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[Manas Ranjan Sahu](https://orcid.org/0009-0009-6467-8420), [Akiko Yamamoto](https://orcid.org/0000-0002-9182-4886)

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An overview of the recent developments in biodegradable Mg-Zn alloyAvailable online at www.keaipublishing.com/en/journals/journal-of-magnesium-and-alloys/ Journal of Magnesium and Alloys 13 (2025) 486–509 www.elsevier.com/locate/jma Review An overview of the recent developments in biodegradable Mg-Zn alloy Manas Ranjan Sahu ∗, Akiko Yamamoto Research Centre for Macromolecules and Biomaterials, National Institute for Materials Science, 1-1 Namiki, Tsukuba, Ibaraki 305-0044, Japan Received 8 December 2024; received in revised form 10 January 2025; accepted 14 January 2025 Available online 3 February 2025 Abstract The increasing interest in Mg-Zn binary alloys as temporary implant materials is attributed to their outstanding biocompatibility, biodegrad- ability, and favourable mechanical properties. However, their application is constrained by high degradation rates in the physiological envi- ronment, resulting in the release of hydrogen gas and a rapid decline in mechanical properties. Additionally, the material’s biocompatibility is contingent upon its degradability. Researchers have demonstrated that addressing these issues is possible through strategies such as control- ling Zn content, employing thermo-mechanical processing to achieve suitable microstructures, and applying surface coatings. This manuscript provides a comprehensive review of published literature on Mg-Zn alloys, exploring the challenges and outlining future research directions in this field. This is an open access article under the CC BY-NC-ND license ( http://creativecommons.org/licenses/by-nc-nd/4.0/ ) Peer review under responsibility of Chongqing University Keywords: Mg-Zn alloy; Degradation; Temporary implant; Biocompatibility. 1 t  s  t  T  a  d  l  a  m  p  e  [  l  a  p  d  f  s  p t  a  l  p  a  a  (  i  b  h  t  s  m  v  i  h h2r. Introduction Biodegradable metals represent a class of metals designedo be corroded gradually in vivo , with an appropriate host re-ponse elicited by released corrosion products, which can passhrough, be metabolized, or assimilated by cells and tissue.hen, they dissolve completely upon fulfilling the mission tossist tissue healing with no implant residues [ 1 ]. Unlike tra-itional permanent implant materials such as titanium, stain-ess steel, and cobalt-chromium alloys, biodegradable met-ls offer a solution for temporary implants to their com-on complications like stress shielding, bone thickening, im-lant loosening, and chronic inflammation, together with thelimination of secondary removal surgery after tissue healing 2 , 3 ]. The representative temporary implants are cardiovascu-ar stents and fracture fixation devices such as bone platesnd screws. These devices have been commercialized withermanent metallic materials such as a Co-Cr alloy for car-iovascular stents and Ti (grade 2) for microplates and screws,ollowed by the products employing bioabsorbable polymers∗ Corresponding author. E-mail address: SAHU.MANASRANJAN@nims.go.jp (M.R. Sahu) . t  b  a  ttps://doi.org/10.1016/j.jma.2025.01.011 213-9567/This is an open access article under the CC BY-NC-ND license ( http:/esponsibility of Chongqing University uch as poly-L-lactic acids. Table 1 presents the mechanicalroperties of typical implant materials with bone tissue. Among all the available metallic elements in the periodicable, the metals or alloys having basic characteristics suchs complete biodegradability and biocompatibility for each al-oying element with good mechanical properties, are the mostromising candidates to be considered as biodegradable met-ls [ 1 ]. For biodegradability, the metals and/or alloys with standard electrode potential lower than that of hydrogenE0 = 0 V) are considered to be corroded in the physiolog-cal environment by reacting with water in body fluids. Foriocompatibility, the composition of existing elements in theuman body and their threshold concentrations in a healthyissue provide a valuable reference for selecting the compo-ition of biodegradable metals for clinical implant use. Forechanical properties, the biodegradable metals should pro-ide sufficient support (varies as per the implantation site)mmediately after the post implantation until tissue starts toeal. In this context, magnesium (Mg), iron (Fe), zinc (Zn), andheir alloys are the candidates most commonly considered asiodegradable metals for temporary implant applications. Fend its alloys possess outstanding mechanical properties and/creativecommons.org/licenses/by-nc-nd/4.0/ ) Peer review under http://crossmark.crossref.org/dialog/?doi=10.1016/j.jma.2025.01.011&domain=pdfhttp://www.keaipublishing.com/en/journals/journal-of-magnesium-and-alloys/https://doi.org/10.1016/j.jma.2025.01.011http://www.elsevier.com/locate/jmahttp://creativecommons.org/licenses/by-nc-nd/4.0/mailto:SAHU.MANASRANJAN@nims.go.jphttps://doi.org/10.1016/j.jma.2025.01.011http://creativecommons.org/licenses/by-nc-nd/4.0/M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 487 Table 1 The mechanical properties of implant materials with bone tissue. Material Young’s modulus (GPa) Yield strength (MPa) Compression strength (MPa) Tensile strength (MPa) Fracture toughness (MPa.m½) Ref Cortical bone 7 – 30 – 100 – 230 164 – 240 2 – 12 [ 23 , 24 ] Stainless steel 316L 193 170 – 310 480 – 620 540 – 1000 ∼100 [ 23 , 24 ] Ti-6Al-4V 114 760 – 880 – 895 – 930 ∼80 [ 23 , 24 ] Ti (grade 2) 102 275 – 390 – 540 – [ 25–27 ] Co-Cr alloy (ASTM F75) 210 448 – 517 – 655 – 889 – [ 28 ] Co-Cr alloy a (ASTM F562) 232 1500 – 1795 – [ 28 ] PLLA 2.7 – 4.14 – 58.6 15.5 – 150 – [ 29–31 ] Mg 41 – 45.5 51 65 – 100 175 – 235 15 – 40 [ 32–35 ] Zn 90 – 110 285 – 325 – 90 – 200 – [ 32 , 36 ] Fe 204 – 212 108 – 122 – 230 – 345 – [ 32 , 37 , 38 ] a Property of cold worked and aged material. e  l  s  r  l  s  T  e  m  d  t  i  c  m  t  b  t  v  m  p  a  s  n  d   e  a  A  M  t  t  f  i  a  a  p  p  o  i  p  a  a  c  i  r  p  M  e  e  o  t  p  f M  i  i  c  a  v  s2a m  a  i  M  si  b  c  i  t  t  a  txceptional machinability [ 4 ]. However, the main drawbackies in its slow corrosion rate as well as insoluble corro-ion products [ 4 ]. Zn and its alloys have moderate corrosionates and potential bioactivities [ 5 , 6 ]. However, the relativelyow mechanical strength, strain softening behaviour, and highensitivity of Zn ions for bioactivity remain challenges [ 7 ].herefore, Mg has been explored most extensively due to itsxcellent biocompatibility, biodegradation, and comprehensiveechanical properties [ 8–10 ]. The closer elastic modulus andensity of Mg and its alloys (41 - 45 GPa, 1.74 - 2.0 g/cm3 )o those of bone (3 - 20 GPa, 1.8 - 2.1 g/cm3 ) help in reduc-ng the stress shielding effect in orthopaedic implant appli-ation [ 11 ]. In addition, Mg ions play a crucial role in boneetabolism as a beneficial bivalent ion [ 12 ]. They supporthe function of bone osteoclasts and aid in the formation ofiological apatite. However, the rapid degradation of Mg withhe evolution of hydrogen (H2 ) gas in the physiological en-ironment is the major challenge to be used as an implantaterial. The evolved H2 gas leads to the formation of gasockets next to the implant which separate the tissue layernd delay the healing process [ 13 ]. The lower mechanicaltrength of Mg is also a major concern because the implanteeds to be strong enough to maintain its structural integrityuring the degradation before the tissue is adequately healed.To overcome these limitations, Mg based alloys with ad-quate mechanical strength along controlled degradation ratere developed by adding various alloying elements to Mg.mong all the Mg-based alloys, Mg-Al, Mg-Ca, Mg-Mn,g-RE (rare earth), and Mg-Zn alloys grab a lot of atten-ion as potential biodegradable metals [ 8 , 14–16 ]. The addi-ion of Al improves both the mechanical and corrosion per-ormance of Mg-Al alloy, but the suspicion of Al involvementn Alzheimer disease limits its application [ 16 ]. Similarly, theddition of RE elements improves the mechanical propertiesnd corrosion resistance, but the accumulation of RE phos-hates and Nephrogenic Systemic Fibrosis caused by Gd com-ounds are also concerned [ 16 ]. In Mg-Ca alloy, the additionf Ca improves the mechanical properties and biocompatibil-ty. The presence of Ca leads to more intermetallic Mg2 Cahase formation which accelerate the corrosion process andlso hinders the corrosion process by the early deposition ofpatite. Therefore, this dual nature of Ca during corrosion pro-ess resulted a few reported literatures of in vivo data [ 8 ]. Thentermediate electrode potential of Zn in between Mg and Feesulted in the controlled degradation behaviour slower thanure Mg, that led to extensive research and development ofg-Zn alloys [ 17–21 ]. In physiology, Zn is an essential el-ment involved in various biological processes such as genexpression, signal transduction, nucleic acid metabolism, co-rdination of various organic ligand interactions, and apop-osis [ 15 , 22 ]. Therefore, Zn-containing Mg alloys have beenaid more attention and developed as promising candidatesor biomedical applications. In this paper, the research work on the development ofg-Zn alloy-based biomaterials and their advantages and lim-tations have been critically reviewed. The effect of alloy-ng, processing techniques, and surface treatment on the mi-rostructure, mechanical performance, degradation behavior,nd biocompatibility have been discussed. The results of initro and in vivo evaluation of these alloys have also beenummarized. . Processing techniques and microstructure of Mg-Zn lloy The binary Mg-Zn phase diagram ( Fig. 1 ) shows that theaximum solubility of Zn in Mg is 6.2 wt.% ( i.e. 2.5 at.%)t the eutectic temperature of 341 °C [ 39 ]. However, theres a negligible solubility of Mg in Zn at room temperature.g51 Zn20 , Mg21 Zn25 , Mg4 Zn7 , MgZn2 , and Mg2 Zn11 are thetable compounds, and present in the phase diagram. MgZn2 s a well-known non-stoichiometric compound, with a solu-ility limit of around 1 wt.% at around 416 °C and meltsongruently at ∼ 589 °C. The remaining compounds are sto-chiometric in nature. The invariant reactions include one eu-ectoid reaction Mg51 Zn20 � (Mg) + Mg21 Zn25 at 325 °C,wo eutectic reactions L � Mg51 Zn20 + Mg21 Zn25 at 341 °C,nd L � (Zn) + Mg2 Zn11 at 364 °C, four peritectic reac-ions L + (Mg) � Mg51 Zn20 at 341 °C, L + Mg4 Zn7 �488 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 Fig. 1. The binary Mg-Zn phase diagram. Reprinted by permission from Springer Nature: Comment on Mg-Zn (magnesium-zinc), H. Okamoto, JPE 15, 129–130 (1994), https:// doi.org/ 10.1007/ BF02667700 [ 40 ]. M  a2 i  p  Z  p  e  [  s  w  s  T  n  s  w  a  s  i  d  w  w  [  h  a i  t  c  f  l  s  a  [  (  m  c  a2 c  m  x  a  I  ∼  [  w  t  o  n  (  r  c  d  wi  i  e  M  p  g21 Zn25 at 347 °C, L + MgZn2 � Mg2 Zn11 at 381 °C,nd L + MgZn2 � Mg4 Zn7 at 416 °C. .1. Casting As described above, the solubility of Zn in Mg is max-mum as 6.2 wt.% at 341 °C and negligible at room tem-erature [ 39 ]. Therefore, the microstructure of as-cast Mg-n alloy consists of primary α-Mg matrix and secondaryhases precipitating along grain boundaries. However, Ciat al. reported a single-phase microstructure for Mg-1Zn alloy 41 ]. Furthermore, Loftabadi et al. observed only the MgZnecondary phase for Mg-Zn alloy with Zn up to 3 wt.%,hereas the Mg-Zn alloy having Zn more than 6 wt.% hasecondary MgZn phase + Mg51 Zn20 intermetallic phase [ 42 ].he Mg51 Zn20 phase at grain boundaries is attributed to theon-equilibrium solidification. The increased Zn content re-ulted in an increase in the fraction of secondary phaseshich became non-uniformly distributed. In case of Mg-Znlloy above 5 wt.% Zn, the second phase formed a networktructure of dendrite along the grain boundaries (as shownn Fig. 2 ). In Mg-Zn alloy with low Zn content, however,endritic structures with their sizes of 149 μm and 67 μmas observed in Mg-2Zn and Mg-4Zn alloys, respectively,here pure Mg showed dendritic structure sizes of 700 μm 43 ]. Furthermore, Peng et al. observed polygon (pentagon orexagon) petal-shaped secondary dendrites and Mg matrix inll as-cast Mg–xZn ( x = 0.5, 1, 1.5, 2 wt.%) alloys [ 44 ]. Grain size of Mg-Zn alloys decreased with an increasen Zn content up to 5 wt.% [ 41 ] or 6 wt.% [ 45 ], and afterhat, the refinement efficiency of Zn is not significant or de-reased. The segregation of Zn at the front of grain growthorms an intensive constitutional undercooling in a diffusionayer ahead of the advancing solid/liquid interface. The Znegregation restricts the grain growth, promoting the nucle-tion of the primary Mg, and thus, it refines the grain size 46 ]. The growth restriction factor (GRE) is higher for Zn5.31) than Al (4.32) and Y (1.70), indicating that Zn hasore powerful growth restriction and better refinement effi-iency [ 47 ]. The grain sizes of various as-cast Mg-Zn alloysre summarized in Table 2 . .2. Heat treatment Researchers have employed different heat treatment pro-esses to alter the microstructure of the Mg-Zn alloy. Theicrostructure with no second phase were reported for Mg-Zn ( x = 1, 2, 3, 4, 5, and 6 wt.%) alloy after heat treatmentt 380 °C for 10 h in air followed by water quenching [ 48 ].n their study, the grain size was decreased from ∼ 400 to300 μm with increase in the Zn content from 1 to 6 wt.% 48 ], which is attributed to the restriction of the grain growthith high Zn content, as explained in Section 2.1 , promotinghe nucleation of the primary Mg [ 46 ]. A similar observationf no dendrites with more uniform microstructure having aearly equiaxed grain was reported after solution treatmentST) at 400 °C for 10 h, followed by water quenching atoom temperature on Mg-4Zn alloy [ 43 ]. The quenching pro-ess leads to insufficient time for grain growth and resulted inecreased grain size in the specimen. However, α-Mg matrixith secondary phases ( i.e. Mg21 Zn25 , Mg51 Zn20, and MgZn2 ntermetallic) were observed for Mg-12Zn alloy after anneal-ng at 320 °C/20 h with subsequent quenching [ 49 ]. The pres-nce of secondary phase may be attributed to the fact thatg dissolves Zn up to 8 wt.