METAL OXIDE COATING ON

BIODEGRADABLE MAGNESIUM ALLOYS

Pralhad Pesode

School of Mechanical Engineering, Dr. Vishwanath Karad MIT-World Peace

University, Pune-411038, Maharashtra, India

pralhadapesode@gmail.com - https://orcid.org/0000-0001-5604-5740

Shivprakash Barve

School of Mechanical Engineering, Dr. Vishwanath Karad MIT-World Peace

University, Pune-411038, Maharashtra, India

shivprakash.barve@mitwpu.edu.in - https://orcid.org/0000-0002-5372-6310

Sagar V. Wankhede

School of Mechatronics Engineering, Symbiosis Skills and Professional University

Kiwle, Pune-412101, MS, India

svw8890@gmail.com - https://orcid.org/0000-0002-2341-3110

Amar Chipade

Dr. D. Y. Patil Institute of Technology, Pimpri Pune-411018, Maharashtra, India.

amar.chipade@dypvp.edu.in - https://orcid.org/0000-0002-2200-0752

Reception: 22/11/2022 Acceptance: 19/01/2023 Publication: 18/02/2023

Suggested citation:

P., Pralhad, B., Shivprakash, V. W., Sagar and C., Amar (2023). Metal Oxide

Coating On Biodegradable Magnesium Alloys. 3C Empresa. Investigación y

pensamiento crítico, 12(1), 392-421. https://doi.org/

10.17993/3cemp.2023.120151.392-421

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ABSTRACT

Magnesium is a biodegradable metal that has potential in orthopaedics. It has several

advantages over other metallic materials because it is biocompatible and degradable

now being used for biomedical applications, including elimination of stress shielding

effects, enhancing degradation properties and enhancing biocompatibility concern in

vivo, eliminating the second surgery for implant removal. Bioabsorbable magnesium

(Mg) and related alloys have been limited in their usage because of its lower corrosion

resistance. Surface alteration and functionality, in addition to basic alloying, is an

important technique to deal with Mg and its alloys' reduced corrosion resistance.

Magnesium's rapid depreciation however is a double-edged sword because it's critical

to match bone renewal to material corrosion. As a result, calcium phosphate coatings

have been proposed as a way to slow down corrosion. There are various possible

calcium phosphate phases and their coating methods and can give a few distinct

properties to various applications. Despite magnesium's lower melting point and

greater reactivity, calcium phosphate coatings require precise settings to be effective.

Because of their toxicity, non-biodegradability, and much higher cost, the recently

used inorganic conversion coatings are less appealing and their application is limited.

Conversion coatings are a viable alternative technology that is based on a cost-

effective, environmentally friendly, and biodegradable organic component. Surface

chelating functional groups in these compounds allow them to link with the

magnesium/surface hydroxide layer while also providing anchoring groups for the

polymer topcoat. Nanoreservoirs with multilayer inhibitors for active self-healing

corrosion resistance thrive in this environment. This study examines the organic

conversion coatings for Mg and its alloys in depth.

KEYWORDS

Magnesium, Calcium phosphate coating, Conversion coating, Biodegradable,

Biocompatible.

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PAPER INDEX

ABSTRACT

KEYWORDS

1. INTRODUCTION

2. CONVERSION COATINGS ON MAGNESIUM AND ITS ALLOYS

2.1. FLUORIDE CONVERSION COATING

2.2. IONIC CONVERSION COATING

2.3. BIOMIMETIC COATINGS

2.4. HYDROTHERMAL COATING

2.5. ALKALI-HEAT TREATED CONVERSION COATING

3. MAO COATING

3.1. ANTIBACTERIAL MAO COATINGS ON MG ALLOYS

4. BIOCOMPATIBILITY OF CONVERSION COATINGS

5. CONCLUSION

ACKNOWLEDGEMENTS

REFERENCES

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1. INTRODUCTION

To replace and repair damaged body tissue implants are broadly utilized in

biomedical field [1,4]. Depending on type of material use, implants can be

differentiated into three classes [5,6]: (i) ceramic, (ii) metal, and (iii) polymer. Metallic

implants have excellent mechanical properties and good biocompatibility, due to which

they are broadly utilized in dental, cardiovascular and orthopaedics applications [1].

