the OM.High-resolution images of sample surfaces were captured;subsequently,threedimensional image stitching was developed with a report on surface roughness value.The microstructures of the original and laser-processed specimens were characterized using scanning electron microscopy(SEM;5kV and 100 pA)and electron backscatter diffraction(EBSD),which was conducted on a ZEISS MerlinR○(ZEISS Group)instrument,operated at 25kV and 12 nA at a step size of 0.3 μm.Specimens for the EBSD measurements were ground in the order of 800#,1000#,1500# and 2000# sandpapers,polished by a VibroMetR○2 vibratory polisher and then thinned by an ionmilling device(Model 695,Gatan PIPS II),with a beam energy of 6keV and a milling angle of 6° for final polishing.The phase composition of the specimens was determined by X-ray diffraction(XRD,Rigaku Smartlab,operated at 45kV and 200mA)with Cu Kαradiation,at a scanning speed of 2°min and a step size of 0.02°/s from 20° to 90° The surface residual stress was measured using the same equipment,using a diffraction angle of 96.82°,i.e.the(211)plane of Mg.Hooke’s law,as shown in Eq.(2.1),was used to determine the surface residual stress of the as-processed specimen [34]:

Fig.1.(a)Schematic diagram of the SLM system,(b)the laser scan strategy used,and(c)optical microscopy image of the original as-extruded alloy and the disk specimen to be laser processed.

where dφis the diffracting plane determined by the test results,d0is the lattice spacing without stress,E is the elastic modulus of the Mg alloys(=45GPa),andνis the Poisson’s ratio(=0.35).Different angles ofψ(?20°,?17.23°,?14°,?9.85°,0°,9.85°,14°,17.23°,20°)were used to determine the slope of the curve.ψis the tilt angle andσφis the internal stress.Each specimen was tested three times.To study the surface chemistry,X-ray photoelectron spectroscopy(XPS;PHI VersaProbe Ⅲ)with an Al-KαX-ray source(1486.8eV)was conducted on the as-processed samples using a take-off angle of 45°Before collecting XPS data,a 0.11kV Ar-ion beam was used to clean the surface,and each specimen was scanned twice.Peak fitting was performed using the MultiPak software after subtracting the Shirley background from the raw data.

2.3.Electrochemical measurements and immersion tests

The corrosion behaviour of the alloys was studied via electrochemical tests in a simulated body fluid(SBF,see Table S1 of the Supplementary file)[35-37].Before testing,only the asextruded specimens were ground using sandpaper to a surface quality similar to that obtained by laser processing,whereas the other samples were directly tested.In the case of H2evolution,the samples were ground from the bottom and top sides.All the samples were cleaned in acetone for 2min in an ultrasound cleaner.An electrochemical workstation(CS310,Wuhan CorrTest Instruments Corp.,Ltd.)with a single-layer plate corrosion pool was used.A saturated calomel electrode(SCE)was used as the reference electrode.The counter electrode was a 20×20mm platinum mesh.The working electrode was a 3 cm2alloy surface exposed to ～300mL of the SBF solution.Potentiodynamic polarization(PDP)was performed at a scanning rate of 1 mV·s?1from?0.5V in the cathodic direction to +0.5V in the anodic direction based on the open circuit potential(OCP).All the tests were repeated three times to guarantee the reliability of the results.H2evolution measurements were also conducted to evaluate the corrosion rate [38].The specimens were placed into an epoxy mould,which only exposed the top,laser-processed surface to be immersed in the SBF solution.All the specimens were tested for 168 h(=7 days)without changing the SBF solution(400mL)for each specimen.The pH value was 7.4 at the beginning of the test.The temperature was maintained at 37°C during the immersion test using an incubator under air atmosphere.

2.4.Microhardness test

The microhardness was measured using a digital micro vickers hardness tester(Model HX-1000TM/LCD).The test load was 100 gf with duration time of 15s.Six points were tested on the surface of each specimen.The average and corresponding standard deviations were calculated.