% at this heat treatment tem-erature, and upon quenching the excess Zn was precipitatedhttps://doi.org/10.1007/BF02667700M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 489 Fig. 2. The SEM micrograph showing the microstructures of (a) pure Mg, (b) Mg-1Zn, (c) Mg-5Zn, (d) Mg-7Zn, and EDS analysis corresponding to assigned area A (e) and B (f). Reprinted from Materials Science and Engineering: C, vol 32, Shuhua Cai, Ting Lei, Nianfeng Li, Fangfang Feng, Effects of Zn on microstructure, mechanical properties and corrosion behavior of Mg–Zn alloys, page no. 2570–2577, Copyright 2012, with permission from Elsevier [ 41 ]. a  t  o  s d  f  w  i  t  a  m  l  s  T  3  (  o  s secondary phases. Therefore, the heat treatment tempera-ure should be selected to completely dissolve the Zn contentf the alloy in Mg at that temperature in order to obtain aingle-phase microstructure. Furthermore, grain growth with the disappearance of theendrite structure consisting of α-Mg + MgZn was observedor Mg-3Zn alloy after ST at 310 °C for 24 h followed byater quenching [ 50 ]. This indicates that the prolonged heat-ng period causes the grain growth of the alloy. The solutionreatment followed by quenching is referred as T4 treatmentnd T4 treatment followed by aging is termed as T6 treat-ent. ST (160 °C for 7–102 h) and T6 treated Mg-4Zn al-oys had grain growth as 76% and 94% of as-cast one, re-pectively [ 43 ]. Xian-bin et al. reported a similar trend in4 and T6 Mg-3Zn alloys as their grain sizes of 240 μm and00 μm, respectively, which are larger than that of as-cast one120 μm) [ 51 ]. A homogeneous and decreased amount of sec-ndary MgZn phases were observed in the T4 Mg-3Zn alloy,490 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 Table 2 Summary of grain sizes and the mechanical properties of various as-cast Mg-Zn alloys. Materials Grain size ( μm) Modulus (GPa) Yield strength (MPa) Ultimate tensile strength (MPa) Elongation (%) Compression strength (MPa) Hardness References Pure Mg 350 1.86 29.88 100.47 7.43 183.09 37.10 HB [ 41 ] 700 – – – – – – [ 43 ] Mg-0.5Zn 185.2 – – – – – – [ 42 ] 780 – – – – – – [ 66 ] 600–800 – – – – – – [ 67 ] – – 38 ± 4 95 ± 13 4.2 ± 1.8 – – [ 44 ] Mg-1Zn 280 – – – – – – [ 54 ] 640 – – – – – – [ 66 ] 198.6 ± 28.9 – – – – – – [ 68 ] 100 24.23 60.62 187.73 13.77 329.60 47.33 HB [ 41 ] – – 20 ± 2 101.5 ± 3 6.96 ± 0.5 – – [ 82 ] – – ∼ 28 ∼ 125 ∼ 18 – – [ 88 ] – – 42 ± 3 99 ± 10 6.1 ± 1.9 – – [ 44 ] – – 56 111.6 1.2 – – [ 69 ] Mg-1.5Zn – – 51 ± 3 109 ± 8 5.9 ± 1.8 – – [ 44 ] 250 – – – – – – [ 70 ] Mg-2Zn 149 – – – – – – [ 43 ] 580 – – – – – – [ 66 ] ∼500 – – – – – – [ 71 ] – – 65 ± 3 121 ± 11 5.3 ± 1.9 – – [ 44 ] – – 27 ± 2 145.9 ± 5 12.2 ± 1.5 – – [ 82 ] Mg-2.65Zn 150 – 45 145 12 – – [ 72 ] Mg-3Zn 135.5 – – – – – – [ 42 ] 120 – – – – – – [ 51 ] 490 – – – – – – [ 66 ] 150–200 – – – – – – [ 67 ] – – 47 ± 1.5 167.8 ± 10 13.7 ± 1.0 – – [ 82 ] – – – – – – 48.8 ± 1.8 Hv [ 50 ] 86.7 – 76.5 166.6 9.3 – – [ 69 ] Mg-4Zn 29 – – – – – [ 43 ] – – 58 ± 1.0 216.8 ± 15 15.8 ± 5.5 – – [ 82 ] – – – 156.93 7.38 – – [ 89 ] Mg-5Zn 133.8 – – – – – – [ 65 ] 88.8 ± 7.9 – – – – – – [ 68 ] 55 36.47 75.60 194.59 8.50 334.12 53.80 HB [ 41 ] 129 – 120 212 10 – 52 Hv [ 73 ] – – 68 ± 1.5 185 ± 5 9.2 ± 0.5 – – [ 82 ] 56.9 – 77.6 188.6 7.7 – – [ 69 ] Mg-6Zn – – 69 ± 1.5 182 ± 5 7.2 ± 0.5 – – [ 82 ] Mg-7Zn 70.3 ± 6.4 – – – – – – [ 68 ] 56 39.60 67.28 135.53 6.00 353.11 56.26 HB [ 41 ] w  o  e  f  5  1  M  b  a  t  d  i  p  p  a  T  o  o  s  i  p  a  g  f  c  t  0  p s  b  r  hereas in the T6 Mg-3Zn alloy, more precipitates formedn matrix and grain boundaries as chain structure [ 51 ]. How-ver, no secondary Mgx Zny phase was observed in T4 (450 °Cor 2 h followed by water quench at room temperature) Mg-Zn alloy, whereas T6 (T4 + aging at 230 °C for 4, 6, and0 h followed by cooling in air) Mg-5Zn alloy showed thegx Zny secondary phases uniformly precipitated on the grainoundaries, whose dimensions became smaller than that in thes-received Mg-5Zn alloy [ 52 ]. The quenching process in T4reatment gives insufficient time to precipitate, resulting in theecreased or no secondary phase, whereas the slow coolingn T6 treatment resulted in higher number of secondary phaserecipitates. Additionally, many short bar, white precipitationhases were observed inside the crystal grains of T6 Mg-5Znlloy [ 52 ]. With an increase in aging time from 4 to 10 h for6 Mg-5Zn alloy, there was no change in the characteristicsf second phase precipitates at grain boundaries [ 52 ], becausef grain boundary hinderance for the growth. However, thehort bar, white precipitation phases within the crystal grainsncreases [ 52 ] due to no obstruction during the growth. Therecipitation phases present a filament characteristic with anpproximate length of 2 μm [ 52 ]. The precipitation phasesrow along the same direction in one crystal grain, but dif-erent directions in different crystal grains [ 52 ]. With the in-rease in aging time from 10 h to 144 h for T6 Mg-3Zn alloy,he volume fraction of the precipitation phase increased from.04 ± 0.02% to 3.03 ± 0.17%, with the growth of rod shaperecipitates [ 50 ]. In the case of Mg-xZn ( x = 3, 6 wt.%) alloys with T4 solidolution heat treatment at 340 °C for 6, 12, and 18 h followedy hot water quenching at around 50 °C, the microstructureemained unchanged with an almost similar amount of sec-M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 491 o  3  t  t  t  m  t  t  a m  c  3  s  a  t  o  f  g  m  6  c  a  r  g  p  c  s  t  p  o  h  Z  s  i2 f  c  c  o  l  w  H  g  i  a  t e  e  r  [  m  i  c  t  s  t  w  s  t  t  t  w  a  e  c  m  4  T  p  a t  a  t  t  r  [  t  m  t  t  1  c  f  o  o  w  t  f  a  T  d  F p  s  c  t  s  e  t3 i  t  d  D  ndary Mg12 Zn13 phases in both as-cast and T4 treated Mg-Zn alloys [ 53 ]. However, in T4 Mg-6Zn alloy after 6 h,he number of precipitates significantly changed, indicatinghe decomposition of the secondary phase (Mg51 Zn20 ) intohe matrix and Mg12 Zn13 [ 53 ]. An increase in the heat treat-ent time to 12 and 18 h causes no significant change inhe amount of secondary phases [ 53 ], which can also be at-ributed to the increase in Zn solubility in Mg up to 8 wt.%t the heat treatment temperature. In summary, the literature indicates that the heat treat-ent temperature, time, and cooling method influence the mi-rostructure. The heat treatment temperature should be above40 °C as MgZn precipitates are unstable and would be dis-olved in the matrix at temperatures higher than about 340 °Cccording to the Mg-Zn phase diagram ( Fig. 1 ). The T4reatment is more suitable than the T6 treatment to obtainptimized microstructure with controlled grain size and uni-orm distribution of the secondary phase, because the agingives the continuous secondary phase which may be detri-ental for corrosion properties. The T4 treatment time up to h would be most promising, as beyond 6 h, less signifi-ant changes were observed in the microstructure of Mg-Znlloys. Additionally, prolonged aging time in T6 specimensesulted more dispersion of secondary phases in the crystalrain which grow monotonically with time. The quenchingrocess leads to insufficient time for grain growth and pre-ipitation of secondary phases, resulting in decreased grainize with reduced or no precipitation of secondary phases inhe specimen. The solubility limit at the heat treatment tem-erature along with the cooling process influences the sec-ndary phase in the microstructure. For similar conditions ofeat treatment, the grain size decreases with an increase inn content up to 6 wt.% with no significant effect on theecondary phases, whereas for Zn content > 6 wt.% needednvestigation. .3. Deformation techniques As Mg-Zn alloys have excellent processability, various de-ormation techniques have been applied to modify the mi-rostructure of the Mg-Zn alloy. A microstructure having re-rystallized grain structure comprising equiaxed grains werebserved after rolling on Mg-1Zn [ 54 ] and Mg-3Zn [ 55 ] al-oys. The grain size of the specimens decreases by 82.62%ith an increase in the number of cycles from 0 to 5 [ 55 ].owever, fine homogenous dynamic recrystallization (DRX)rains with intermetallic phase were observed after perform-ng a high strain rate rolling process (HSRR) on Mg-4Znlloy [ 56 ]. The intermetallic phase after HSRR is attributedo the stress-induced precipitation during HSRR. Similar to (relatively low strain rate) rolling, uniformquiaxed grains with no second phase were observed for hotxtruded Mg-Zn alloy with Zn up to 6 wt.%. [ 57 , 58 ]. Grainefinement was not observed with an increase in Zn content 58 , 59 ]. In contrast to this, Peng et al. reported grain refine-ent in backward extruded Mg-Zn alloy with an increasen Zn content from 0.5 to 2 wt.% [ 44 ]. Inhomogeneous mi-rostructure with white precipitates were also observed in ex-ruded Mg-xZn ( x = 2, 3, 4, and 5 wt.%) alloy [ 60 ]. Theecond phases (Mgx Zny ) present a discrete distribution alonghe grain boundaries and their volume fractions gradually riseith increasing Zn concentration from 2 to 5 wt.% [ 60 ]. Theecondary phases, with the help of the pinning effect, preventhe growth of DRXed grains and resulted in a fine microstruc-ure in the deformed sample [ 43 ]. Inhomogeneous microstruc-ure having fine equiaxed grains and row elongated grainsith plenty of strike-like and coarse intermetallic phases werelso observed along the extrusion direction in sintered + hotxtruded Mg-6Zn alloy [ 61 ]. The intermetallic phases becomeoarser with an increase in Zn content [ 61 ]. Hot extrusion andulti-directional forging (MDF) reduced the grain size of Mg-Zn alloy to 80% and 73% of as-cast one, respectively [ 43 ].he volume fractions of the precipitates in extruded and MDFrocessed specimens are 0.4% and 0.6%, respectively, whichre lower than that of the as-cast one (1.3%) [ 43 ]. Severe plastic deformation techniques were also employedo refine the microstructure of Mg-Zn alloy. Equal-channelngular pressing with applied back pressure (ECAP-BP) upo 4 passes on Mg-xZn ( x = 6, 12 wt.%) alloy resulted inhe microstructure consisted of highly-deformed and partially-ecrystallized regions with MgZn2 and Mg21 Zn25 particles 62 ]. The ECAP-BP Mg-12Zn alloy showed a majority ofhe Zn from the supersaturated solid solution of the α-Mgatrix, which is consumed mainly by the MgZn2 nanopar-icles, resulting in the inhomogeneous distribution of Zn inhe α-Mg matrix [ 62 ]. The Mg21 Zn25 microparticles in Mg-2Zn exhibited distinct forms in the α-Mg matrix that wereharacterized as a single-crystalline form, a nano-crystallineorm, and a broken-up form [ 63 ]. Differences in the sizef MgZn2 nanoparticles were reported in the α-Mg matrixf ECAP-BP Mg-12Zn alloy [ 64 ]. The Mg21 Zn25 particlesere observed in a highly deformed α-Mg matrix [ 64 ]. Fric-ion stir processing (FSP) on Mg-5Zn alloy resulted in uni-orm, fine grain of size 1.3 μm, whereas the cast Mg-5Znlloy showed non-uniform, coarse grain of 133.8 μm [ 65 ].he Large precipitates in as-cast specimens were brokenown into refined and uniformly distributed precipitates afterSP. The microstructure of the specimen after the deformationrocess has resulted in grain refinement ( Fig. 3 ). The grainizes of various Mg-Zn alloys resulting from different pro-essing methods are summarized in Table 3 . In some cases,he inhomogeneity of the microstructure of the deformedpecimens has been observed, which is related to the non-qual rate of dynamic recrystallization in different grains, andhus, DRX is found to be incomplete in some grains. . Mechanical properties of Mg-Zn alloy As described earlier in Section 1 , the appropriate mechan-cal properties of the temporary implant are a fundamen-al requirement to maintain its structural integrity during theegradation process before the tissue is adequately healed.ue to the insufficient mechanical properties of pure Mg, al-492 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 Table 3 Summary of grain sizes and the mechanical properties of various deformed Mg-Zn alloys. Materials Processing methods Grain size ( μm) Modulus (GPa) Yield strength (MPa) Ultimate tensile strength (MPa) Elongation (%) Compression strength (MPa) Hardness (Hv) References Pure Mg Hot pressed at 28 MPa for 1–2 h at 535 °C 242.6 – – – – – – [ 83 ] Mg-0.4Zn Hot rolled (at 450 °C) + Annealed 60 – – – – – – [ 74 ] Mg-0.5Zn Back extruded at 420 °C with an extrusion ratio of 12.25 47.5 – 62 ± 1 145 ± 8 17.2 ± 1.3 – – [ 44 ] Mg-1Zn Back extruded at 420 °C with an extrusion ratio of 12.25 37.5 – 91 ± 1 169 ± 9 18.7 ± 1.4 – – [ 44 ] Hot extruded at 300 °C with an extrusion ratio of 25:1 17.9 – 116 254 16.2 – – [ 58 ] hot rolling (at 400 °C) + annealing (at 400 °C for 5 min) 11 – – – – – – [ 54 ] Hot extruded at 400 °C 35.3 [ 75 ] Hot Rolled at 400 °C – – ∼ 160 ∼ 240 ∼ 23 – – [ 88 ] Mg-1.5Zn Back extruded at 420 °C with an extrusion ratio of 12.25 36 – 101 ± 1 190 ± 7 17.2 ± 1.5 – – [ 44 ] hot rolling (at 250 °C) + annealing (at 350 °C for 1 h) 25.82 – – – – – – [ 70 ] hot rolling (at 350 °C) + annealing (at 350 °C for 1 h) 20.85 – – – – – – [ 70 ] hot rolling (at 450 °C) + annealing (at 350 °C for 1 h) 27.1 – – – – – – [ 70 ] Mg-2Zn Back extruded at 420 °C with an extrusion ratio of 12.25 – – 111 ± 1 198 ± 6 15.7 ± 1.6 – 79 [ 44 ] Hot extruded at 300 °C with an extrusion ratio of 25:1 – – 109 258 16.4 – – [ 58 ] Mg-2.4Zn Hot extruded at 210 °C with an extrusion ratio of 18:1 15 – – – – – – [ 94 ] Mg-2.9Zn Hot pressed at 28 MPa for 1–2 h at 535 °C 106.5 – – – – – – [ 83 ] Mg-3Zn Hot extruded at 300 °C with an extrusion ratio of 25:1 17.9 – 98 276 18.5 – – [ 58 ] Mg-3.3Zn Hot pressed at 28 MPa for 1–2 h at 535 °C 155.6 – – – – – – [ 83 ] Mg-4Zn Hot extruded at 300 °C with an extrusion ratio of 25:1 17.6 – 89 297 20.6 – – [ 58 ] Hot extruded at 350 °C with an extrusion ratio of 11:1 – 198.4 ±3.4 301.1 ± 4.1 33.9 – 44.5 ± 3.3 [ 90 ] Hot extruded at 350 °C with an extrusion ratio of 9:1 7.6 ± 2.6 – – – – – – [ 76 ] MDF (after 3 pass) – 142 228 8 – 84 [ 91 ] MDF (after 5 pass) 2.3 – 135 196 8.7 – – [ 91 ] High strain rate rolling (HSRR) 4.5 – – – – – – [ 56 ] Hot pressed at 28 MPa for 1–2 h at 535 °C 70.2 – – – – – – [ 83 ] Mg-4.4Zn Hot pressed at 28 MPa for 1–2 h at 535 °C 112.9 – – – – – – [ 83 ] Mg-5Zn FSP 1.3 – – – – – – [ 65 ] Extruded 33.6 – – – – – – [ 60 ] ( continued on next page ) M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 493 Table 3 ( continued ) Materials Processing methods Grain size ( μm) Modulus (GPa) Yield strength (MPa) Ultimate tensile strength (MPa) Elongation (%) Compression strength (MPa) Hardness (Hv) References Mg-6Zn MDF (after 1 pass) 72 ± 4 – 130 ± 6 222 ± 10 8.3 ± 1 – 78 ± 6 [ 84 ] MDF (after 3 pass) 28 ± 2 – 145 ± 6 230 ± 10 6.2 ± 0.7 – 86 ± 6 [ 84 ] MDF (after 5 pass) 3.8 ± 0.5 – 138 ± 7 193 ± 9 7 ± 0.9 – 82 ± 5 [ 84 ] Hot extruded at 250 °C 42.3 ± 0.1 169.5 ±3.6 279.5 ± 2.3 18.8 ± 0.8 433.7 ± 1.4 – [ 57 ] Rolling – 91.3 ± 40 124.1 ± 37.9 0.5 ± 0.3 – 48.32 ± 1.5 [ 92 ] l  p  t  [  i  o  a  a  a  p  t  t  c  t  o  5  f  g  s  i  c  r  m  t  h  i  s i  i  e  M  S  t d  Ffoying and microstructure control are major solutions to im-rove them. For Mg-Zn alloys, their hardness was reportedo increase with an increase in Zn content (shown in Fig. 4 ) 42 , 44 , 45 , 79 ]. In Mg-Zn alloy with Zn content < 5 wt.