Conventional metallic implant materials incorporate Ti alloy, Co-Cr alloy and SS [7,

131]. However, there is a stress shielding issue because of the substantial difference

in modulus of elasticity of metallic implants and human bone, as indicated in table 1,

which leads to osteoporosis and bone fusion [1]. Also, metallic implant generates

metallic irons due to corrosion or erosion near to implant if these implants have been

there for longer period of time, this might prompt issues, like inflammation of tissue [8].

At last, additional medical surgery may require to eliminate implants [6]. Availability of

biodegradable materials have prompted new improvements in implant innovation.

Table 1. Cortical bone mechanical properties as well as implant materials [1].

The primary purpose of biodegradable implants is to encourage tissue

development, cure specific injuries, and then vanish in vivo by a breakdown process

with little tissue damage [1]. Among orthopaedic application, bone screw and metallic

bone plates are ordinarily used to fix crack sites before new bone development [5].

During rehabilitation, it has been discovered that the implant's strength gradually

diminishes as the strength of the new bone grows [1,132]. Magnesium-based alloys

are currently getting a lot of consideration and are being explored as more recent

types of biodegradable material [8,9]. This is due to a number of factors. For starters,

magnesium-based alloys have superior mechanical qualities for load-bearing

applications than polymers, including high strength and malleability, in addition, they

are biodegradable in vivo [9, 10]. Second, magnesium-based alloys have modulus of

elasticity that are near to natural bone as compared to ordinary metallic implants,

eliminating the stress shielding effect [11]. Magnesium-based alloys offer excellent in-

vivo degradability; hence, using magnesium alloy to remove a temporary implant can

avoid the need for additional medical operations, so significantly reducing a patient's

suffering [7]. Magnesium-based alloys also exhibit good biocompatibility, osteogenesis

induction, anti-inflammatory characteristics, and other bio functional qualities [12,13].

Materials

Young’s Modulus

(GPa)

UTS (MPa)

Yield strength

(MPa)

% Elongation

Cortical bone 5-23 35-283 -- 1.07-2.10

Stainless steel

193-200

480-620

190

Titanium alloys 100-125 550-985 420-780 12-16

Pure iron

195-230

200

150

Co-Cr alloys 210 450-960 310-440 10.7-18.5

Pure Mg (As-cast)

41-45

90-190

20.9

DL-PLA 70-120 40-200 -- 3.10

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Furthermore, it containing specific elements, for example Cu, Ag, Ga, and Sn may

likewise have good antibacterial characteristics [14,15]. It indicates that magnesium-

based alloys have extraordinary development potential as biodegradable materials.

Researchers and industry interest in magnesium- based alloy as biodegradable

materials emphatically grown in the course of the last many years. Magnesium based

alloys possess very good biocompatibility, excellent corrosion behaviour, and

excellent mechanical properties when proper process parameters and right alloying

elements are utilized. Likely uses of magnesium-based alloys are temporary

cardiovascular devices and structural alloys for orthopaedic applications. Once these

implants have served their limited purpose, the body consumes them, such as

scaffolding, mechanical support and attaching to living tissue. Researchers and

industry are exceptionally talking about the connection between magnesium-based

alloys in-vitro and in-vivo characteristics, that could assist in reducing animal testing

and backing simulations to choose alloys. Mechanical characteristics are typically

portrayed by hardness test and tensile test. Additionally, information on stress

corrosion and fatigue is anticipated to give a complete idea of stability over the course

of degradation. Alloys made with magnesium are prone to pitting corrosion. A

consistent corrosion morphology needs to be given special consideration since

corrosion pits increase the intensity of stress during mechanical strain and contribute

to the premature failure of implants [2, 132]. First and primarily, the mechanical and

physical characteristics of magnesium-alloys do not yet match that of implant

materials, and their rapid disintegration can cause mechanical instabilities before the

bone healing process is completed [9,16]. Second, the rate of breakdown of

magnesium-based alloys is very quick, particularly in a chloride medium like a human

physiological fluid. The degradation product of Mg alloys is probably going to create

some issues, for example, tissue inflammation [17]. Finally, corrosion of magnesium-

based alloys is not uniform, which could lead to premature implant failure [18]. Various

tests have been carried out to date with the goal of breaking through these barriers.