2.5.Biocompatibility evaluation

To investigate the biocompatibility,MG-63 cells were seeded onto the specimens in 24-well tissue culture polystyrene plates(TCPs)at a density of 5×104cells/specimen.After 2h incubation,each well was washed thrice using phosphate buffer saline(PBS).The cells were fixed with 2.5% glutaraldehyde for 12h at 4°C.The specimens were washed 3 times with PBS for 5min each.The cells were then dehydrated using gradient ethanol solutions(10%,30%,50%,70%,80%,90%,95%,and 100%),for 10min each.The treated samples were dried overnight in a freeze dryer,and sputter coated with platinum.Subsequently,they were observed via SEM.

3.Results

3.1.Surface roughness and laser processing optimization

Fig.2(a)shows the OM images of the specimens after laser processing.Using a fixed scan speed(800mm/s),different laser powers yielded different surface morphologies.The scan speed adopted was predetermined by preliminary experiments(details omitted herein)and will be analysed in the discussion section.The surface roughness was used as a primary indicator to optimize the processing parameters since it can affect the corrosion behaviour.A high roughness value can break the surface protective layer and increase the probability of pitting corrosion [13].Furthermore,a high roughness value should be avoided from a surface engineering perspective [39].

Fig.2.(a)OM images of laser processed alloys(scan rate:800mm/s),(b)and(c)surface roughness of laser-processed Mg-0.6Ca and Mg-0.5Zn-0.3Ca,respectively.

The results are summarized in detail as follows:(a)for the Mg–0.6Ca alloy(Figs.2(a)and 2(b)),the roughness,Rz(the average of absolute values of five highest peaks and five deepest valleys)was 26.7±2 μm at 150W and 800mm/s,and Ra(defined as the contour arithmetic mean deviation)was 3.84±0.9 μm.At 100W and 800mm/s,the Rzvalue was slightly lower,at 26.3±8 μm and Ra=3.1±1.5 μm.Excessive laser energy,for example,280W and 800mm/s,caused a large degree in deformation of the entire specimen;(b)the Mg–0.5Zn–0.3Ca alloy(Figs.2(a)and 2(c))exhibited Rz=22.58±0.5 μm and Ra=3±0.56 μm when processed at 150W and 800mm/s,and Rz=21.41±2.5 μm,Ra=3±0.5 μm at 100W and 800mm/s.It was observed that smoother surfaces were achieved when the laser power was in the range of 100–150W(see Supplementary Fig.S1 for the case of using 100W and 800mm/s).An excessive laser energy input resulted in a larger surface roughness and more fluctuations,which may result in the strong evaporation of the surface material [40].The trends reflected by Rz and Ra are consistent with each other.Combining the Rz and Ra values with the OM images,150W and 800mm/s were determined as a good set of parameters,and they were selected to process samples for further studies on the corrosion behaviour and microhardness.

3.2.Microstructural analysis

XRD analyses were conducted to determine whether any phase transformation occurred due to laser processing.Fig.3 shows that no significant changes occurred in the phase constitution before and after laser processing.For both the alloys,the dominant phase wasα?Mg,whose peak positions shifted slightly towards higher 2θangles after the laser processing.No peaks were associated with intermetallic compounds such as Mg2Ca and Ca2Mg6Zn3[15].Their absence in the laser processed surfaces can be attributed to the rapid solidification associated with SLM,low concentration of the alloying elements,and relatively low detectability of XRD itself [43].

To analyse the grain size and morphology variation before and after SLM processing,EBSD analysis was conducted on the sample surface,which was perpendicular to the laser beam in the laser-processed specimens.The extruded Mg–0.6Ca alloy comprised equiaxed crystals,as depicted in Fig.4(a).The average grain size was 7.6±0.1 μm.After laser processing,the grains elongated,and the average grain size increased to 16.0±0.1μm,Fig.4(b).In comparison,for the Mg–0.5Zn–0.3Ca alloy,Figs.4(c)and 4(d)indicate that its original average grain size was 32.0±0.1 μm,which decreased to 28±0.1 μm.The difference between the two alloys indicates that rapid solidification due to SLM processing can induce a decreased grain size,but the effect is limited and also material dependant.Aside from sample surface,there was an SLM-induced HAZ in the depth direction.Fig.5 shows the typical results for the Mg–0.5Zn–0.3Ca alloy processed at 150W and 800mm/s.A laser modified layer is highlighted in Fig.5(a)by an SEM secondary electron image,whose EBSD result is presented in Fig.5(b);as shown,the depth is approximately 90 μm.The grains were slightly elongated and showed a reduced grain size.A theoretical analysis of the HAZ is provided in the discussion section.