%, thencreased hardness is due to the solid solution strengtheningf Zn dissolved in the Mg matrix [ 42 ]. However, in Mg-Znlloy with Zn content > 5 wt.%, the hardness enhancement isttributed to the evolution of the secondary phase (Mg51 Zn20 )t the grain boundary in addition to the MgZn intermetallichase [ 80 ]. The increased hardness with increased Zn con-ent in Mg-Zn alloy indicate the probability of decrease inhe wear of the implant [ 81 ]. Furthermore, the tensile andompressive strength was improved up to 5 wt.% Zn dueo grain strengthening, solid solution strengthening, and sec-nd phase strengthening [ 41 ]. The decreased strength above wt.% Zn is attributed to plenty of secondary phases whichormed a network structure with dendritic segregation alongrain boundaries, resulting in residual defects and deterioratedig. 3. The comparison of the grain size distribution for various as-cast and derom [ 41–43 , 51 , 54 , 65–73 ], and deformed from [ 43 , 44 , 56 , 58 , 60 , 70 , 74–78 ]. trength and elongation of the specimen. The Zn-rich regionsn Mg-Zn alloy were also often prone to the formation of mi-roporosity, resulting in reduced mechanical properties. Someesearchers [ 43 , 82 , 83 ] reported a similar trend of improvedechanical properties but up to 4 wt.% Zn, and over that,hey declined. Therefore, the Mg-Zn alloy with 4 wt.% Znaving highest strength and hardness is more suitable as anmplant material to prevent fractures and improve functionaltability [ 81 ]. The alteration in microstructure due to heat treatment alsonfluences the mechanical properties. For Mg-4Zn alloy, anncrease in aging time from 0 to 62 h enhances the Vick-rs hardness from 27 to 41 due to the precipitation of moregZn2 phase [ 43 ]. However, decreased USS was reported forT Mg-4Zn alloy due to the grain growth and dissolution ofhe precipitates [ 43 ]. The mechanical properties were further improved after theeformation process due to grain refinement ( Fig. 5 ). Forformed Mg-Zn samples. The data sources are following references: as-cast 494 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 Fig. 4. The effect of different Zn additions on microhardness properties of Mg-xZn alloy. Reprinted by permission from Springer Nature: Rare Metals, Thermal characteristics and corrosion behaviour of Mg–xZn alloys for biomedical applications, LOTFABADI A.F., IDRIS M.H., OURDJINI A., KADIR A., FARAHANY S., and BAKHSHESHI-RAD H.R., Bull Mater Sci 36, 1103–1113 (2013), https:// doi.org/ 10.1007/ s12034–013–0566–9 , Copyright 2013 [ 42 ]. Fig. 5. The comparison of (a) UTS, and (b) elongation of various Mg-Zn alloys in as-cast and deformed conditions. The data sources are following references: as-cast from [ 41 , 44 , 69 , 73 , 82 , 88 , 89 ], and deformed from [ 44 , 57 , 58 , 84 , 88 , 90–92 ]. M  a  t  t  d e  d  d  i  w  o  d  w  e  r  a c  t  g-4Zn alloys, extrusion and MDF improved USS by 26%nd 18%, respectively, than that of the as-cast one, due toheir fine microstructure [ 43 ]. The enhanced properties are at-ributed to higher strain, grain refinement strengthening, andispersion strengthening mechanisms. In case of Mg-6Zn alloy, the hardness and strength werenhanced up to 3 passes of the MDF, and over 4 passes, theyecreased due to dynamic recrystallization resulting in twinensity reduction [ 84 ]. For the hot extruded specimens, anncrease in Zn content from 1 to 4 wt.% decreased YS alongith increased UTS and elongation [ 58 ]. This enhancementf elongation and decrease of YS might be attributed to theecrease in basal texture intensity with increase in Zn content,hereas the increase of UTS is considered to relate with thenhanced strain hardening ability [ 85 , 86 ]. Researchers alsoeported improved mechanical properties after ECAP-BP [ 62 ],nd hot rolling [ 55 , 87 ] on Mg-Zn alloys. Degradable implant materials have deterioration in the me-hanical properties during their degradation process. Tensileest in 0.9 wt.% NaCl solution revealed about 19% loss inhttps://doi.org/10.1007/s12034-013-0566-9M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 495 s  m  T  M  b  l  T  n t  o  T  e  i  t  m  g  a  T  p  s  M4 t  i  d  t4 r  i  (  t  r  o  r  i  t  r  p  l  m  t  a  b  l  p  g Z  Z  Z  w  l  l  c  d  i  a  l  c  t  o  s  tA  CZM i  h  b  s  v  s  T  n  b  t  s  c  d  c  cr  s o  r  i  t  7  o  T  a  e  t  e  trength for the HSRR Mg-4Zn specimen after 7 days of im-ersion, while that of the as-cast sample is about 62% [ 56 ].he tensile test in simulated body fluid (SBF) reported thatg-4Zn alloy is susceptible to stress corrosion cracking at orelow strain rates of 3.6 × 10–4 s-1 with approximately 10%oss in fracture stress as compared to that tested in air [ 93 ].he Mg-4Zn alloy is found to crack at corrosion pits domi-antly by anodic dissolution and hydrogen embrittlement. In summary, the above literature indicates that the Zn con-ent has to be controlled to promote the mechanical propertiesf Mg-Zn alloys more suitable for human implant materials.he Zn content up to 4 wt.% enhances the mechanical prop-rties due to grain strengthening, solid solution strengthen-ng, and second-phase strengthening mechanisms. The heatreatment temperature, time, and cooling process affect theechanical properties. Similarly, the strain rate, the extent ofrain refinement, and the dispersion of the secondary phasefter the deformation process alter the mechanical properties.able 2 and 3 show the grain size and various mechanicalroperties, tensile strength, ultimate tensile strength, compres-ive strength, and elongation, of different as-cast or deformedg-Zn alloys, respectively. . Degradation properties of Mg-Zn alloy In the human body, a biodegradable implant is expectedo degrade at a suitable rate matching to the recovery rate ofnjured tissues. Hence, it is essential to understand its degra-ation process and underlying mechanisms, together with theechniques to regulate degradation behaviour. .1. Degradation mechanism of Mg-Zn alloy It is well known that the microstructure influences the cor-osion behaviour of Mg and its alloys. The corrosion processnitiated on the α-Mg matrix phase includes reactions (1) and2) [ 95 , 96 ]. In Mg-Zn alloy, the lower standard electrode po-ential of Mg (−2.36 V) than Zn (−0.76 V) resulted in moreelease of Mg2 + than Zn2 + in an aqueous medium. The an-dic dissolution of Mg released Mg2 + , whereas the cathodiceaction evolved hydrogen gas (H2 ) and released hydroxideons (OH−) resulting in an increase in local pH. Accordingo the Pourbaix diagram, the released Mg2 + reacts with OH−,esulting in magnesium hydroxide [Mg(OH)2 ] formation atH > 11.5. However, after surgery, the pH is nearly 7.4 orower at the interface of bone and implant due to circulation,etabolic, and resorptive processes [ 97 ]. As a result, a par-ially protective, unstable Mg(OH)2 layer was formed on thelloy surface. With the progress of time, the Mg(OH)2 layerecomes denser which inhibits the ion penetration through theayer, retarding the growth of it. Therefore, the dissolution-recipitation mechanism plays the predominant role in therowth of Mg(OH)2 layer [ 98 ]. The released Zn2 + can also react with OH− to formn(OH)2 [ 99–101 ]. The solubility product constant ofn(OH)2 is much smaller than that of Mg(OH)2 [ 102 ] andn(OH)2 is converted to more stable ZnO (reaction (5)),hich can enhance the protectiveness of the insoluble saltayer (ISL) [ 59 , 101 ]. The chloride ion (Cl- ) in the physio-ogical fluids reacts with the deposited Mg(OH)2 and disso-iates it by forming MgCl2 (reaction (4)). The smaller ra-ius of Cl- enables to penetrate through ISL which causests preferential adsorption by replacement of OH− [ 61 ]. As result, Cl- shifts the dynamic balance between the disso-ution and formation of ISL, resulting in acceleration of theorrosion rate. The presence of Ca2 + , CO3 2- and PO4 3- inhe physiological fluids results in the predominant depositionf calcium/magnesium phosphate/carbonate on the specimenurface due to their lower solubilities, which further retardshe anodic dissolution of Mg. nodi c r eact i on : Mg ( s) → M g2+ ( aq ) + 2e− (1)athodi c r eact i on : 2H2 O + 2e− → H2 ( g) ↑ + 2OH− ( aq ) (2) n ( s) → Z n2+ ( aq ) + 2e− (3) g ( OH ) 2 ( s) + 2Cl− ( aq ) → MgC l2 ( s) + 2OH− ( aq ) (4) Zn ( OH ) 2 ( s) → ZnO + H2 O (5) The selection of in vitro physiological solution plays anmportant role in evaluating the degradation behaviour. In theuman body, interstitial fluid has a similar composition tolood plasma; plasma penetrates through the capillary ves-el wall to be interstitial fluid, which comes back to plasmaia lymph flow. Hence, the in vitro physiological solutionhould have a similar composition to that of blood plasma.he physiological solution should contain essential compo-ents such as inorganic ingredients, organic components, anduffering system, since this influence severely the degrada-ion behaviour of biodegradable metals [ 14 ]. Table 4 repre-ents the list of common physiological solution and their ioniconcentrations along with that of human blood plasma. Asescribed above, the Cl- and SO4 2- contributes to the disso-iation of Mg(OH)2 layer, accelerating the degradation pro-ess [ 103 ] whereas HPO4 2- /PO4 3- , HCO3 - /CO3 2- , and Ca2 + etard the degradation process by the precipitation of theiralts [ 103 , 104 ]. The deposition of the ISL on the specimen surface dependsn the local pH of the corrosion environment. Since OH- iseleased accompanying to Mg anodic dissolution, the buffer-ng ability of the physiological solution severely influenceshe ISL deposition. The pH of the blood is maintained as.4 mainly by carbonate buffer system under the atmospheref 5% CO2 , which is higher than that in air (0.4%) [ 105 ].herefore, the employment of carbonate buffer is necessarys the main buffer system in the physiological solution forvaluating the corrosion behaviour of Mg alloys. Because ofhis reason, phosphate buffered saline (PBS) is not preferableven though its pH was around 7.4. SBF and HBSS contains496 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 Table 4 The main components of commonly used simulated physiological solutions in comparison with human blood plasma. Medium Na+ (mM/L) K+ (mM/L) Ca2 + (mM/L) Mg2 + (mM/L) Cl- (mM) HCO3 - (mM/L) H2 PO4 - /HPO4 2- /PO4 3- (mM/L) SO4 2- (mM/L) Amino acids (g/L) Vitamins (g/L) Proteins (g/L) Glucose (g/L) Phenol red (mM/L) Ref Blood plasma a 142.0 5.0 2.5 1.5 103.0 27.0 1.0 0.5 0.25−0.4 unknown 63−80 0.65−1.1 –[104 [ 108–111 ] PBS 154 4.1 ––140.6 –9.5 ––––––[ 112 ] SBF 142 5.0 2.5 1.5 148.8 4.2 1.0 0.5 –––––[ 108 ] Revised SBF b 142 5.0 2.5 1.5 103 27 1.0 0.5 –––––[ 108 ] HBSS 143 5.8 1.26 0.90 146.7 4.2 0.78 0.4 –––1.0 0.03 [ 113–115 ] EBSS b 144.4 5.3 1.8 0.8 126.2 26.2 1.0 0.8 –––1.0 0.03 [ 116 ] MEM b 143 5.4 1.8 0.8 125 26.2 1.0 0.8 0.87 0.008 –1.0 0.03 [ 110 , 118 ] EMEM b 144 5.4 1.8 0.8 125 26.2 0.9 0.8 0.86 0.009 –1.0 0.02 [ 104 ] DMEM b 157 5.3 1.8 0.8 121.1 44.0 0.9 0.8 1.61 0.032 –4.5 0.04 [ 119 , 120 ] α-MEM b 144.8 5.3 1.8 0.8 128.2 26.2 1.0 0.8 1.27 0.060 –1.0 0.03 [ 117 ] The compositions of the simulated body fluids (except SBFs) are based on the basic version of the commercially available products. The catalogue number of a representative one is shown in the brackets for reference. PBS: Dulbecco’s phosphate buffered saline (ThermoFisher Scientific 14190), SBF: simulated body fluid, HBSS: Hanks’ balanced salt solution (ThermoFisher Scientific 24020); EBSS: Earle’s balanced salt solution (ThermoFisher Scientific 24010); MEM: Minimum Essential Medium (ThermoFisher Scientific 11095), EMEM: Eagle’s Minimum Essential Medium (AccuDiaTM Eagle’s MEM1 ©); DMEM: Dulbecco’s Modified Eagle’s Medium (ThermoFisher Scientific 11965), α-MEM: Minimum Essential Medium alpha (ThermoFisher Scientific 12561). a Being equilibrated with 5% CO2 . b Should be used under 5% CO2 atmosphere. b  p  t  p  T  l  s t  [  s  [  s  t  d  E  p  c  a  s  a  ic  i4a o  d  c  [  c  p  s  h  s  a  d  s  d  c  r  l  t  l  T  [  0  a (  c  q  i  t  icarbonate, but its concentration is much lower than that inlasma for their use in air, not under 5% CO2 , suggestingheir inadequacy for the use as corrosion testing solution forreclinical evaluation. The simple buffering system such asris–HCl, maintains the pH of the corrosion testing solutionow, inhibits the ISL formation, and accelerates the Mg dis-olution [ 106 ]. Protein retards the Mg degradation in the initial stage andhe effect weaken dramatically with the prolonged period 106 ]. Organic molecules like amino acids reduce ISL on thepecimen surface, resulting in acceleration of Mg corrosion 104 ]. Cell culture media supplemented with protein mixtureuch as serum along with 5% CO2 condition are regarded ashe most suitable physiological solution for evaluating the Mgegradation behaviour [ 107 ]. Among them, EMEM or MEM-arle is preferable since it was developed based on the com-osition of human blood plasma, that is, it has the closestomposition. DMEM, one of the modified MEMs, containslmost double amount of NaHCO3 than that of blood plasma,uggesting the acceleration of carbonate precipitation in ISLs well as its buffering ability exceeding that in the humannterstitial fluid. The appropriateness of this high NaHCO3 oncentration of DMEM should be studied in comparison ton vivo or clinical data. .2. Effect of amount of Zn on the degradation of Mg-Zn lloy In Mg-Zn alloys, the concentrations of Zn and the naturef the secondary phase have a significant role in their degra-ation behaviour. The degradation rate decreases with an in-rease in Zn content up to 5 wt.