Regardless of this advantageous property, magnesium alloys used in biomedical

applications must be carefully monitored for their ability to corrode when in contact

with ECF containing Cl iron. Due to low electronegativity of magnesium alloy, their

corrosion rate is high in in-vivo. From one viewpoint, this implies that according to a

thermodynamic perspective, a biomedical implant developed from magnesium alloy

probably won't be viewed as a reasonable alternative in the exceptionally corrosive

body fluid.

In general, three methods can be used to protect biomedical magnesium against in

vivo corrosion: (i) adding defensive coatings to isolate body fluid and implant. (ii)

alloying with biocompatible elements to protect surface from corrosion (iii)

microstructural surface alteration. During initial phase of implantation, high corrosion

resistance, and even the establishment of a homogeneous, controlled, and anticipated

corrosion rate, are critical. Surface modifications of magnesium alloy for corrosion

control includes micro arc oxidation (MAO), hydrothermal treatments, anodization,

electrophoretic deposition (EPD), physical vapor deposition (PVD), electrochemical

deposition (ECD), sol-gel deposition, magnetron sputtering, and a few other, less

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known process for example, electrospinning, cold gas dynamic showering (CGDS),

phosphatization, laser cladding [3,19,20]. However, regardless of numerous multi-

pronged methodologies towards improvements of magnesium alloy for biomedical

applications, significantly more exploration is needed to overcome their insufficiencies

in demanding biomedical field. This argument is supported by the astounding amount

of effort, resources, funds, and research capacity put into adapting a chemical like

magnesium, which is fundamentally inappropriate for use in medicinal applications. At

the end, simply by applying protective coatings along with suitable alloying elements,

it became conceivable to use Mg alloy in biomedical applications because of its

promising mechanical and natural properties as a cutting-edge biomaterial. Currently,

much research and development is focused on the development of magnesium alloy

for cutting-edge biomaterials, To address the stress shielding effect, future isoelastic

arthroplastic implants could be employed [3, 21–25] if they have lower young's

modulus and are acceptable with the nearby cortical bone. Paper is focusing on metal

oxide coating on magnesium alloy to improve its performance in biomedical

applications.

2. CONVERSION COATINGS ON MAGNESIUM AND ITS

ALLOYS

Surface coating is formed by electrochemical or chemical treatment in the

conversion coating process. Coating transforms a metal's or alloy's surface layer into

a thin coating of metallic oxide or any other chemically related substances. The

coating creates a corrosion-resistant surface that better adheres to the topcoat. A

successful conversion coating should be (i) inert/insoluble (ii) self-healing (iii) resistant

to mechanical damages, (iv) impermeable to liquids/gas (v) eco-friendly (vi) cost

effective. Chemical conversion coatings are manufactured via non-electrolytic

chemical reactions without the use of electricity, whereas electrochemical conversion

coatings use an electrolyte and electricity [26]. The primary benefits of chemical

conversion coatings are process speed, simplicity, low capital and working expenses,

lower energy utilization, low-temperature process, high efficiency and less treatment

time [26,27]. It was found that corrosion fatigue properties were improved by

conversion coating [28]. A conversion coating, in general, functions by driving

interfacial reactions and resultant coating/precipitation formation. A reduction in pH

and a rise in Mg2+ concentration at the solution/metal contact initiate the interaction.

The chemical has a minor dissolving effect on the hydroxide/oxide layer, which will aid

further solution penetration into the layer. Many variables have impacts on the quality

of conversion coating, for example, phase or compound structure, types of pre-

treatment processes, concentration and composition of bath solution, pH of bath,

temperature of bath, time of immersion, degree of stirring, and condition of post

treatment [29]. By and large, different functional additives are used to accelerate

coating formation and to control bath pH value. By changing experimental condition,

thickness of coating can be varied from several hundred nano meter to a micro meter.

The thickness of the coating might range from a few hundred nanometers to a few

micrometres, depending on the trial circumstances used. It was observed that

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thickness of coating can be varied by addition of nanoparticles to the conversion

solutions.