Fig.3.XRD results of the extruded and SLM processed specimens.The standard PDF card for Mg is also provided for comparison.

Fig.4.EBSD results of the extruded and SLM processed specimens:(a)and(b)for Mg-0.6Ca,(c)and(d)for Mg-0.5Zn-0.3Ca.The observations were made on the cross section of the extruded sample.

Fig.5.(a)SEM secondary electron image and(b)EBSD image of the SLM processed Mg-0.5Zn-0.3Ca in the depth direction(i.e.the longitudinal section).

Fig.6.Microhardness of the extruded and SLM processed specimens.

3.3.Microhardness measurements

Fig.6 shows the microhardness of the specimens before and after laser processing.The hardness of the extruded Mg–0.6Ca alloy was 46±1 HV.Laser processing increased the hardness to 56±1 HV,indicating a 20% increase.For the Mg–0.5Zn–0.3Ca alloy,the microhardness increased from 47±3 HV to 55±3 HV after laser processing,i.e.a 17%increase had been achieved.

3.4.Residual stress evaluation

Fig.7 shows the surface residual stress evaluations before and after laser processing.For the Mg–0.6Ca alloy(Figs.7(a)and 7(b)),the residual tensile stress of the extruded sample was 3±4MPa,which increased to 66±27MPa after laser processing.For the Mg–0.5Zn–0.3Ca alloy(Figs.7(c)and 7(d)),the original residual tensile stress was 23±6MPa,which increased to 56±13MPa after laser processing.

3.5.Corrosion resistance evaluation:H2 release in SBF solution

The corrosion rate was first evaluated from an H2release test,see Fig.8 for the SBF immersion results.For the Mg–0.6Ca alloy(Fig.8(a)),the H2gas that evolved from the extruded specimen increased almost linearly with time,indicating a stable corrosion rate.By contrast,the corrosion rate of the specimens after laser processing up to approximately 96 h was much slower.For the Mg–0.5Zn–0.3Ca alloy(Fig.8(b)),only at the beginning of the immersion,the H2amount of the laser-processed specimen was slightly higher than that of the extruded specimen.After ～60h of immersion,the H2gas evolution amount became much lower than that of the extruded sample.

The corrosion rate,PH(mm/y),was evaluated using PH=2.088VH,where VH(mL·cm?2d?1)is the H2evolution rate[41].The results are listed in Table 2.The as-extruded Mg–0.6Ca exhibited a corrosion rate of 2.1±0.2mm/y.After laser processing,the corrosion rate was 1.0±0.1mm/y within the first ～96 h(4 days).The as-extruded Mg–0.5Zn–0.3Ca corroded at a rate of 1.6±0.1mm/y.After laser processing,the corrosion rate was 1.2±0.1mm/y within the first 72 h and then decreased to 0.7±0.2mm/y.Previous reports showed that the corrosion rate of the Mg–Ca alloy was in the range of 1.4–4mm/y [42]and 0.3–2mm/y for Mg–Zn–Ca [43,44].The current results are consistent with the values reported in literature.

Fig.7.Surface residual stress results of the extruded and SLM specimens:(a)and(b)for Mg-0.6Ca,(c)and(d)for Mg-0.5Zn-0.3Ca.

Fig.8.H2 evolution rate from immersion test in the SBF solution:(a)Mg-0.6Ca,(b)Mg-0.5Zn-0.3Ca.

Table 2 Electrochemical parameters and corrosion rates determined for the original and SLM samples.

Fig.9.Potentiodynamic polarization curves of(a)Mg-0.6Ca and(b)Mg-0.5Zn-0.3Ca.