%, but over that, it increases 41–43 , 45 , 53 ]. The increase of Zn content over 5 wt.% in-reased continuous network structure of intermetallic MgZnhase in the matrix, leading to the formation of anode-cathodeites and resulting galvanic corrosion. Mg-Zn alloys alwaysave a lower degradation rate than pure Mg [ 41 , 43 , 121 ]. Theurface of pure Mg (as shown in Fig. 6 ) collapsed severelynd showed lamellar microstructures in SBF (a corrosion me-ia), indicating a substantial corrosion rate during the immer-ion test [ 41 ]. The surface of Mg-Zn alloys showed severaleep pits (indicated by arrows) superimposed on superficialorrosion over the specimen surfaces, implying localized cor-osion attacks during the immersion process in SBF [ 41 ]. Theocalized corrosion for Mg-Zn alloy was also confirmed fromhe breakdown potential (Eb ) (as shown in Fig. 7 ) on the po-arization curve obtained by the electrochemical test in SBF.he more positive Eb indicates less risk of localized corrosion 122 ], where Mg-5Zn alloy showed slightly higher Eb ( = -.96 V in Fig. 7 ) than those of Mg-1Zn alloy (Eb = −1.09 V)nd Mg-7Zn alloy (Eb = −1.25 V). Additionally, the electrochemical impedance spectroscopyEIS) of Mg-Zn alloys suggests the involvement of two-timeonstants at the high-frequency region (HF) and the low fre-uency region (LF) (as shown in Fig. 8 ) [ 41 ], suggesting thenterference of ISL formed on the Mg-Zn alloy surface duringhe immersion in SBF. The capacitive loop in the high fre-M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 497 Fig. 6. The SEM microstructure of corroded appearance after corrosion product removal for (a) pure Mg, (b) Mg–1Zn, (c) Mg–5Zn and (d) Mg–7Zn immersed in SBF for 5 days. Reprinted from Materials Science and Engineering: C, vol 32, Shuhua Cai, Ting Lei, Nianfeng Li, Fangfang Feng, Effects of Zn on microstructure, mechanical properties and corrosion behavior of Mg–Zn alloys, page no. 2570–2577, Copyright 2012, with permission from Elsevier [ 41 ]. q  a  b  a  q  s  s  t  c  p4a c  d  d  t  c  s  i  M  M  f  s  m  p  w  a  d  m  /  4  b  p  uency region can be attributed to the charge transfer reactiont the metal surface and the capacitance of the electric dou-le layer formed at the interface between the metal surfacend the corrosive medium. The capacitive loop in the low fre-uency region can be attributed to the mass transport in theolid phase, such as the diffusion of ions through the insolublealt layer. The EIS analysis also confirms the lower degrada-ion rate of Mg-Zn alloys than pure Mg. Table 5 shows theorrosion rate of as-cast pure Mg and Mg-Zn alloys in varioushysiological solutions. .3. Effect of heat treatment on the degradation of Mg-Zn lloy It has been reported that heat treatment alters the mi-rostructures which play a pivotal role in tailoring the degra-ation rate of Mg-Zn alloys. T4 Mg-6Zn alloy has a higher re-uction rate in corrosion than T4 Mg-3Zn alloy in comparisono those of corresponding as-cast alloys [ 53 ]. Although the as-ast Mg-6Zn alloy (having Mg12 Zn13 and Mg51 Zn20 phases)howed higher corrosion rate than as-cast Mg-3Zn alloy (hav-ng Mg12 Zn13 phases), the decomposition of Mg51 Zn20 in T4g-6Zn alloy resulted in the lower corrosion rate than the T4g-3Zn alloy [ 53 ]. The increased heating time in T4 processrom 6 to 18 h has a less significant influence on the corro-ion behaviour on all the Mg-Zn alloys due to the unchangedicrostructure after heat treatment beyond 6 h [ 53 ]. The im-roving effect of solution treatment on corrosion resistanceas also confirmed by other studies. ST Mg-3Zn alloy gave lower corrosion rate (3.05 ± 0.20 mL /cm2 / day) due to theissolution of ( α-Mg + MgZn) eutectic phases, which reducesicrogalvanic corrosion than as-cast one (3.50 ± 0.20 mLcm2 / day) [ 50 ]. Comparing to as-cast Mg-4Zn alloys, ST Mg-Zn alloy has a 37% lower hydrogen evolution rate (HER),ecause the ST treatment decreases the volume fraction ofrecipitation phases in Mg-4Zn alloys, which reduces the mi-498 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 Fig. 7. The potentiodynamic polarization curves of pure Mg and Mg-Zn alloys in SBF. Reprinted from Materials Science and Engineering: C, vol 32, Shuhua Cai, Ting Lei, Nianfeng Li, Fangfang Feng, Effects of Zn on microstructure, mechanical properties and corrosion behavior of Mg–Zn alloys, page no. 2570–2577, Copyright 2012, with permission from Elsevier [ 41 ]. Table 5 The list of corrosion data of as-cast various Mg-Zn alloys. Materials Electrolyte Ecorr (VSCE ) Icorr ( μA/cm2 ) Electrochemical corrosion rate (mm/y) Immersion corrosion rate (mm/y) Hydrogen evolution rate (ml/cm2 /day) Ref Pure Mg PBS −1.80 14.6 0.33 – 3.44 (after 11 days) [ 43 ] Pure Mg SBF −1.581 680.1 15.30 34.78 (after 5 days) – [ 41 ] Pure Mg SBF −1.886 86.06 1.94 – – [ 88 ] Pure Mg HBSS −1.533 15.98 0.36 – – [ 88 ] Mg-0.5Zn SBF −1.840 144 2.1 5.19 (after 6 days) [ 42 ] Mg-1Zn SBF −1.527 23.4 0.53 2.01 (after 5 days) – [ 41 ] Mg-1Zn SBF −1.822 67.30 1.52 – – [ 88 ] Mg-1Zn HBSS −1.609 10.47 0.24 – – [ 88 ] Mg-1Zn Ringer’s soln. −1.432 300.18 54.30 162.5 (after 4 days) 213.7 (after 4 days) [ 69 ] Mg-1.25Zn SBF −1.762 282.7 6.45 3.17 (after 6 days) – [ 123 ] Mg-1.5Zn SBF −1.740 134 1.67 2.91 (after 6 days) [ 42 ] Mg-2Zn PBS −1.75 8.9 0.20 – 3.37 (after 11 days) [ 43 ] Mg-2.5Zn SBF −1.744 242.5 5.54 – – [ 123 ] Mg-3Zn SBF −1.731 228.3 5.21 1.95 (after 7 days) – [ 53 ] Mg-3Zn SBF −1.701 103 1.55 1.84 (after 6 days) – [ 42 ] Mg-3Zn SBF −1.521 12.6 – – – [ 124 ] Mg-3Zn Ringer’s soln. −1.621 3.1 0.57 2.03 (after 4 days) 0.65 (after 4 days) [ 69 ] Mg-4Zn PBS −1.73 7.8 0.18 – 2.73 (after 11 days) [ 43 ] Mg-4Zn SBF −1.710 212.4 4.85 2.14 (after 6 days) – [ 123 ] Mg-5Zn SBF −1.477 11.72 0.26 1.26 (after 5 days) – [ 41 ] Mg-5Zn HBSS −1.56 33.80 – – – [ 73 ] Mg-5Zn Ringer’s soln. −1.591 3.03 0.55 8.04 (after 4 days) 5.02 (after 4 days) [ 69 ] Mg-6Zn SBF −1.759 122 1.83 7.23 (after 6 days) – [ 42 ] Mg-6Zn SBF −1.759 270.8 6.18 3.48 (after 7 days) – [ 53 ] Mg-7Zn SBF −1.543 51.79 1.17 3.18 (after 5 days) – [ 41 ] Mg-9Zn SBF −0.856 147 2.21 11.78 (after 6 days) – [ 42 ] M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 499 Fig. 8. (a) Nyquist plots for pure Mg and Mg-Zn alloys after soaking in SBF for 1 h and (b) equivalent circuit used for modelling experimental EIS data of Mg-Zn alloys. Rs indicates the corrosive media resistance, Rt indicates the charge transfer resistance, Rf indicates the insoluble layer resistance, CPE1 indicates the double layer capacitance, and CPE2 indicates the insoluble salt layer capacitance. Reprinted from Materials Science and Engineering: C, vol 32, Shuhua Cai, Ting Lei, Nianfeng Li, Fangfang Feng, Effects of Zn on microstructure, mechanical properties and corrosion behavior of Mg–Zn alloys, page no. 2570–2577, Copyright 2012, with permission from Elsevier [ 41 ]. c  [  a  m  p  [  t  t  p  e  r  T  Z o  o  h  l  m  a  t  l  d4M m  o  a  d  t  c  d  s  t  s  c  a  s  c  [  p  t  t  b  6  T  t  c  h  l  t  a  f  b  d  d  c  [  k  c  t  a  t  d  t  t  d  s  s  d  a  f  e  T  Z  t  i  drogalvanic corrosion, leading to a decreased degradation rate 43 ]. However, T6 Mg-4Zn alloy has 30% higher HER thans-cast one due to the dispersion of more MgZn2 precipitatesainly at the grain boundaries where the corrosion processreferentially begins and continues from the grain boundaries 43 ]. The corrosion rate increases monotonically with agingime during T6 treatment, indicating that an increase in agingime leads to an increase in the volume fraction of nano-scalerecipitates in Mg-Zn alloy, which results to micro-cathodicffects and increases the degradation rate [ 50 , 52 ]. Similaresults were also reported by other researchers [ 51 , 52 , 125 ].able 6 shows the corrosion rate of various heat-treated Mg-n alloys in various physiological solutions. In summary, the heat treatment has significant influencen the corrosion behaviour of Mg-Zn alloy with Zn contentf more than 4 wt.%. The T6 treated specimen showed aigher corrosion rate than that of T4 treated one due to thearger volume fraction of Mgx Zny secondary phases. Besides,ore Mgx Zny phases dispersed at the grain boundaries whichre preferential site for the corrosion initiation. Therefore, T4reatment up to 6 h of heat treatment is more preferable toower the corrosion rate in Mg-Zn alloy as the microstructureid not change after heat treatment beyond 6 h. .4. Effect of deformation techniques on the degradation of g-Zn alloy The deformation process leads to much smaller grain sizes,ore uniform microstructures, and a lower volume fractionf the precipitates in the deformed samples, which results in lower degradation rate. Gu et al. observed a lower degra-ation rate for the hot rolled pure Mg and Mg-1Zn alloyhan the corresponding as-cast ones due to the finer grain mi-rostructure [ 88 ]. The finer microstructure resulted in a higherensity of grain boundaries in the deformed samples compen-ating the volumetric mismatch in an atomic scale betweenhe Mg matrix and its oxide, making the oxide layer moretable. Additionally, grain boundaries provide excellent nu-leation sites for the formation of oxide layer, along withcting as physical corrosion (diffusion) barriers against corro-ion propagation. Furthermore, an increase in the rolling cy-le resulted in an increased degradation rate for Mg-3Zn alloy 55 ]. Zou et al. observed uniform corrosion with small depthits for the HSRR Mg-4Zn sample in 0.9% NaCl, whereashe interconnected deeper pits with non-uniform corrosion forhe as-cast sample [ 56 ]. Similarly, Mg-4Zn samples deformedy MDF or extrusion showed decreased degradation rates by3% and 74% respectively than that of the as-cast one [ 43 ].he higher degradation rate in the MDF-processed samplehan the extruded one is attributed to the inhomogeneous mi-rostructure consisting of some un-recrystallized grains withigh dislocation density and high energy together with someow-energy DRXed grains in the MDF specimen. The recrys-allized grains act as anode and the un-recrystallized grainss cathodic sites, resulting in microgalvanic corrosion. There-ore, the density of dislocations can influence the corrosionehaviour of the material. Generally, the regions with a higherislocation density act as anode, whereas those with a lowerislocation density act as cathode. This might result in the mi-rogalvanic effect increasing the corrosion rates in Mg alloys 126 ]. MDF-processed Mg-6Zn alloy has decreased cathodicinetics and growth of oxide layer along with refined mi-rostructure, which resulted in a lower degradation rate thanhe homogenised sample [ 84 ]. Furthermore, the electrode re-ction kinetics of the alloy decreased with the increase in theotal length of the grain boundaries [ 127 ]. Similarly, a reducedegradation rate with uniform corrosion was also observed af-er FSPed [ 65 ] and Hot Extrusion + ECAP-stimulated solu-ion treatment [ 77 ] of Mg-Zn samples. However, an increasedegradation rate for the Mg-Zn specimen was observed afterome deformation process such as on ECAP-BP [ 62 ], and ten-ile and compressive deformation [ 72 ]. The increase in degra-ation rate by some deformation process may be attributed toccumulation of structural defects ( e.g. dislocations and de-ormation twins), increased residual stress, which leads to thenhancement of intergranular corrosion and stress corrosion.able 7 shows the corrosion rate of various deformed Mg-n alloys in various physiological solutions. Fig. 9 indicateshe corrosion rates of Mg-Zn alloys in simulated body flu-ds, revealing that Mg-5 wt.% Zn alloys showed the lowestegradation rate among other as-cast alloys. 