Figure 1. Potential advantages of organic conversion coatings [26]

Inorganic coatings, in which the chemical reacting with the surface of metal to

generate highly adhesive coatings and corrosion resistance coatings, are among the

various chemical conversion coatings explored and used in practise. Figure 1 depicts

the most widely used inorganic conversion coatings on magnesium alloy. Chrome

manganese, phosphate treatments, chrome-pickle, dichromate and ferric nitride pickle

are most normal industrial conversion coatings. For biomedical applications fluoride

and calcium phosphate containing conversion coatings are significant [27]. Chemical

breakdown of rare earth and phosphate-based conversion coatings occurs at acidic

and alkaline pH levels, respectively [30]. Phosphates in water may also promote

nutritional replenishment. Molybdate conversion coatings decreased oxidising

capacity isn't up to par. Heavy metal contamination is the main source of worry. When

reactive magnesium comes into contact with certain inorganic materials, it can

become anodic. As a result, research into non-toxic and environmentally acceptable

conversion coatings for magnesium and related alloys is crucial. In this case, organic

conversion coatings (OCCs) are crucial. Precipitation and chemical dissolution

interact to form chemical conversion coatings. Chemical conversion coating generally

carried out in bath of phosphate, fluoride, chromate and carbonate [32]. Because of its

low cost and convenience of application, conversion coating can be employed in a

various biomedical field. Regardless of the fact that chromate-based coatings offer

outstanding corrosion resistance to Mg alloys, they have been outlawed because of

concerns about the human health and environment. As a result, biomedical

applications of phosphate conversion coatings [33, 34], fluoride [35], and MAO coating

have gotten a lot of interest. Figure 2 depicts a few phosphate-based conversion

coatings, which are typically Zn, Mg, Mn, Ce, Sr, and Ca phosphates. Ce, Sr, K, F, and

Ca doped phosphates are rare. Regrettably, several conversion coatings have

riverbed-like topologies on surfaces, limiting their robustness. In this case, it's also

necessary to change the coating. Due to high biocompatibility, high temperature

resistance, insoluble in water and chemical stability of calcium phosphate and zinc

phosphate, they have been accounted as a possible option in contrast to chromate

coating for biomedical applications [36]. The morphology and structure of those layers

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exceptionally relies upon the temperature of process, composition of conversion bath

and pH value [37].

Figure 2. Mg alloys with phosphate-based conversion coatings [31]

A lot of strategies have been utilized for improving Mg-based biomedical implants

and devices. Calcium phosphate (Ca-P) coatings are non-toxic and have excellent

osteoconductivity. As a result, numerous researchers are concentrating their efforts on

Ca-P coatings in bone replacement, orthopaedics, and other fields. The elements P

and Ca combine to develop a HA layer and is the primary component of natural bone.

The critical aspect is changing the phase and increasing the concentration of HA in

Ca-P coatings, which are fundamentally non-crystalline and contain relatively lesser

amount of HA. HA can develop across specified range of Ca/P proportions under sub-

alkaline and neutral circumstances, as per calculation of the thermodynamic chemical

reaction [34], HA is exclusively produced in a narrow range of Ca/P proportions in

sub-alkaline and neutral circumstances.

One of the researchers [38] established a Ca-P coating on magnesium consisting

of Mg (H2 PO4)2, Ca3(PO4)2, Ca (H2 PO4)2, and Mg 3 (H2 PO4)2 to reduce

corrosion of AZ31B in SBF. The morphology of the surface, EDS finding, and cross-

sectional area of the Ca-P coating on a surface with a thickness of roughly 20 m

reveal a regular petal-like crystal made up of small long blocks. In comparison to its

substrate, the Ca-P coating has a corrosion current density that is two times lower. As

the immersion period in SBF grows, the hemolytic rate of AZ31B increases without

and with coating at first, then reduces to a safe level. The load bearing capacity of the

Ca-P coating for AZ31B decreases noticeably while submerged in SBF, and after 120

days of submersion, the loading capacity was reduced to 85%. Even yet, it's

impossible to say which elements are significant in the progression of corrosion

resistance.