3.6.Corrosion resistance evaluation:potentiodynamic polarization study

The potentiodynamic polarization was also used to evaluate the corrosion behaviour.The corrosion current density was calculated using the Stern–Geary formula [45].For the Mg–0.6Ca alloy(Fig.9(a)),the specimens before and after laser processing exhibited similar electrolyte-specific electrochemical behaviours.The data in Table 2 suggests a higher corrosion potential(Ecorr)for the laser processed specimen.For the Mg–0.5Zn–0.3Ca alloy(Fig.9(b)),the results showed a similar tendencies.In the Table,the values of Icorr can be much influenced by the varying slopes of the data curves recorded.Corrosion resistance of the alloys can be compared by a cross-check between the H2evolution study and the Ecorr values,which show the same tendency,as revealed in Fig.8 and Fig.9.

3.7.Surface chemistry evaluation

XPS was conducted to determine if any variation occurred on the surface due to SLM.For the Mg–0.6Ca alloy,the main elements were Mg and Ca for the as-extruded specimen(Figs.10(a)-10(d)).O was detected due to the reactive nature of Mg;N was not detected by the survey scan or fine scan.The core level binding energies of metallic Mg(2p)is ～49.77eV [46];for the metallic Ca(2p1/2),it is ～348.6eV [47].After the laser processing,the peaks appeared at ～50.8eV and ～349.8eV,respectively,which shifted towards the higher binding energies corresponding to oxidized states,e.g.oxides.This indicates that the surface chemistry changed due to the laser processing in the case of the Mg–0.6Ca alloy,demonstrating increased presence of,in particular,MgO oxides.For Mg–0.5Zn–0.3Ca(Figs.10(e)–10(h)),the chemical states of Mg and Ca did not change significantly before and after laser processing,indicating that the alloy exhibited a better oxidation resistance than Mg–0.6Ca.Zn was not detected,which should be attributed to its low concentration in the alloy.

3.8.Biocompatibility test

Biocompatibility was evaluated for the alloys before and after SLM.The cell adhesion results are shown in Fig.11.

The MG-63 cells exhibited spherical shapes on the surface of the extruded sample,and the filopodia of the cells were clearly visible(Figs.11(a)and 11(c)).In comparison,the MG-63 cells bound tightly to the substrate and extended to a polygonal shape on the laser processed specimen,see Figs.11(b)and 11(d)for Mg–0.6Ca and Mg–0.5Zn–0.3Ca,respectively.Generally speaking,when an implant material is in contact with cells,proteins on the cell membrane quickly adhere to their surface;transmit surface information into biological signals and then passing them back to cells.A good biocompatibility facilitates stable cell adhesion,which was observed in all samples including the laser-processed ones[48-50].

Fig.10.XPS survey and fine spectra for the alloys before the after SLM:(a)-(d)for Mg-0.6Ca,(e)-(h)for Mg-0.5Zn-0.3Ca.

Fig.11.SEM images of cell spread and adhesion on:(a)extruded Mg-0.6Ca,(b)SLM Mg-0.6Ca at 150W and 800mm/s,(c)extruded Mg-0.5Zn-0.3Ca,and(d)SLM Mg-0.5Zn-0.3Ca at 150W and 800mm/s.

4.Discussion

4.1.Theoretical analysis of laser surface modification

The scan speed(i.e.dwell time of the laser beam)and laser power,the two most important processing parameters,were theoretically analysed as follows from the surface modification perspective.

Regarding the laser dwell time,a study has shown that laser heating induced melting occurs if the dwell time is sufficiently long [51]:

where this the heating time,rfis the beam radius at the focal point,andχis the thermal diffusivity.Data for pure Mg were used for the present estimation because of the low alloying concentrations in the Mg alloys studied.Tm(melting temperature)of pure Mg is 923K.The value ofχis 0.874 cm2/s at 27°C [52].The surface temperature is a function of the laser radiation intensity [51]:

where TSis the surface temperature;αis the absorption coefficient,which is 0.26,q is the radiation intensity absorbed by the surface,andκis the thermal conductivity,which is 153.6556 W/m·K.