500 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 Table 6 The list of corrosion data of various heat-treated Mg-Zn alloys. Materials Processing methods Medium Ecorr (VSCE ) Icorr ( μA/cm2 ) Electro-chemical corrosion rate (mm/y) Immersion corrosion rate (mm/y) Hydrogen evolution rate (ml/cm2 /day) Ref Mg-3Zn T4 for 6 h SBF −1.730 223.6 5.10 1.92 (after 7 days) – [ 53 ] Mg-3Zn T4 for 12 h SBF −1.724 216.2 4.94 1.87 (after 7 days) – [ 53 ] Mg-3Zn T4 for 18 h SBF −1.718 210.7 4.81 1.85 (after 7 days) – [ 53 ] Mg-4Zn ST PBS −1.71 5.8 0.13 – 1.71 (after 11 days) [ 43 ] Mg-4Zn T6 PBS −1.69 50.0 1.14 – 3.54 (after 11 days) [ 43 ] Mg-6Zn T4 for 6 h SBF −1.702 205.2 4.68 1.42 (after 7 days) – [ 53 ] Mg-6Zn T4 for 12 h SBF −1.692 198.4 4.53 1.38 (after 7 days) – [ 53 ] Mg-6Zn T4 for 18 h SBF −1.682 191.5 4.37 1.36 (after 7 days) – [ 53 ] Table 7 The list of corrosion data of various deformed Mg-Zn alloys. Materials Processing methods Medium Ecorr (VSCE ) Icorr ( μA/cm2 ) Electro- chemical corrosion rate (mm/y) Immersion corrosion rate (mm/y) Hydrogen evolution rate (ml/cm2 /day) Ref Pure Mg Hot Rolled at 400 °C SBF −1.796 37.24 0.84 – – [ 88 ] Pure Mg Hot Rolled at 400 °C HBSS −1.544 9.58 0.22 – – [ 88 ] Mg-1Zn Hot Rolled at 400 °C SBF −1.805 40.78 0.92 – – [ 88 ] Mg-1Zn Hot Rolled at 400 °C HBSS −1.549 7.55 0.17 – – [ 88 ] Mg-4Zn MDF PBS −1.65 2.6 0.06 – 1.02 (after 11 days) [ 43 ] Mg-4Zn Extruded PBS −1.53 1.2 0.03 – 0.72 (after 11 days) [ 43 ] Mg-4Zn Extruded PBS −1.65 ±0.03 60.3 ± 3.3 1.37 ± 0.07 – – [ 76 ] Mg-6Zn Extruded SBF −1.62 45.0 – ∼0.07 (after 30 days) – [ 121 ] Mg-6Zn Hot extruded at 250 °C SBF −1.56 – 0.16 0.07 ± 0.02 (after 30 days) – [ 57 ] 4  a t  Z  [  a  f  i  l  t  t  c  c  t m  [  a  b  a  [  p  d  a  6  t  t  z  o  a  i  c  t  b  M  i  i  c  t  C  [  h  .5. Effect of surface treatment on the degradation of Mg-Znlloy Several researchers further employed different surfacereatment techniques to control the degradation rate of Mg-n alloy by postponing the start of the degradation process 132 ]. The coated Mg-Zn alloy showed a lower corrosion rates compared to the uncoated one because the coating sur-ace hinders the transport of ions. The corrosion resistancencreases with an increase in the thickness of the coatingayer. In some cases, the decreased corrosion resistance withhicker coatings as compared to the thinner coatings is due tohe presence of defects such as pores and microcracks on theoating surface. The heat treatment performed after coatingonsolidates the coating layer and thereby further increaseshe corrosion resistance. For Mg-Zn alloys for biomedical application, the coatingaterial can be divided into bioinert or bioactive coating 133 ]. The bioactive coating is more preferable as it helpsn implant to mimic the natural properties of organ. Forone tissue application, various form of Ca-P coating suchs brushite (DCPD) [ 134 , 135 ], hydroxyapatite (HA) [ 68 , 134 ] 136 ], and fluoridated hydroxyapatite (FHA) [ 134 ] were ex-lored to reduce the degradation rate in Mg-Zn alloy. In ad-ition to controlling the degradation rate, the Ca-P coatinglso helps in healing of bone tissue. The FHA-coated Mg-Zn alloy offers long-term stability and corrosion resistancehan DCPD and HA coating due to its largest stack density,he lowest solubility products and a closely packed organi-ation [ 134 ]. Furthermore, more biomineralization behaviourf FHA-coated Mg-Zn alloy was observed due to the appear-nce of more and better crystallized Ca-P formation on thenterface of biomaterials [ 137 ] The carbonate apatite (CAp)oated Mg-Zn specimen also showed lower corrosion ratehan HA-coated one, indicating the more corrosion resistanceehaviour of CAp-coating [ 138 ]. Researchers also performedAO coatings [ 139 ] and MgCO3 ·3H2 O (nesquehonite) coat-ng [ 128 ] on Mg-Zn alloy, improving corrosion resistancen coated ones. Recently, researchers employed double-layeroating and composite coatings such as nicotinic acid pre-reatment (NA)/calcium phosphate composite coating [ 124 ],a-P/chitosan coating [ 140 ], and HA coating on MAO-treated 141 ] Mg-Zn alloys to improve the surface properties and en-anced the corrosion resistance of the alloys. Furthermore, aM.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 501 c  a  s  a55 s  d  r  l  f  a  i  s  t  e5 o  b  [  i  a  i  1  o  a  B  c  m  o c  t  s  m  u  p  n  r  e  o  o  m  r5  h p  Z  i  s  c  n  t  s  i  e  Z  l  o  a  3  [  l  t  m  a  o  t  o  e  o5 d  r  s  o  c  a  a  a  [  s  t  b  Z  t  [  Z  c5 c  w  c  a  a  v  e  f  v  3  ombination of both deformation and coating techniques suchs extrusion + hydrothermal coating showed a lower corro-ion rate for the Mg-4Zn alloy than the extruded Mg-4Znlloy [ 76 ]. . Evaluation of biological safety and functionality .1. Basic principle of biological safety evaluation The biocompatibility of metallic materials depends on itstability in the biological environment. It can be evaluated atifferent levels including cellular, tissue, and human/clinical-elated biocompatibility [ 1 ]. Metal ions and wear debris re-eased by implant degradation may influence the biologicalunctions of cells and tissue adjacent to implanted materi-ls. In the case of a biodegradable metal, its corrosion ratenside the human body will strongly influence its biologicalafety. Therefore, recreation of the corrosion environment inhe implanted site is the key issue for its biological safetyvaluation. .2. In vitro cell culture studies of Mg-Zn alloy Cytotoxicity test is incorporated as the basic evaluationf the biological safety of chemicals since the correlationetween cytotoxicity and human acute toxicity is confirmed 142 ]. In the case of implant materials’ evaluation, cytotox-city correlates to the thickness of the fibrous tissue formedround the implant materials, which is the parameter for thenflammation level induced by implantation [ 142 , 143 ]. ISO0993–5 offers a basic procedure for cytotoxicity evaluationf biomaterials, indicating to use of established cell lines suchs L929 (NCTC clone 929), Balb/3T3, MRC-5, WI-38, Vero,HK-21 and V-79. It also indicates the two methods forytotoxicity testing; a direct contact method and an extractethod. In the extract method, extract condition is crucial tobtain reasonable results for a biodegradable metal. Like ISO 10993–5, the most of cytotoxicity tests employellular proliferation or viability as their endpoints. Beyondhese kinds of tests, materials’ effects on cellular functionuch as specific biomolecule synthesis, mineral nodule for-ation, and messenger RNA expression are also investigatedsing in vitro cell culture technique. These kinds of cytocom-atibility tests become popular in recent years but it should beoticed that their relevance to in vivo or clinical implantationesults is unproved. Similarly, primary cells are frequentlymployed to cytocompatibility tests since establish cell linesften reported their loss in some cellular function which isbserved by in vivo experiments. However, availability of pri-ary cells is limited, resulting in uncertain reproducibility andeliability of their results. This is the reason that ISO 10993– recommends to use the well-established cell lines for theirigh reproducibility and reliability on their results. Tables 8 indicates the results of cytotoxicity and cytocom-atibility evaluation of Mg-Zn alloys. The results for Mg-n specimens with surface treatment are also summarizedn Table 9 . Regarding the uncoated Mg-Zn alloys, the re-ults of the extract method gave no significant cytotoxicityomparing to those without the extract solution either as-castor deformed samples. However, by the direct method, ex-ruded Mg-6Zn alloy specimen suppresses the adhesion andpreading of murine calvarial osteoblastic cells MC3T3-E1n comparison to PLLA [ 121 ]. Therefore, surface coating isffective to improve cell viability on direct contact to Mg-n alloy surface [ 136 , 138 , 139 , 144 ]. HA-coated Mg-3Zn al-oy surface gave 41% higher cell viability of MG-63 (humansteosarcoma-derived cells) than the uncoated one at days 1nd 3 [ 136 ]. However, the cell viability decreases after day than day 1, indicating the influence of substrate corrosion 136 ]. In addition to cytotoxicity of FHA-coated Mg-6Zn al-oy by direct contact, Li et al. confirmed the higher osteoblas-ic differentiation and activity on the coated surface using hu-an bone marrow stromal cells (hBMSCs) [ 144 ]. This resultgrees with that Mg2 + stimulates the osteogenesis [ 145 ]. Thesteogenic activity of the Mg-Zn alloy can be partly ascribedo the Zn2 + , which can stimulate bone formation and increasesteogenic function in osteoblasts by stimulating cell prolif-ration, alkaline phosphatase activity, collagen synthesis, andsteoblast marker gene expression [ 146 ]. .3. Antibacterial effect of Mg-Zn alloy The implants possessing antibacterial behaviour are moreesirable for better clinical performance as they reduce theisk of implant-associated infection. ISO 22196 offers the ba-ic protocols for the evaluation of the antibacterial activityf plastics and non-porous surfaces. The contact condition isrucial for evaluating the antibacterial activity of biodegrad-ble metals [ 150 ]. The antibacterial activity of the Mg-Znlloy was possibly attributed to the synergistic effects of highlkalinity and released Zn2 + due to its degradation process 147 ]. The high alkalinity (pH ≥ 9) decreases the bacterialurface hydrophobicity [ 151 ] and prevented biofilm forma-ion on Mg surface [ 152 ]. Bacterial adhesion and growth cane inhibited by reactive oxygen species (ROS) generated bynO [ 153 ], and Zn2 + are able to inhibit transmembrane pro-on translocation, glycolysis, and acid tolerance in bacteria 154 ]. Therefore, ZnO on the surface of the Mg-Zn alloy andn2 + released by the degradation of the Mg-Zn alloy mayontribute to enhancing its antibacterial properties. .4. Blood compatibility of Mg-Zn alloy The blood compatibility plays a determined role in thelinical applications as the implant inevitably come in contactith blood during implantation. When the implant comes intoontact with blood, plasma protein adsorption occurs within very short time, and the adsorbed plasma protein plays major role in the blood reactions (including platelet acti-ation, hemolytic reaction, and coagulation factor activation,tc.) [ 155 , 156 ]. ISO 10993–4:2002 offers the basic protocolsor the evaluation of the blood compatibility for medical de-ices. Extruded Mg-6Zn alloy showed a hemolysis rate of.4%, which is less than 5%, indicating the non-destructive502 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 Table 8 The results in cytocompatibility evaluation of Mg-Zn alloys. Materials Processing Methods a Cells Inoculated number Incubation period End point Control Results Ref Mg-1Zn As-cast Extract L929 5 × 103 in 100 μL 2, 4, and 7 days Viability Pure Mg No cytotoxicity with enhanced cell viability than pure Mg. [ 88 ] NIH3T3 MC3T3-E1 ECV304 VSMC Mg-5.6Zn As-cast Extract Rat BMSC 3 × 103 per 96 well 1, 3, and 5 days Viability Ti (i) Similar cell viability to control. (ii) More mineralized nodules and higher ALP activity for Mg-Zn extract than control. (iii) Increased expression of differentiation markers for Mg-Zn extract at day 7 but no difference at day 14. [ 147 ] 1 × 105 per 6 well 2 and 3 weeks Calcification 4, 7, and 10 days ALP 7 and 14 days RT-PCR Mg-6Zn Extrusion Direct MC3T3-E1 1 × 104 per 24 well (0.5 ∼1 mL) 2 h Adhesion (morphology) PLLA Mg-6Zn is able to support earlier adhesion of cells, but they do not spread as sufficiently as those on PLLA. [ 121 ] Mg-5.6Zn Extrusion Direct MC3T3-E1 1 × 105 /cm2 (in 24 well) 1, 2, 4h Attachment PLLA (i) More cells attached to Mg-Zn alloy than PLLA. (ii) Greater area of mineralized nodules formed on Mg-Zn alloy than PLLA. (iii) collagen (COL) 1 α1 and osteocalcin (OC) mRNA were at a higher level for Mg-Zn group than PLLA at day 3 and 12, respectively. [ 148 ] 1 × 105 /well 21 days Mineralization Every 3 days RT-PCR Mg-6Zn Extrusion Extract L929 2.5 × 103 in 100 μL 2, 4, and 7 days Relative growth rate (RGR) No sample No significant difference in RGR between the extracts and negative control. [ 57 ] Mg-6Zn Extrusion + aging at 150 °C for 24h Extract L929 not mentioned 1, 4, and 7 days Relative growth rate (RGR) No sample No significant difference in RGR between the extracts and negative control. [ 61 ] a The method how the cells contact to the testing sample; the extract method employs extract solution added into culture medium, and in the direct method, cells are directly inoculated onto the material surface and cultured on it.VSMC, vascular smooth muscle cells. BMSC: bone mesenchymal stem cell. M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509503 Table 9 The list of cytocompatibility evaluation of surface-coated Mg-Zn alloys. Materials Surface modification Methods a Cells Inoculated number Incubation period End point Control Results Ref Mg-1Zn Mg-5Zn Mg-7Zn CAp-coated HA-coated Direct MC3T3-E1 2 × 104 /cm2 (in 2 mL) 3 days Viability Uncoated sample CAp and HA coatings enhanced the cell viability on the Mg-Zn alloys, but the Cap coatings improved it more effectively than the HA coating. [ 138 ] Mg-3Zn HA-coated Direct (after pre-incubation in SBF or D-MEM) MG-63 5 × 104 / 50 μL 4 h 1, and 3 days Adhesion, viability and morphology Uncoated alloy and TCPS (i) Coated sample had better initial cell adhesion even than TCPS. (ii) The coated ones had 41% higher cell viability than the uncoated ones at days 1 and 3, but the cell viability at day 3 decreased than day 1 and TCPS. [ 136 ] Mg-6Zn FHA-coated Direct hBMSC 1 × 104 /disc 24h Morphology Uncoated Mg-6Zn (i) Cells on both uncoated and coated samples spread well. (ii) Higher proliferation observed on coated ones at days 2 and 3. (iii) Higher ALP levels appeared on coated ones at days 14 and 21. (iv) The up-regulation was suggested on coated ones after 21 days of culture. [ 144 ] 5 × 103 /disc 1–4 days Viability 5 × 103 /disc 7, 14, 21 days ALP 2 × 104 /disc 7, 14, 21 days RT-PCR Mg-6Zn FHA-coated Extract hBMSC 5 × 103 per 96 well 2, 4, and 7 days Viability and morphology No sample (i) No significant difference in cell viability at days 2 and 4, but the higher viability on the extracts than the negative control at day 7. (ii) No morphological change was observed in 100% extract. [ 149 ] Mg-6Zn MAO-coated Direct Rabbit BMSC 5 × 104 /cm2 6, 12, and 24 h Adhesion Uncoated sample (i) Coated samples showed improved the cell viability, cell adhesion and spreading. (ii) ALP activity is enhanced with increase in extract concentration of coated samples. (iii) Col-I expression increased on both uncoated and coated samples. [ 139 ] Extract 1 × 103 /100 μL 1, 3, and 5 days Viability 1 × 104 /mL 14 days ALP Col-I a In the extract method, extract solution of the testing materials was added into culture medium. In the direct method, cells are inoculated onto the testing material.FHA, Fluoridated Hydroxyapatite; HA, Hydroxyapatite; CAp, carbonate apatite; MAO, Microarc Oxidation; TCPS, tissue culture polystyrene; ALP, alkaline phosphatase activity assay; RT-PCR, Reverse transcription-polymerase chain reaction assay; hBMSC, human bone marrow-derived stem cell. 504 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 Fig. 9. The corrosion rate of various Mg-Zn alloys after the electrochemical test and immersion test in SBF. E and I indicate electrochemical and immersion tests, respectively. The data sources are following references: E-as-cast from [ 41 , 42 , 44 , 53 , 85 , 88 , 123 , 128 ], E-deformed from [ 44 , 55 , 57 , 88 , 129 ], E-heat-treated from [ 53 ], I-as-cast from [ 41 , 42 , 44 , 53 , 55 , 123 , 130 ], I-deformed from [ 44 , 55 , 131 ], and I-heat-treated from [ 53 ]. e  h  d  c  t  e  o  t  w  F  t  a  f  l  a  c  c  c  f5 t  d  t  Z  b  S r  d  e  [  t  e  i  e  h  b w  s  i  b  i  a  t  a  s  b  w  h t  d  p  b  p  t  f  T  m  p  ffect on red blood cells (RBCs) [ 121 ]. However, pure Mgas a high hemolysis rate of 59.24% because of its higheregradation rate [ 157 ]. The larger increase in local pH ac-ompanying to higher degradation rate of pure Mg promoteshe binding between hemoglobin and cellular membrane andnhances the uptake of Ca2 + by RBCs, resulting in the rupturef RBCs leading to hemolysis. Zn2 + are reported to reducehe fragility of RBCs and to maintain the enzyme activities,hich is contributing to reduction in hemolysis [ 158 , 159 ].urthermore, the less corrosion of Mg-Zn alloys and the an-icoagulant properties of Zn2 + may contribute to less plateletdhesion and activation, that means better anticoagulant per-ormance [ 160 ]. Therefore, the alloying of Mg with Zn mayead to better blood compatibility. The surface treatment suchs MAO coating on Mg-6Zn alloy further improved the bloodompatibility by reducing the hemolysis rate from 4 (withoutoating) to 1.9% [ 139 ], suggesting the further reduction oforrosion reaction and resulting local pH shift is preferableor the devices contacting to blood for long period of time. .5. In vivo implantation studies of Mg-Zn alloy In vivo implantation is a widely used method to estimatehe biological performance and safety of a biomedical deviceue to the assumption of nearly similar physiological condi-ions in the testing animals and humans. Regarding the Mg-n alloys, only one alloy, Mg-5.6Zn was extensively studiedy the group of Shanghai Jiaotong University and Shanghaiixth People’s Hospital. Implantation of a Mg-5.6Zn rod into the distal femur mar-ow cavity of the New Zealand rabbit resulted in the materialegradation of 87% after 14 weeks with no measurable influ-nce in serum magnesium, or on liver and kidney functions 57 , 161 ]. The serum uric acid and creatine kinase increasedransiently after the surgery but returned to preoperative lev-ls 1 weeks after the surgery [ 161 ]. There were no changesn the histology of the heart, liver, kidney, or spleen postop-ratively [ 161 ]. These observations suggest that Mg-Zn alloyas good biocompatibility with no damages on vital organsy femur implantation. The degradation behaviour of the implanted Mg-5.6Zn rodas confirmed radiographically; the macroscopic degradationtarts in the first 3 weeks after post-implantation with the ev-dence of fuzzy implant edges accompanied by subcutaneousubbles ( Fig. 10 a), and gradual progress in degradation wasndicated by too blurry radiograph of the implant ( Fig. 10 b)t 12 weeks after post-implantation [ 57 , 161 ] After 3 weeks,he tissues around the gas bubble are composed of two layers,n inner compact one and an outer loose connective tissue (ashown in Fig. 11 ) [ 57 ]. No adverse effects due to these gasubbles were observed and they disappeared after 6 weeksithout any intervention [ 57 , 161 ], indicating the diffusion ofydrogen gas into the surrounding tissues. Observation of the cross-sections of the implanted rods af-er 12 weeks of post-implantation also confirmed the markedegradation of Mg-5.6Zn rod with the irregular shapes accom-anied by the layer of degradation products and newly formedone [ 148 ]. Neither inflammation nor fibrous membrane wasresented at the interface between the Mg-5.6Zn rods andhe bone tissue, whereas continuous fibrous membranes wereormed for PLLA rods [ 148 ]. Furthermore, in comparison toi Kirschner wires, Mg-5.6Zn rods had more new bone for-ation and enhanced bone-to-implant contact around the im-lants after 8 weeks of post-implantation into the distal femurM.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 505 Fig. 10. Radiographs of hot extruded Mg-5.6Zn rod at 3 weeks (a) and 12 weeks (b) post-implantation. Reprinted from Acta Biomaterialia, vol 6, Issue 2, Shaoxiang Zhang, Xiaonong Zhang, Changli Zhao, Jianan Li, Yang Song, Chaoying Xie, Hairong Tao, Yan Zhang, Yaohua He, Yao Jiang, Yujun Bian, Research on an Mg–Zn alloy as a degradable biomaterial, page no. 626–640, Copyright 2010, with permission from Elsevier [ 57 ]. Fig. 11. Hematoxylin and eosin (HE) stained tissues around the gas bubble at 3 weeks post-implantation. Reprinted from Acta Biomaterialia, vol 6, Issue 2, Shaoxiang Zhang, Xiaonong Zhang, Changli Zhao, Jianan Li, Yang Song, Chaoying Xie, Hairong Tao, Yan Zhang, Yaohua He, Yao Jiang, Yujun Bian, Research on an Mg–Zn alloy as a degradable biomaterial, page no. 626–640, Copyright 2010, with permission from Elsevier [ 57 ]. o  i  i  p  a m  t  1  a  t  p f  M  a  F  p  c  a  i  i  o  t  t  t  o  e  t6 h  l  I  t  t  i  g  m  t  t  w  b  p  t  fi  o  t d  c  i  d  s  i  i  f the rat [ 147 ]. Mg-5.6Zn pin had better healing with reducednflammation than Ti pin after 4 weeks of post-implantationnto the cecum of the Sprague-Dawley rats [ 162 ]. At this im-lantation site, no emphysema was formed due to the rapiddsorption of hydrogen gas by the intestine [ 147 ]. Implantation of an extruded Mg-5.6Zn tube into the com-on bile duct (CBD) of adult New Zealand rabbits confirmedhe majority of the Mg-5.6Zn tube remained in the CBD after-week post-implantation, whereas 91% of the tube degradedfter 3 weeks of post-implantation [ 163 ]. These data indicatehat the degradation rate of the implant depends on the im-lantation site. To suppress the quick degradation and improve the inter-ace bioactivity of the implant, FHA coating was applied tog-5.6Zn alloy [ 149 ]. Implantation into the femoral shaft ofdult New Zealand rabbits confirmed more direct contact onHA-coated alloy than uncoated one after 1 month post im-lantation [ 149 ]. From the above literature, it can be con-luded that Mg-5.6Zn alloy degraded in a controlled mannerccompanied by new bone formation with enhanced bone-to-mplant contact ( Table 10 ). The released Mg2 + and Zn2 + dur-ng the degradation of Mg-5.6Zn alloy may help the formationf new bone near the implant site. The degradation rate ofhe implant also depends on the implantation site. However,he implantation condition into the animal body differs fromhat into human body in terms of size, body weight, amountf body fluid, etc. which sometimes leads to undesired resultsven though they had good compatibility by in vivo implan-ation tests. . Future areas and challenges This section addresses the difficulties that the researchersave encountered in the studies of biodegradable Mg-Zn al-oys and the topics that require additional focus in this field.n the future, the combination of thermo-mechanical deforma-ion and surface modification techniques could attract atten-ion. As described Section 3 , the highest UTS of Mg-Zn alloys 301 MPa with 4 wt.% Zn, which is lower than that of Tirade 2, which is employed for conventional nondegradableicroplates and screws as temporary implants. The applica-ion of the material with lower mechanical properties bringshe increase in the device thickness, which is generally un-elcome due to the spatial limitation in the tissue for theone fixture implant. Therefore, deformation techniques arerimarily utilized to increase mechanical properties, which ishe basic requirement for the application to temporary bonexture devices. The effect of different processing techniquesn the degradation behaviour and following antibacterial ac-ivity of Mg-Zn alloy also needs attention. The surface modification is mandatory to avoid the earlyegradation of the implant, which causes the formation of gasavity in surrounding tissue, resulting in retardation of heal-ng process. In one of the in vivo implantation studies intro-uced in Section 5.4 , the tunnel of the rabbit femur made byham operation in control group healed within 3 weeks, butn the Mg-5.6Zn implant group, the large gaps between themplant and bone tissue were observed even after 14 weeks506 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 Table 10 The list of results from the in vivo test of Mg-Zn alloys. Materials Processing Animal Implantation site Implantation period Control material Results Ref Mg-5.6Zn Extruded Rabbit Femoral marrow cavity 18 weeks Without implant (i)New bone formation with 87% of implant material degradation (ii)The gas bubbles were disappeared after 6 weeks of post implantation indicating the diffusion of hydrogen gas into the surrounding tissues [ 57 ] Mg-5.6Zn (Extruded) Rabbit Femoral marrow cavity 14 weeks Without implant Mg-Zn alloy is resorbable (degraded 87%) when implanted in bone. It has no effect on the chromaticness, structure or function of heart, liver, kidney, or spleen. [ 161 ] Mg-5.6Zn (Extruded) Rabbit Femoral marrow cavity 12 weeks Poly-L-lactic acid (PLLA) (i)The Mg-5.6Zn alloy not only degraded faster but was accompanied by more new bone formation than PLLA (ii) The fibrous membrane was observed only at interface between bone tissue and PLLA implant [ 148 ] Mg-5.6Zn (Extruded) Rat Femoral marrow cavity 8 weeks Ti Kirschner wire More new bone formation and enhanced bone-to-implant contact were observed around the Mg-Zn alloy implants than around the Ti implants [ 147 ] Mg-5.6Zn Extruded Rabbit Common bile duct (CBD) 3 weeks Stainless steel After 1 week most of the stent structure retained whereas after 3 weeks 91% of Mg-5.6Zn stent degraded [ 163 ] Mg-5.6Zn Extruded Rat Cecum 4 weeks Ti (i)Better healing and reduced inflammation for Mg-6Zn than Ti (ii)H2 gas was rapidly absorbed by the intestine and not formed emphysema in the implantation site [ 162 ] Mg-5.6Zn FHA coated Rabbit Femoral condyle 4 weeks – FHA coating enhance the interface bioactivity mainly by inducing quicker differentiation and decreasing the degradation rate to make a better contacted interface [ 149 ] o  p  p  t  t  c  t  o  t  o a  d  p  p  e  o  g  w  s  t  e b  e  h  o  a  s  t  c  v  b  g  t  i7 m  t  c  a  o  H  r  s  s  p  h  s  c  o  M h  t  f implantation [ 161 ]. To improve this, FHA-coating was em-loyed and succeeded to have direct contact between the im-lant and bone tissue at 1-month post-operation [ 149 ]. Moreechniques will be studied for the control of implant degrada-ion rate and interfacial activity. Research on the compositeoating layers is one of them, where the inner layer reduceshe rate of substrate corrosion and the outer layer delays thenset of the corrosion process. A thorough investigation intohe impact of different ceramic-polymer coating combinationsn the corrosion process is also necessary. The human interstitial fluid contains inorganic ions as wells organic components such as amino acids, lipids, carbohy-rates, and proteins. Its pH is influenced by the gas com-osition inside the body; 5% CO2 , 5% O2 , and 90% N2 ineripheral tissue. Blood flow supplies nutrition and oxygen tovery tissue via capillary network, which also helps diffusionf metabolites and waste molecules. The fluid composition,as partial pressure, and diffusion rate by circulation differith implanting tissue and organ. Therefore, more research istill needed to analyse the mechanical and corrosion proper-ies of the Mg-Zn alloy in the intricate human physiologicalnvironment. Even though a few in vivo studies on Mg-Zn alloys haveeen conducted, more thorough in vivo investigations are nec-ssary prior to conducting clinical trials. The mechanism ofydrogen absorption and whether the hydrogen can be metab-lized or will accumulate in certain organs are still unknownnd need further detailed investigation. Furthermore, the re-ults of animal studies not always give the same results tohose in human clinical cases. The recreation of pathologicalondition is a critical issue to lead the success of device de-elopment in an efficient manner. Not only an in vivo modelut also in vitro experimental techniques including tissue en-ineering should be investigated for better understanding ofhe requirements for the implant device treating a specificnjury or diseases. . Conclusions An ideal biodegradable material should possess sufficientechanical properties, a degradation rate aligning with theissue healing process, and excellent biocompatibility. In thease of Mg-Zn alloy, the degradation rate decreases with theddition of up to 5 wt.% Zn, achieved through the formationf ZnO, which is more stable than MgO in aquas condition.owever, further additions lead to an increased degradationate due to the formation of a more continuous network-liketructure of the secondary phase, leading to localized corro-ion. Similarly, mechanical properties exhibit significant im-rovement up to 4 wt.% Zn, beyond which no notable en-ancement is observed due to the formation of a networktructure with dendritic segregation along grain boundariesaused by a higher number of second phases. Therefore, theptimal composition for implant applications is identified asg-Zn alloy with 4.0-5.0 wt.% Zn. Various thermo-mechanical processing techniques can en-ance mechanical properties and corrosion resistance, but op-imization of processing parameters is crucial. DecreasingM.