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To seal the structural design, one of the researchers [39] applied a Ca-P coating to

MAO oxidised Mg by immersing it in a Ca-P bath. The upper layer is made up of

dicalcium phosphate dihydrate (DCPD) and HA, according to the findings. At various

temperatures, flake-like and round-shaped morphologies appeared on the surface of

the substrate, demonstrating that temperature has an impact on coating structure. The

volume of hydrogen produced in MAO and soaked MAO after 300 hours of immersion

in SBF was around 17.75 and 1.11 mL/cm2, respectively. After submerging the coated

specimen in SBF, bone-like apatite formed. The thickness was about 42 m after 4

weeks of soaking in SBF Ca-P coatings. It has been observed that corrosion

performance of coating enhanced. Instead of the Zn-P coating's dry-riverbed-like

structure, one of the researchers [36] coated Mg alloy AZ31 with a Ca-doped zinc

phosphate (Zn-Ca-P) coating that had a flower-like structure. The structure of Zn-Ca-P

coatings were found to be tightly connected to the microstructure, including particle

size and secondary phases, as well as the chemical composition of the substrate [40].

Form, orientation, and size of AlMn particles were discovered to take a significant

effect on Zn-Ca-P coatings characteristics. Corrosion of the -Mg network near an

upright AlMnSi particle almost likely results in an obstructed and deeper hole.

Horizontal particles, on the other hand, may form a flat, open, and shallow trench.

Because of this, small pits rather than deeper holes are where phosphate nuclei can

easily accumulate. The outer layer of AM30 therefore develops a coarse grain coating.

Under magnetic field, one of the researchers [41] developed Mn-P conversion coating

on AZ91D Mg alloy. The findings demonstrated that, during the production of the

phosphate conversion layer, superpositioning magnetic fields might speed up the

synthesis of minute hydrogen gas bubbles as well as their speedy desorption from the

surface. Following that, it was discovered that Mg2+ cations are evenly distributed

throughout the alloy, regardless of its microstructure. Immersion in a solution

treatment bath can provide a smooth and homogeneous phosphate conversion

coating when the magnetic field is delivered perpendicular to the substrate rather than

parallel to the substrate. It's worth mentioning that Ce-P, Zn-Ca-Ce-P, and Mn-P

coatings have yet to be shown biocompatible. The biocompatibility of Sr-p ZnCaP and

CaP coatings, on the other hand, is excellent [33,42].

2.1. FLUORIDE CONVERSION COATING

The use of fluoride conversion coatings as biomaterials appears to be a viable

option [43, 44]. Fluoride conversion coating is created by a chemical reaction with Mg

alloys in a hydrofluoric acid (HF) solution bath. Fluorine in bones can help with

calcium and phosphorus digestion and improve bone strength. The major component

of fluoride conversion coatings, magnesium fluoride (MgF2), is water insoluble and

binds to the Mi-alloys surface satisfactorily. Because of their great biocompatibility,

outstanding corrosion resistance, and cell responsiveness, magnesium alloys with

MgF2 coatings have been explore in the biomedical appplications [45]. It has been

observed that HF concentration has great impact on protective efficiency and

performance of coating. One of the studies [46] suggested that a MgF2 coating on the

LAE442 alloy could decrease the corrosion rate in-vivo. Despite the fact that the

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substrate was slightly protected, the MgF2 layer deteriorated completely after a month

of implantation. Despite this, fluoride levels in the adjacent bone did not rise over the

first six weeks after implantation. In MgF2 coated Mg implants, there was restricted

pitting erosion but no subcutaneous gas pits. Regardless of the fact that fluoride

conversion coating could offer safety during the initial phases of implantation, due to

the relatively thin coating, in terms of prolonged use, it must be very effective. The

detrimental effects of high fluoride on bone are also present at the same time, fluoride

release during Mg implants breakdown and its harmfulness for implanted tissue has

been unclear yet. One of the researchers [47] used a novel method to make fluoride

conversion coatings by immersing an AZ61 sample in Na [BF4] liquid salt for varied

durations of time at 430°C and 450°C, then heating the sample to eliminate any

remaining salts and external layer. The two-layer structure of the 2-µ

m coating made

of a larger interior MgF2 layer and a minor exterior NaMgF3 layer [47]. As the

treatment period increases from 2 h at 450 °C, the corrosion current density, i corr, of

the coatings decreases, implying higher corrosion resistance in SBF. The untreated

substrate had an i corr that was around four times lower than the material that had