The HAZ zone where the laser modifies the material is not only limited to the surface,but also to the specimen depth in the Z direction,which can be calculated using Eqs.(3.3)–(3.5)[51]:

where T0is the sample temperature far from the laser spot;x,y and z represent the spatial three-dimensional coordinates of the area affected by the laser.t is the time,is the laser powder,d is the laser beam diameter,v is the scan velocity and Tzis the phase transition temperature.Adding known constants into the equations above can help evaluate the relationship between the laser power,scan speed,and surface state.After incorporating the latent heat of melting(=8.7kJ/mol),it was estimated that the depth of the HAZ was ～59 μm using 100W and 800mm/s,and ～80 μm using 150W and 800mm/s,generally consistent with Fig.4 and Fig.5.

4.2.Laser processing induced improvement in corrosion resistance

The corrosion rate results in Fig.8 and Fig.9 indicate a change in the corrosion resistance of both the alloys.The underlying mechanisms are discussed in terms of the variations in the grain size,surface chemistry and residual stress status due to laser processing.

The Mg–0.6Ca alloy contained a low level of impurity elements namely Fe,Ni,and Cu(Table 1).Moreover,it was revealed that impurity elements,such as Fe,should be kept as low as possible [53],since they can easily deteriorate the corrosion resistance of Mg alloys by forming secondary phases.Hence the high cooling rate associated with SLM processing may be beneficial because under a rapid solidification,both the formation of secondary phases and their growth will be significantly restricted.Furthermore,the results in Fig.10 indicate that the surface chemistry changed because of the laser processing for the Mg–0.6Ca alloy,demonstrating an increased presence of MgO.It is suggested that MgO can increase the self-corrosion potential of Mg alloys[54],implying that it might also be a reason for the improved corrosion resistance.The MgO phase,however,is often not dense and forms a continuous layer on Mg and its alloys [55].Corrosion can still occur with a prolonged immersion time,consistent with the increased corrosion tendency when the immersion time is longer than 96h,as shown in Fig.8(a).This increase in the corrosion rate was not observed in the ternary alloy.However,the solubility of Ca in Mg was lower than that of Zn,and the higher Ca content in the binary alloy promoted the formation of this thermodynamically stable Mg2Ca phase during the first solidification [56,57].This phase formed a weak galvanic coupling with the Mg matrix and can promoted corrosion [15].Both the effects may have already compensated for the negative effect imposed by the high residual stress detected(66±27MPa)in the surface of the laser-processed Mg–0.6Ca.

For the Mg–0.5Zn–0.3Ca alloy,the chemical states of Mg and Ca did not change notably before and after laser processing.Zn element in the alloy’s chemistry may have contributed to the improved oxidation resistance,which is resulted from a lower vaporization temperature of Zn(～1380K vs.1180K of Mg).The generated vapour phase of Zn can hinder the oxidation reaction between environmental oxygen and the Mg matrix[58].This is an evident difference between the two Mg alloys and it helps to explain the opposite,initial H2evolution trends as shown in Fig.8.

Since there was no obvious chemistry change in the surface condition,the corrosion resistance of Mg–0.5Zn–0.3Ca may be a result of several other factors combined,as follows:(a)it is suggested that a refined microstructure generally produces a lower corrosion rate for Mg alloys [59],which is consistent with the decreased grain size due to the SLM of the Mg–0.5Zn–0.3Ca alloy,as shown in Fig.4;(b)similar to the case of the laser-processed Mg–0.6Ca alloy,impurity and alloying elements appeared in the microstructure as a solid solution,which likely contributed to the enhanced corrosion resistance;and(c)residual stress may be a factor as well.The higher level of surface tensile residual stress detected in the laser-processed specimens typically decreases the corrosion resistance by encouraging cracking of protective surface oxides [60].A compressive residual stress corresponding to this surface tensile residual stress appeared which maintained the entire specimen in stress equilibrium.Such a variation in the residual stress may explain the initial slightly higher H2release rate,followed by a slightly decreased corrosion rate during the latter period of the immersion test when the surface residual stress was removed by corrosion resulting in a residual compressive stress.

4.3.Laser processing induced increase on microhardness

The hardness of pure Mg was relatively low,in the range of 32–40 HV [61].Alloying can increase the hardness to 50–53 HV [62].The present study reports a further increase in hardness(e.g.～55 HV for Mg–0.6Ca)owing to SLM,as shown in Fig.6.The increased hardness is expected to provide a higher wear resistance [63].