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 507 g  a  a  e  o  t  p  t i  s  t  i  a  b  a  r a  W  r  gD fi  aC –  o  W  s  tDR                                        rain size and reducing the secondary phase through suit-ble deformation processes improve mechanical propertiesnd increase corrosion resistance. The decreased grain sizenhanced surface homogeneity, while a lower amount of sec-ndary phase results in a discontinuous distribution, reducinghe degradation rate. However, in some cases, lower grain sizeost-deformation may not decrease the degradation rate dueo the presence of high dislocation densities. Coating Mg-Zn alloy reduces corrosion rate by hinderingon transport. Mg-Zn alloys processed by suitable techniquesuch as deformation and proper surface treatment showed bet-er cell adhesion and proliferation, and excellent biocompat-bility compared to the as-cast alloy. The Mg2 + produced as result of in vivo degradation help in the formation of newone near the implant site. The presence of Zn in the Mg-Znlloy may contribute to strengthen the bone-implant contactegion. In conclusion, biodegradable Mg-Zn alloys show promises potential candidates for temporary implant applications.ith further development, they are anticipated to partiallyeplace traditional metallic implants and serve as next-eneration implants. eclaration of competing Interest The authors declare that they have no known competingnancial interests or personal relationships that could haveppeared to influence the work reported in this paper. RediT authorship contribution statement Manas Ranjan Sahu: Writing – review & editing, Writingoriginal draft, Visualization, Validation, Software, Method-logy, Investigation, Conceptualization. Akiko Yamamoto:riting – review & editing, Writing – original draft, Supervi-ion, Resources, Methodology, Funding acquisition, Concep-ualization. ata availability Data will be available on request. eferences [1] Y. Liu, et al., Adv Funct Mater 29 (18) (2019), doi: 10.1002/adfm.201805402. [2] J. Venezuela, M.S. Dargusch, Curr Opin Solid State Mater Sci 24 (3)(2020) 100822, doi: 10.1016/j.cossms.2020.100822. [3] C. Shuai, S. Li, S. Peng, P. Feng, Y. Lai, C. Gao, Mater Chem Front3 (4) (2019) 544–562, doi: 10.1039/C8QM00507A. [4] A. Francis, Y. Yang, S. Virtanen, A.R. Boccaccini, J Mater Sci MaterMed 26 (3) (2015) 138, doi: 10.1007/s10856- 015- 5473- 8 . [5] J. Venezuela, M.S. Dargusch, Acta Biomater 87 (2019) 1–40, doi: 10.1016/j.actbio.2019.01.035 . [6] A.R. Khan, N.S. Grewal, C. Zhou, K. Yuan, H.-J. Zhang, Z. Jun, ResEngineering 20 (2023) 101526, doi: 10.1016/j.rineng.2023.101526 . [7] F. Gao, et al., Corros Sci 240 (2024) 112439, doi: 10.1016/j.corsci.2024.112439 . [8] M.R. Sahu, T.S.S. Kumar, U. Chakkingal, J Magnes Alloys 10 (8)(2022) 2094–2117, doi: 10.1016/j.jma.2022.08.002. [9] A. Dehghanghadikolaei, H. Ibrahim, A. Amerinatanzi, M. Elahinia,Biodegradable Magnesium Alloys, 2nd ed., Elsevier Ltd., 2019, doi: 10.1016/b978- 0- 08- 102666- 3.00009- 2. [10] C. Pan, et al., J Magnes Alloys 11 (1) (2023) 48–77, doi: 10.1016/j.jma.2022.12.017 . [11] Y. Lu, S. Deshmukh, I. Jones, Y.-L. Chiu, Biomater Transl 2 (3) (2021)214–235, doi: 10.12336/biomatertransl.2021.03.005 . [12] Y. Yamasaki, et al., Biomaterials 24 (27) (2003) 4913–4920, doi: 10.1016/S0142- 9612(03)00414- 9 . [13] V. Tsakiris, C. Tardei, F.M. Clicinschi, J Magnes Alloys 9 (6) (2021)1884–1905, doi: 10.1016/j.jma.2021.06.024. [14] Y.F. Zheng, X.N. Gu, F. Witte, Mater Sci Engineering R: Reports 77(2014) 1–34, doi: 10.1016/j.mser.2014.01.001 . [15] M. He, L. Chen, M. Yin, S. Xu, Z. Liang, Journal of Mater ResTechnol 23 (2023) 4396–4419, doi: 10.1016/j.jmrt.2023.02.037 . [16] D. Liu, D. Yang, X. Li, S. Hu, J Mater Res Technol 8 (1) (2019)1538–1549, doi: 10.1016/j.jmrt.2018.08.003 . [17] M. Hashemi, R. Alizadeh, T.G. Langdon, J Magn Alloys 11 (7) (2023)2260–2284, doi: 10.1016/j.jma.2023.07.009 . [18] M. Sabbaghian, R. Mahmudi, K.S. Shin, J Magnes Alloys 7 (4) (2019)707–716, doi: 10.1016/j.jma.2019.11.001 . [19] Y. Chen, S. Zhang, J. Li, Y. Song, C. Zhao, X. Zhang, Mater Lett 64(18) (2010) 1996–1999, doi: 10.1016/j.matlet.2010.06.011 . [20] Y. Zhang, J. Li, J. Li, J Alloys Compd 728 (2017) 37–46, doi: 10.1016/j.jallcom.2017.08.264. [21] H. Li, Q. Peng, X. Li, K. Li, Z. Han, D. Fang, Mater Des 58 (2014)43–51, doi: 10.1016/j.matdes.2014.01.031 . [22] K. Szentmihályi, P. Vinkler, J. Fodor, J. Balla, B. Lakatos, Orv Hetil150 (15) (2009) 681–687, doi: 10.1556/oh.2009.28591 . [23] A. Bandyopadhyay, I. Mitra, S.B. Goodman, M. Kumar, S. Bose,Prog Mater Sci 133 (2023) 101053, doi: 10.1016/j.pmatsci.2022.101053 . [24] K. Prasad, et al., Materials (Basel) 10 (8) (2017) 884, doi: 10.3390/ ma10080884. [25] H. Breme, J. Helsen, Metals As Biomaterials, 1998, pp. 1–35 . [26] T. Hanawa, in: Metallic Biomaterials Metals For Medical devices, CRCPress, Boca Raton, 2024, pp. 148–179, doi: 10.1201/9781003357827 . [27] R. Osman, M. Swain, Materials (Basel) 8 (3) (2015) 932–958, doi: 10.3390/ma8030932. [28] T. Hanawa, in: Metallic Biomaterials Metals For Medical devices, CRCPress, Boca Raton, 2024, pp. 180–192, doi: 10.1201/9781003357827 . [29] K. Van de Velde, P. Kiekens, Polym Test 21 (4) (2002) 433–442,doi: 10.1016/S0142- 9418(01)00107- 6 . [30] S.C. Cifuentes, et al., Mater Lett 74 (2012) 239–242, doi: 10.1016/j.matlet.2012.01.134. [31] N. Eawwiboonthanakit, M. Jaafar, Z.A.A. Hamid, M. Todo, B. Lila,Adv Mat Res 1024 (2014) 179–183, doi: 10.4028/www.scientific.net/ AMR.1024.179 . [32] G. Gąsior, J. Szczepański, A. Radtke, Materials (Basel) 14 (12) (2021)3381, doi: 10.3390/ma14123381 . [33] C.C. Ng, M.M. Savalani, M.L. Lau, H.C. Man, Appl Surf Sci 257 (17)(2011) 7447–7454, doi: 10.1016/j.apsusc.2011.03.004. [34] Y. Zhou, et al., J Alloys Compd 687 (2016) 109–114, doi: 10.1016/j.jallcom.2016.06.068 . [35] M.P. Staiger, A.M. Pietak, J. Huadmai, G. Dias, Biomaterials 27 (9)(2006) 1728–1734, doi: 10.1016/j.biomaterials.2005.10.003 . [36] P. Wen, et al., Mater Des 155 (2018) 36–45, doi: 10.1016/j.matdes.2018.05.057 . [37] B. Song, S. Dong, S. Deng, H. Liao, C. Coddet, Opt Laser Technol56 (2014) 451–460, doi: 10.1016/j.optlastec.2013.09.017 . [38] B. Song, S. Dong, Q. Liu, H. Liao, C. Coddet, Mater Design (1980-2015) 54 (2014) 727–733, doi: 10.1016/j.matdes.2013.08.085 . [39] P. Jiang, C. Blawert, M.L. Zheludkevich, Corros Mater Degradation 1(1) (2020) 92–158, doi: 10.3390/cmd1010007 . [40] H. Okamoto, J \ Phase Equilibria 15 (1) (1994) 129–130, doi: 10.1007/ BF02667700. [41] S. Cai, T. Lei, N. Li, F. Feng, Mater Sci Engineering: C 32 (8) (2012)2570–2577, doi: 10.1016/j.msec.2012.07.042. https://doi.org/10.1002/adfm.201805402https://doi.org/10.1016/j.cossms.2020.100822https://doi.org/10.1039/C8QM00507Ahttps://doi.org/10.1007/s10856-015-5473-8https://doi.org/10.1016/j.actbio.2019.01.035https://doi.org/10.1016/j.rineng.2023.101526https://doi.org/10.1016/j.corsci.2024.112439https://doi.org/10.1016/j.jma.2022.08.002https://doi.org/10.1016/b978-0-08-102666-3.00009-2https://doi.org/10.1016/j.jma.2022.12.017https://doi.org/10.12336/biomatertransl.2021.03.005https://doi.org/10.1016/S0142-9612(03)00414-9https://doi.org/10.1016/j.jma.2021.06.024https://doi.org/10.1016/j.mser.2014.01.001https://doi.org/10.1016/j.jmrt.2023.02.037https://doi.org/10.1016/j.jmrt.2018.08.003https://doi.org/10.1016/j.jma.2023.07.009https://doi.org/10.1016/j.jma.2019.11.001https://doi.org/10.1016/j.matlet.2010.06.011https://doi.org/10.1016/j.jallcom.2017.08.264https://doi.org/10.1016/j.matdes.2014.01.031https://doi.org/10.1556/oh.2009.28591https://doi.org/10.1016/j.pmatsci.2022.penalty -@M 101053https://doi.org/10.3390/ma10080884http://refhub.elsevier.com/S2213-9567(25)00028-3/sbref0025https://doi.org/10.1201/9781003357827https://doi.org/10.3390/ma8030932https://doi.org/10.1201/9781003357827https://doi.org/10.1016/S0142-9418(01)00107-6https://doi.org/10.1016/j.matlet.2012.01.134https://doi.org/10.4028/www.scientific.net/AMR.1024.179https://doi.org/10.3390/ma14123381https://doi.org/10.1016/j.apsusc.2011.03.004https://doi.org/10.1016/j.jallcom.2016.06.068https://doi.org/10.1016/j.biomaterials.2005.10.003https://doi.org/10.1016/j.matdes.2018.05.057https://doi.org/10.1016/j.optlastec.2013.09.017https://doi.org/10.1016/j.matdes.2013.08.085https://doi.org/10.3390/cmd1010007https://doi.org/10.1007/BF02667700https://doi.org/10.1016/j.msec.2012.07.042508 M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509                                                              [[  [  [  [42] A.F. Lotfabadi, M.H. Idris, A. Ourdjini, M.R. Abdul Kadir, S. Fara-hany, H.R. Bakhsheshi-Rad, Bullet Mater Sci 36 (6) (2013) 1103–1113,doi: 10.1007/s12034- 013- 0566- 9 . [43] M. Asadollahi, E. Gerashi, R. Alizadeh, R. Mahmudi, J Mater ResTechnol 21 (2022) 4473–4489, doi: 10.1016/j.jmrt.2022.11.041 . [44] Q. Peng, X. Li, N. Ma, R. Liu, H. Zhang, J Mech Behav BiomedMater 10 (2012) 128–137, doi: 10.1016/j.jmbbm.2012.02.024. [45] J. Chen, et al., Metals (Basel) 6 (11) (Oct. 2016) 259, doi: 10.3390/ met6110259 . [46] E. Zhang, D. Yin, L. Xu, L. Yang, K. Yang, Mater Sci Engineer: C29 (3) (2009) 987–993, doi: 10.1016/j.msec.2008.08.024. [47] Y.C. Lee, A.K. Dahle, D.H. StJohn, Metallurg Mater Trans A 31 (11)(2000) 2895–2906, doi: 10.1007/BF02830349 . [48] S. Keerti, A. Gokhale, J. Jain, E.-W. Huang, Tribol Lett 65 (4) (2017)140, doi: 10.1007/s11249- 017- 0925- 6 . [49] M. Němec, V. Gärtnerová, M. Klementová, A. Jäger, Mater Charact106 (2015) 428–436, doi: 10.1016/j.matchar.2015.05.038 . [50] Y. Lu, A.R. Bradshaw, Y.L. Chiu, I.P. Jones, J Alloys Compd 614(2014) 345–352, doi: 10.1016/j.jallcom.2014.06.078 . [51] X. Liu, D. SHAN, Y. SONG, E. HAN, Trans Nonferrous MetalsSoc China 20 (7) (2010) 1345–1350, doi: 10.1016/S1003-6326(09)60302-2. [52] Y. Song, E.-H. Han, D. Shan, C.D. Yim, B.S. You, Corros Sci 60(2012) 238–245, doi: 10.1016/j.corsci.2012.03.030. [53] H.R. Bakhsheshi-Rad, et al., Mater Corrosion 65 (10) (2014) 999–1006, doi: 10.1002/maco.201307492. [54] M. Masoumi, M. Pekguleryuz, Mater Sci Engineering: A 529 (2011)207–214, doi: 10.1016/j.msea.2011.09.019 . [55] S. Nayak, B. Bhushan, R. Jayaganthan, P. Gopinath, R.D. Agarwal,D. Lahiri, J Mech Behav Biomed Mater 59 (2016) 57–70, doi: 10.1016/j.jmbbm.2015.12.010. [56] Z. Zou, J. Chen, H. Yan, B. Su, X. Gong, J Mater Eng Perform 25(5) (2016) 1974–1985, doi: 10.1007/s11665- 016- 2041- 4. [57] S. Zhang, et al., Acta Biomater 6 (2) (2010) 626–640, doi: 10.1016/j.actbio.2009.06.028 . [58] C. Zhao, et al., J Mater Sci Technol 35 (1) (2019) 142–150, doi: 10.1016/j.jmst.2018.09.015 . [59] H.-Y. Ha, J.-Y. Kang, J. Yang, C.D. Yim, B.S. You, Corros Sci 75(2013) 426–433, doi: 10.1016/j.corsci.2013.06.027 . [60] Y. Song, E.-H. Han, D. Shan, C.D. Yim, B.S. You, Corros Sci 65(2012) 322–330, doi: 10.1016/j.corsci.2012.08.037 . [61] Y. Yan, et al., J Alloys Compd 693 (2017) 1277–1289, doi: 10.1016/j.jallcom.2016.10.017 . [62] M. Němec, A. Jäger, K. Tesař, V. Gärtnerová, Mater Charact 134(2017) 69–75, doi: 10.1016/j.matchar.2017.10.017 . [63] M. Němec, V. Gärtnerová, A. Jäger, Mater Charact 119 (2016) 129–136, doi: 10.1016/j.matchar.2016.07.016 . [64] M. Němec, V. Gärtnerová, in: European Microscopy Congress2016:Proceedings, Wiley, 2016, pp. 241–242, doi: 10.1002/9783527808465.EMC2016.5004. [65] F. Long, G. Chen, M. Zhou, Q. Shi, Q. Liu, J Magn Alloys 11 (6)(2023) 1931–1943, doi: 10.1016/j.jma.2021.08.029 . [66] E. Koç, M.E. Turan, Mater Res Express 6 (8) (2019) 0865b5, doi: 10.1088/ 2053-1591/ ab1955 . [67] E. Koç, M.B. Kannan, M. Ünal, E. Candan, J Alloys Compd 648(2015) 291–296, doi: 10.1016/j.jallcom.2015.06.227 . [68] D.N. Pham, S. Hiromoto, M. O, E. Kobayashi, Surf Coat Technol 421(2021) 127414, doi: 10.1016/j.surfcoat.2021.127414. [69] A.R. Setiawan, A.A. Putera Satria, A. Basuki, U.A. Wibowo, IOP ConfSer Mater Sci Eng 553 (1) (2019) 012005, doi: 10.1088/1757-899X/ 553/ 1/ 012005 . [70] P. Liu, H. Jiang, Z. Cai, Q. Kang, Y. Zhang, J Magn Alloys 4 (3)(2016) 188–196, doi: 10.1016/j.jma.2016.07.001 . [71] Z. Zareian, M. Emamy, M. Malekan, H. Mirzadeh, W.J. Kim, A. Bah-mani, Mater Sci Engineer: A 774 (2020) 138929, doi: 10.1016/j.msea.2020.138929 . [72] Y. Zheng, Y. Li, J. Chen, Z. Zou, Corros Sci 90 (2015) 445–450,doi: 10.1016/j.corsci.2014.10.043 . [73] M. Cheng, et al., J Alloys Compd 691 (2017) 95–102, doi: 10.1016/j.jallcom.2016.08.164. [74] Z.R. Zeng, M.Z. Bian, S.W. Xu, C.H.J. Davies, N. Birbilis, J.F. Nie,Mater Sci Engineer: A 674 (2016) 459–471, doi: 10.1016/j.msea.2016.07.049 . [75] X. Qian, et al., Mater Des 224 (2022) 111322, doi: 10.1016/j.matdes.2022.111322. [76] M. Zohrevand, M. Mohammadi-Zerankeshi, F. Nobakht-Farin, R. Al-izadeh, R. Mahmudi, J Mater Res Technol 20 (2022) 1204–1215,doi: 10.1016/j.jmrt.2022.07.072. [77] K. Yan, et al., J Magn Alloys 7 (2) (2019) 305–314, doi: 10.1016/j.jma.2019.02.006 . [78] K. Yan, et al., Mater Sci Engineering: A 739 (2019) 513–518, doi: 10.1016/j.msea.2018.09.007 . [79] X. Luo, S. Yang, M. Li, Z. Tang, S. Wang, B. Huang, TransIndian Instit Metals 74 (12) (2021) 3063–3073, doi: 10.1007/ s12666- 021- 02373- 9 . [80] D. YIN, E. ZHANG, S. ZENG, Trans Nonferrous Metals Soc China18 (4) (2008) 763–768, doi: 10.1016/S1003- 6326(08)60131- 4. [81] M. Saini, World J Clin Cases 3 (1) (2015) 52, doi: 10.12998/wjcc.v3.i1.52. [82] B.P. Zhang, Y. Wang, L. Geng, Biomaterials - Physics and Chemistry,InTech, 2011, doi: 10.5772/23929 . [83] C.J. Boehlert, K. Knittel, Mater Sci Engineering: A 417 (1–2) (2006)315–321, doi: 10.1016/j.msea.2005.11.006 . [84] S. Ramesh, G. Anne, H.S. Nayaka, S. Sahu, M.R. Ramesh,J Mater Eng Perform 28 (4) (2019) 2053–2062, doi: 10.1007/ s11665- 019- 04007- 0. [85] M. Kavyani, G.R. Ebrahimi, M. Sanjari, M. Haghshenas, J Magn Al-loys 4 (2) (2016) 89–98, doi: 10.1016/j.jma.2016.05.002. [86] Q. Li, G.J. Huang, X.D. Huang, S.W. Pan, C.L. Tan, Q. Liu, J MagnAlloys 5 (2) (2017) 166–172, doi: 10.1016/j.jma.2017.06.001 . [87] N. Stanford, M.R. Barnett, in: Magnesium Technology 2012, John Wi-ley & Sons, Inc., Hoboken, NJ, USA, 2012, pp. 207–211, doi: 10.1002/ 9781118359228.ch39 . [88] X. Gu, Y. Zheng, Y. Cheng, S. Zhong, T. Xi, Biomaterials 30 (4)(2009) 484–498, doi: 10.1016/j.biomaterials.2008.10.021 . [89] H. Jia, X. Feng, Y. Yang, J Magn Alloys 3 (3) (2015) 247–252, doi: 10.1016/j.jma.2015.08.006 . [90] M. Sabbaghian, R. Mahmudi, K.S. Shin, Mater Sci Engineering: A792 (2020) 139828, doi: 10.1016/j.msea.2020.139828 . [91] G. Anne et al., 2019, pp. 339–343. doi: 10.1007/978- 3- 030- 05789- 3_50. [92] N. El Mahallawy, A. Ahmed Diaa, M. Akdesir, H. Palkowski, Mater-wiss Werksttech 47 (1) (2016) 37–46, doi: 10.1002/mawe.201500468 . [93] D.B. Prabhu, S. Dhamotharan, G. Sathishkumar, P. Gopalakrishnan,K.R. Ravi, Mater Sci Engineering: A 730 (2018) 223–231, doi: 10.1016/j.msea.2018.06.002. [94] H. Somekawa, A. Singh, T. Mukai, Scr Mater 60 (6) (2009) 411–414,doi: 10.1016/j.scriptamat.2008.11.017 . [95] C. Shuai, S. Li, S. Peng, P. Feng, Y. Lai, C. Gao, Mater Chem Front3 (4) (2019) 544–562, doi: 10.1039/c8qm00507a. [96] G. Song, A. Atrens, Adv Eng Mater 5 (12) (2003) 837–858, doi: 10.1002/adem.200310405 . [97] F. Witte, et al., Biomaterials 26 (17) (2005) 3557–3563, doi: 10.1016/ j.biomaterials.2004.09.049 . [98] J.H. Nordlien, S. Ono, N. Masuko, K. Nisancioglu, Corros Sci 39 (8)(1997) 1397–1414, doi: 10.1016/S0010- 938X(97)00037- 1 . [99] R. Udhayan, D.P. Bhatt, J Power Sources 63 (1) (1996) 103–107,doi: 10.1016/S0378- 7753(96)02456- 1 . 100] K. Aramaki, Corros Sci 43 (8) (2001) 1573–1588, doi: 10.1016/ S0010- 938X(00)00144- X. 101] D. Persaud-Sharma, A. McGoron, J Biomim Biomater Tissue Eng 12(2012) 25–39, doi: 10.4028/ www.scientific.net/ JBBTE.12.25 . 102] D.R. Lide (Ed.), CRC Handbook of Chemistry and Physics, CRC press,Boca Raton, 2005 . 103] Y. Xin, K. Huo, H. Tao, G. Tang, P.K. Chu, Acta Biomater 4 (6)(2008) 2008–2015, doi: 10.1016/j.actbio.2008.05.014. https://doi.org/10.1007/s12034-013-0566-9https://doi.org/10.1016/j.jmrt.2022.11.041https://doi.org/10.1016/j.jmbbm.2012.02.024https://doi.org/10.3390/met6110259https://doi.org/10.1016/j.msec.2008.08.024https://doi.org/10.1007/BF02830349https://doi.org/10.1007/s11249-017-0925-6https://doi.org/10.1016/j.matchar.2015.05.038https://doi.org/10.1016/j.jallcom.2014.06.078https://doi.org/10.1016/S1003-6326(09)penalty-@M 60302-2https://doi.org/10.1016/j.corsci.2012.03.030https://doi.org/10.1002/maco.201307492https://doi.org/10.1016/j.msea.2011.09.019https://doi.org/10.1016/j.jmbbm.2015.12.010https://doi.org/10.1007/s11665-016-2041-4https://doi.org/10.1016/j.actbio.2009.06.028https://doi.org/10.1016/j.jmst.2018.09.015https://doi.org/10.1016/j.corsci.2013.06.027https://doi.org/10.1016/j.corsci.2012.08.037https://doi.org/10.1016/j.jallcom.2016.10.017https://doi.org/10.1016/j.matchar.2017.10.017https://doi.org/10.1016/j.matchar.2016.07.016https://doi.org/10.1002/9783527808465.EMC2016.5004https://doi.org/10.1016/j.jma.2021.08.029https://doi.org/10.1088/2053-1591/ab1955https://doi.org/10.1016/j.jallcom.2015.06.227https://doi.org/10.1016/j.surfcoat.2021.127414https://doi.org/10.1088/1757-899X/553/1/012005https://doi.org/10.1016/j.jma.2016.07.001https://doi.org/10.1016/j.msea.2020.138929https://doi.org/10.1016/j.corsci.2014.10.043https://doi.org/10.1016/j.jallcom.2016.08.164https://doi.org/10.1016/j.msea.2016.07.049https://doi.org/10.1016/j.matdes.2022.111322https://doi.org/10.1016/j.jmrt.2022.07.072https://doi.org/10.1016/j.jma.2019.02.006https://doi.org/10.1016/j.msea.2018.09.007https://doi.org/10.1007/s12666-021-02373-9https://doi.org/10.1016/S1003-6326(08)60131-4https://doi.org/10.12998/wjcc.v3.i1.52https://doi.org/10.5772/23929https://doi.org/10.1016/j.msea.2005.11.006https://doi.org/10.1007/s11665-019-04007-0https://doi.org/10.1016/j.jma.2016.05.002https://doi.org/10.1016/j.jma.2017.06.001https://doi.org/10.1002/9781118359228.ch39https://doi.org/10.1016/j.biomaterials.2008.10.021https://doi.org/10.1016/j.jma.2015.08.006https://doi.org/10.1016/j.msea.2020.139828http://10.1007/978-3-030-05789-3_50https://doi.org/10.1002/mawe.201500468https://doi.org/10.1016/j.msea.2018.06.002https://doi.org/10.1016/j.scriptamat.2008.11.017https://doi.org/10.1039/c8qm00507ahttps://doi.org/10.1002/adem.200310405https://doi.org/10.1016/j.biomaterials.2004.09.049https://doi.org/10.1016/S0010-938X(97)00037-1https://doi.org/10.1016/S0378-7753(96)02456-1https://doi.org/10.1016/S0010-938X(00)00144-Xhttps://doi.org/10.4028/www.scientific.net/JBBTE.12.25http://refhub.elsevier.com/S2213-9567(25)00028-3/sbref0102https://doi.org/10.1016/j.actbio.2008.05.014M.R. Sahu and A. Yamamoto / Journal of Magnesium and Alloys 13 (2025) 486–509 509 [  [  [  [  [  [   [[    [   [[  [  [  [  [  [  [  [  [   [  [  [  [  [  [  [  [   [[  [  [  [  [[[  [  [  [   [  [  [  [  [  [[[  [  [  [  [  [   [  [  [  [  [  [[[  [  104] A. Yamamoto, S. Hiromoto, Mater Sci Engineering: C 29 (5) (2009)1559–1568, doi: 10.1016/j.msec.2008.12.015 . 105] M.K.I. Kanai, Kanai’s Manual of Clinical Laboratory Medicine, 31sted., Kinbara-shuppan, Tokyo, 1998 . 106] Y. Xin, T. Hu, P.K. Chu, Acta Biomater 7 (4) (2011) 1452–1459,doi: 10.1016/j.actbio.2010.12.004. 107] D. Mei, S.V. Lamaka, X. Lu, M.L. Zheludkevich, Corros Sci 171(2020) 108722, doi: 10.1016/j.corsci.2020.108722. 108] T. Kokubo, H. Takadama, Biomaterials 27 (15) (2006) 2907–2915,doi: 10.1016/j.biomaterials.2006.01.017 . 109] I. Marco, F. Feyerabend, R. Willumeit-Römer, O. Van der Biest,in: TMS2015 Supplemental Proceedings, Wiley, 2015, pp. 497–506,doi: 10.1002/9781119093466.ch63 . 110] L. Wei, Z. Gao, RSC Adv 13 (12) (2023) 8427–8463, doi: 10.1039/ D2RA07829E. 111] D.F. Williams, R.L. Williams, in: B.D. Ratner, A.S. Hoffman,F.J. Schoen, J.E. Lemons (Eds.), Biomaterials Science An Introduc-tion to Materials in Medicine, 2nd ed., Elsevier, 2004, pp. 430–439.ch. 6 . 112] A.C. Hänzi, I. Gerber, M. Schinhammer, J.F. Löffler, P.J. Uggowitzer,Acta Biomater 6 (5) (2010) 1824–1833, doi: 10.1016/j.actbio.2009.10.008 . 113] N.I. Zainal Abidin, et al., Corros Sci 75 (2013) 354–366, doi: 10.1016/ j.corsci.2013.06.019 . 114] J.H. Hanks, R.E. Wallace, Exp Biol Med 71 (2) (1949) 196–200,doi: 10.3181/00379727- 71- 17131 . 115] N.I. Zainal Abidin, A.D. Atrens, D. Martin, A. Atrens, Corros Sci 53(11) (2011) 3542–3556, doi: 10.1016/j.corsci.2011.06.030. 116] N.T. Kirkland, et al., J Mater Sci Mater Med 23 (2) (2012) 283–291,doi: 10.1007/s10856- 011- 4517- y. 117] X.B. Chen, et al., Acta Biomater 10 (3) (2014) 1463–1474, doi: 10.1016/j.actbio.2013.11.016 . 118] N.T. Kirkland, J. Lespagnol, N. Birbilis, M.P. Staiger, Corros Sci 52(2) (2010) 287–291, doi: 10.1016/j.corsci.2009.09.033 . 119] I. Marco, et al., Eur Cell Mater 33 (2017) 90–104, doi: 10.22203/eCM.v033a07 . 120] F. Feyerabend, et al., J Mater Sci Mater Med 23 (1) (2012) 9–24,doi: 10.1007/s10856- 011- 4490- 5 . 121] S. Zhang, et al., Mater Sci Engineering: C 29 (6) (2009) 1907–1912,doi: 10.1016/j.msec.2009.03.001 . 122] J.-W. Chang, P.-H. Fu, X.-W. Guo, L.-M. Peng, W.-J. Ding,Corros Sci 49 (6) (2007) 2612–2627, doi: 10.1016/j.corsci.2006.12.011 . 123] H.R. Bakhsheshi-Rad, et al., Mater Corros 65 (12) (2014) 1178–1187,doi: 10.1002/maco.201307588 . 124] Y. Song, D. Shan, E.-H. Han, Mater Sci Engineering: C 33 (1) (2013)78–84, doi: 10.1016/j.msec.2012.08.008 . 125] S. Zhang, J. Li, Y. Song, C. Zhao, C. Xie, X. Zhang, Adv Eng Mater12 (5) (2010), doi: 10.1002/adem.200980052. 126] F. Cao, K. Deng, K. Nie, J. Kang, H. Niu, J Alloys Compd 770 (2019)1208–1220, doi: 10.1016/j.jallcom.2018.08.191 . 127] R. Udhayan, D.P. Bhatt, J Power Sources 63 (1) (1996) 103–107,doi: 10.1016/S0378- 7753(96)02456- 1 . 128] D.B. Prabhu, P. Gopalakrishnan, K.R. Ravi, J Alloys Compd 812(2020) 152146, doi: 10.1016/j.jallcom.2019.152146 . 129] S. Zhang, et al., Mater Sci Engineering: C 68 (2016) 414–422, doi: 10.1016/j.msec.2016.06.017 . 130] M.E. Moussa, H.I. Mohamed, M.A. Waly, G.S. Al-Ganainy,A.B. Ahmed, M.S. Talaat, J Alloys Compd 792 (2019) 1239–1247,doi: 10.1016/j.jallcom.2019.03.363 . 131] J. Dong, et al., J Magn Alloys 10 (9) (2022) 2491–2509, doi: 10.1016/ j.jma.2021.11.018 . 132] Z.-Z. Yin, et al., J Magn Alloys 8 (1) (2020) 42–65, doi: 10.1016/j.jma.2019.09.008 . 133] M.N. Sarian, et al., Bioact Mater 12 (2022) 42–63, doi: 10.1016/j.bioactmat.2021.10.034. 134] Y. Song, S. Zhang, J. Li, C. Zhao, X. Zhang, Acta Biomater 6 (5)(2010) 1736–1742, doi: 10.1016/j.actbio.2009.12.020. 135] S.-X. Zhang, J.-N. Li, Y. Song, C.-L. Zhao, X.-N. Zhang, Front MaterSci China 4 (2) (2010,) 116–119, doi: 10.1007/s11706- 010- 0026- z. 136] S. Singh, et al., JOM 67 (4) (2015) 702–712, doi: 10.1007/ s11837- 015- 1364- 1 . 137] J. Li, et al., Bioinorg Chem Appl 2011 (2011) 1–7, doi: 10.1155/2011/ 192671 . 138] D.N. Pham, S. Hiromoto, T. Yamazaki, M. O, E. Kobayashi, ACS ApplBio Mater 4 (9) (2021) 6881–6892, doi: 10.1021/acsabm.1c00594. 139] S. Yang, et al., Trans Indian Instit Metals 76 (8) (2023) 2273–2283,doi: 10.1007/s12666- 023- 02928- y. 140] I. Kozina, H. Krawiec, M. Starowicz, M. Kawalec, Int J Mol Sci 22(15) (2021) 8301, doi: 10.3390/ijms22158301 . 141] I.M. Ghayad, M.A. Maamoun, W.A. Metwally, Z.M. El-Baradie,A.N. Abdel-Azim, J Mater Eng Perform 25 (10) (2016) 4171–4180,doi: 10.1007/s11665- 016- 2258- 2. 142] N.S. Goud, in: A Comprehensive Guide to Toxicology in Nonclin-ical Drug Development, Elsevier, 2017, pp. 825–840, doi: 10.1016/ B978- 0- 12- 803620- 4.00031- 1 . 143] S. Capuani, G. Malgir, C.Y.X. Chua, A. Grattoni, Bioeng Transl Med7 (3) (2022), doi: 10.1002/btm2.10300. 144] J. Li, et al., Biomaterials 31 (22) (2010) 5782–5788, doi: 10.1016/j.biomaterials.2010.04.023 . 145] S. Yoshizawa, A. Brown, A. Barchowsky, C. Sfeir, Acta Biomater 10(6) (2014) 2834–2842, doi: 10.1016/j.actbio.2014.02.002. 146] X. Wang, A. Ito, Y. Sogo, X. Li, A. Oyane, Acta Biomater 6 (3) (2010)962–968, doi: 10.1016/j.actbio.2009.08.038 . 147] W. Yu, et al., RSC Adv 6 (51) (2016) 45219–45230, doi: 10.1039/ C6RA03998G. 148] Y. He, Int J Mol Med (2011), doi: 10.3892/ijmm.2011.707 . 149] J. Li, et al., Mater Sci Engineering: B 176 (20) (2011) 1785–1788,doi: 10.1016/j.mseb.2011.05.029 . 150] E. Zhang, X. Zhao, J. Hu, R. Wang, S. Fu, G. Qin, Bioact Mater 6(8) (2021) 2569–2612, doi: 10.1016/j.bioactmat.2021.01.030. 151] A. Nostro, et al., APMIS 120 (9) (2012) 733–742, doi: 10.1111/j.1600-0463.2012.02900.x. 152] H. Qin, et al., Biomaterials 53 (2015) 211–220, doi: 10.1016/j.biomaterials.2015.02.096 . 153] H. Hu, W. Zhang, Y. Qiao, X. Jiang, X. Liu, C. Ding, Acta Biomater8 (2) (2012) 904–915, doi: 10.1016/j.actbio.2011.09.031 . 154] T.-N. Phan, T. Buckner, J. Sheng, J.D. Baldeck, R.E. Marquis, Oral Mi-crobiol Immunol 19 (1) (2004) 31–38, doi: 10.1046/j.0902-0055.2003.00109.x. 155] C. Pan, R. Xu, J. Chen, Q. Zhang, L. Deng, Q. Hong, Int J BiolMacromol 271 (2024) 132487, doi: 10.1016/j.ijbiomac.2024.132487 . 156] M.B. Gorbet, M.V. Sefton, Biomaterials 25 (26) (2004) 5681–5703,doi: 10.1016/j.biomaterials.2004.01.023 . 157] J. Gao, L. Qiao, L. Li, Y. Wang, Trans Nonferrous Metals Soc China16 (3) (2006) 539–544, doi: 10.1016/S1003- 6326(06)60094- 0. 158] J. Chen, et al., Prog Org Coat 186 (2024) 108058, doi: 10.1016/j.porgcoat.2023.108058 . 159] Y. Zhang, et al., Nat Med 22 (10) (2016) 1160–1169, doi: 10.1038/nm.4162. 160] C. Pan, et al., Int J Biol Macromol 279 (2024) 135166, doi: 10.1016/ j.ijbiomac.2024.135166 . 161] Y. He, et al., Sci Bull (Beijing) 54 (3) (2009) 484–491, doi: 10.1007/ s11434- 009- 0080- z. 162] J. Yan, et al., J Mater Sci Mater Med 24 (6) (2013) 1515–1525, doi: 10.1007/s10856- 013- 4906- 5 . 163] Y. Chen, et al., Mater Sci Engineering: C 42 (2014) 116–123, doi: 10.1016/j.msec.2014.05.014. https://doi.org/10.1016/j.msec.2008.12.015http://refhub.elsevier.com/S2213-9567(25)00028-3/sbref0105https://doi.org/10.1016/j.actbio.2010.12.004https://doi.org/10.1016/j.corsci.2020.108722https://doi.org/10.1016/j.biomaterials.2006.01.017https://doi.org/10.1002/9781119093466.ch63https://doi.org/10.1039/D2RA07829Ehttp://refhub.elsevier.com/S2213-9567(25)00028-3/sbref0111https://doi.org/10.1016/j.actbio.2009.10.008https://doi.org/10.1016/j.corsci.2013.06.019https://doi.org/10.3181/00379727-71-17131https://doi.org/10.1016/j.corsci.2011.06.030https://doi.org/10.1007/s10856-011-4517-yhttps://doi.org/10.1016/j.actbio.2013.11.016https://doi.org/10.1016/j.corsci.2009.09.033https://doi.org/10.22203/eCM.v033a07https://doi.org/10.1007/s10856-011-4490-5https://doi.org/10.1016/j.msec.2009.03.001https://doi.org/10.1016/j.corsci.2006.penalty -@M 12.011https://doi.org/10.1002/maco.201307588https://doi.org/10.1016/j.msec.2012.08.008https://doi.org/10.1002/adem.200980052https://doi.org/10.1016/j.jallcom.2018.08.191https://doi.org/10.1016/S0378-7753(96)02456-1https://doi.org/10.1016/j.jallcom.2019.152146https://doi.org/10.1016/j.msec.2016.06.017https://doi.org/10.1016/j.jallcom.2019.03.363https://doi.org/10.1016/j.jma.2021.11.018https://doi.org/10.1016/j.jma.2019.09.008https://doi.org/10.1016/j.bioactmat.2021.10.034https://doi.org/10.1016/j.actbio.2009.12.020https://doi.org/10.1007/s11706-010-0026-zhttps://doi.org/10.1007/s11837-015-1364-1https://doi.org/10.1155/2011/192671https://doi.org/10.1021/acsabm.1c00594https://doi.org/10.1007/s12666-023-02928-yhttps://doi.org/10.3390/ijms22158301https://doi.org/10.1007/s11665-016-2258-2https://doi.org/10.1016/B978-0-12-803620-4.00031-1https://doi.org/10.1002/btm2.10300https://doi.org/10.1016/j.biomaterials.2010.04.023https://doi.org/10.1016/j.actbio.2014.02.002https://doi.org/10.1016/j.actbio.2009.08.038https://doi.org/10.1039/C6RA03998Ghttps://doi.org/10.3892/ijmm.2011.707https://doi.org/10.1016/j.mseb.2011.05.029https://doi.org/10.1016/j.bioactmat.2021.01.030https://doi.org/10.1111/j.1600-0463.2012.02900.xhttps://doi.org/10.1016/j.biomaterials.2015.02.096https://doi.org/10.1016/j.actbio.2011.09.031https://doi.org/10.1046/j.0902-0055.2003.00109.xhttps://doi.org/10.1016/j.ijbiomac.2024.132487https://doi.org/10.1016/j.biomaterials.2004.01.023https://doi.org/10.1016/S1003-6326(06)60094-0https://doi.org/10.1016/j.porgcoat.2023.108058https://doi.org/10.1038/nm.4162https://doi.org/10.1016/j.ijbiomac.2024.135166https://doi.org/10.1007/s11434-009-0080-zhttps://doi.org/10.1007/s10856-013-4906-5https://doi.org/10.1016/j.msec.2014.05.014 An overview of the recent developments in biodegradable Mg-Zn alloy 1 Introduction 2 Processing techniques and microstructure of Mg-Zn alloy 2.1 Casting 2.2 Heat treatment 2.3 Deformation techniques 3 Mechanical properties of Mg-Zn alloy 4 Degradation properties of Mg-Zn alloy 4.1 Degradation mechanism of Mg-Zn alloy 4.2 Effect of amount of Zn on the degradation of Mg-Zn alloy 4.3 Effect of heat treatment on the degradation of Mg-Zn alloy 4.4 Effect of deformation techniques on the degradation of Mg-Zn alloy 4.5 Effect of surface treatment on the degradation of Mg-Zn alloy 5 Evaluation of biological safety and functionality 5.1 Basic principle of biological safety evaluation 5.2 In vitro cell culture studies of Mg-Zn alloy 5.3 Antibacterial effect of Mg-Zn alloy 5.4 Blood compatibility of Mg-Zn alloy 5.5 In vivo implantation studies of Mg-Zn alloy 6 Future areas and challenges 7 Conclusions Declaration of competing Interest CRediT authorship contribution statement Data availability References