Regarding the mechanism responsible for the microhardness,the EBSD results in Fig.4 show an increased grain size in the SLM of Mg–0.6Ca but a decreased grain size in the SLM of Mg–0.5Zn–0.3Ca.Changing the hatch space may be a possible way to decrease the grain size.Meanwhile,such elongated grains were observed in other alloys processed through SLM because of the associated thermal gradient [64].The Hall-Petch relationship indicates that a smaller grain size results in an increased hardness/strength [65].This indicates that grain refinement alone may only partially explain the observed increase in the microhardness for the SLM of Mg–0.5Zn–0.3Ca.There must be other contributing factor(s)to the microhardness observed in the laser processed Mg–0.6Ca.Fig.7 revealed that there was residual stress in the laser processed alloys,because of the nonuniform temperature distribution during laser processing.Carlsson and Larsson [66]have found a direct relationship between the hardness and residual stress,suggesting that the residual stress field influences the hardness substantially and that the presence of residual stress can increase the surface hardness [67].Moreover,a large amount of dislocation is often exhibited in the laser processed materials,which can be an atomic-level,stress-induced microstructural change [68].

4.4.General implications of the current study

Fig.12 summarizes the observed phenomena in SLM as a surface engineering approach for modifying Mg alloys,with the Mg–0.5Zn–0.3Ca alloy as a typical example.Applying SLM to the as-extruded Mg alloy changed the original microstructure from equiaxed to elongated grains with decreased grain size.The rapid solidification associated with SLM produced residual stress on the surface of the specimen,which most likely enhanced the microhardness,and cause the Hall-Petch effect due to grain refinement.Negative impacts from existing impurity elements can be confined by the rapid cooling rate during SLM,thereby improving the corrosion resistance.The residual stress may have further enhanced the resistance to H2release in the case of Mg–0.5Zn–0.3Ca.In addition,the biocompatibility remained good after SLM.The results suggest that appropriate laser processing parameters for the as-extruded Mg alloys can simultaneously increase the microhardness and decrease the corrosion rate,indicating the potential of SLM for the development of advanced biomedical Mg alloys.

Fig.12.SLM surface modification of Mg alloy(using Mg-0.5Zn-0.3Ca as a typical example)and important realizable results from the processing.

Conclusion

This study demonstrated that SLM surface modification can be employed for developing biomedical Mg alloys with both low corrosion rate and high microhardness.Mg alloys(Mg–0.6Ca and Mg–0.5Zn–0.3Ca)indicated good biocompatibility after SLM.The major conclusions are as follows:

(1)The biomedical Mg-0.6Ca had simultaneously increased microhardness and decreased corrosion rate due to the surface modification by SLM.The main reason for the enhanced microhardness(20%increment)was attributed to the surface residual stress.The corrosion rate decreased because of confined impurity elements,the presence of MgO and the correspondingly increased selfcorrosion potential.

(2)The biomedical Mg–0.5Zn–0.3Ca alloy exhibited increased microhardness(17% increment).Its H2release amount was slightly higher than the as-extruded state in the initial immersion time period but decreased significantly after ～72h of testing,this was attributed to the combined result of varying residual stress states,refined microstructure and confined impurity elements.

(3)For the biomedical Mg–0.6Ca and Mg–0.5Zn–0.3Ca alloys,the suitable laser processing parameters for surface modification were around 150W as the laser power and 800mm/s as scan speed,according to both experimental observations and theoretical analysis.HAZ could reach a depth of ～100 μm under the present processing settings.

Declaration of Competing Interest

The authors declared that we have no conflicts of interest to this work.We declare that we do not have any commercial or associative interest that represents a conflict of interest in connection with the work submitted.

Acknowledgements

This research was funded by the Shenzhen Science and Technology Innovation Commission(JCYJ20180504165824643)and Shenzhen Industrial and Information Technology Bureau(ZDYBH201900000009).Dr.M.Yan thanks the support of Humboldt Research Fellowship for Experienced Researchers.Professor Dargusch would like to acknowledge the support of the Australian Research Council Research Hub for Advanced Manufacturing of Medical Devices(IH150100024).The authors would like to acknowledge the technical support from SUSTech CRF.