been treated for 12 hours at 450°C. Even though a few deformities are seen on the

surface of coating, these imperfections don't arrive at the substrate. The coatings are

a superior alternative for fluoride treatments of magnesium alloys because they are

finer than standard conversion coatings. The biocompatibility of the coating still has to

be established. To increase the corrosion resistance of the conversion coating,

researchers [45] employed a two-step immersion approach to develop a uniform and

thinner MgF 2/polydopamine (PDA) coating on Mg-Zn-Y-Nd alloy. Dopamine and tris-

hydrochloric acid (tris-HCl) bath treatments followed by Mg alloy immersion in HF

bath. The coating is made up of two layers, each of which is roughly 100 nm thick: a

PDA outside layer and a MgF2 inner layer.

2.2. IONIC CONVERSION COATING

Ionic fluids (ILs) are molten organic salts formed entirely of ions at normal

temperature [48]. Because of the combination of a charge delocalized anion and a big

cation, the salts are defined by weak interaction. ILs have been classified as

ecologically friendly or biocompatible compounds in the past. Mg alloys can stay

stable over longer periods time without corroding due to the absence of free H+ or

even other metal cations in ILs, providing ideal environment for active metal film

development control [49]. One of the research groups created an IL film on the basis

of interaction of ILs with highly pure Mg [50], which led to the use of ILs in the creation

of Mg alloy composite films. Various ILs conversion coatings have been proposed in

the past [51-54]. Because of their biocompatibility, phosphonate derivative-containing

coatings have been investigated as a new chemical for Mg alloy corrosion protection.

To decrease deterioration in the human body, it been recommended that the AZ31 Mg

alloy be surface treated in compatible phosphate-based ionic fluids [55]. IL coating's

corrosion resistance and cytotoxicity were investigated, and it was observed that

treatment times have a considerable impact on corrosion resistance. A more

homogenous IL film was generated when the ZE41 Mg alloy was exposed to the IL of

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trihexyl phosphonium diphenyl phosphate at a potential of 200 mV [52]. Because of

the protic ammonium-phosphate and trihexyl tetradecyl phosphonium cation mixed

with organophosphate, it was suggested that ILs, or (CF3SO2)2N- anions, might

reacting with Mg alloy surface to develop an excellent corrosion protection coating

[50,56,57]. The metal phosphates, metal oxides, and carbonaceous compounds in the

IL 300°C conversion coating have a double layer structure having 70-80 nm thickness

and outstanding passivation performance in a 1 wt. percent NaCl. Deep eutectic

solvents (DESs), are another type of ecologically friendly IL [58]. Because they are

non-reactive with water, DESs are easier to synthesise in a pure condition than

conventional ILs. Most of them seem to be biodegradable, and the component's

toxicological qualities are well known [58]. One of the research groups [59] focused on

the creation of DES-based Mg alloy conversion coatings. At 160 degrees Celsius, the

interaction of a ChCl-urea mixture-based DES with the AZ31B Mg alloy was proposed

as a new ionothermal method for generating a corrosion resistant layer. Mg alloy and

DES reacted to produce a MgH2 and MgCO3 conversion coating that replaced a

dangerous process. The conversion coating was also found to be superhydrophobic

after additional reaction with stearic acid. According to electrochemical polarisation

experiments [59], the Mg alloy's corrosion resistance could be increased by the DESs

conversion coating. An electric field was also used by the same research group to

induce the disintegration of DES, which aided the reactivity of the DES/Mg alloy

interface [49]. Surprisingly, the proposed anodic treatment in DESs produced

conversion films on Mg alloy substrates with different nanostructures. It was

discovered that a more corrosion-resistant conversion coating can be created by

increasing anodic current density. On the resulting conversion layer,

superhydrophobic and sliding surfaces can be made to even further improve corrosion

resistance [49]. Excellent conversion coatings with double capacity of super-

hydrophobicity and self-healing may also be produced on the AZ31B Mg alloy in DES-

based media [60]. Using varied external fields at the IL/substrate contact to build more

effective coatings on Mg alloy could be promising.

2.3. BIOMIMETIC COATINGS

Phases of CaP are precipitated out of solution and developed on the necessary

sample in simulated bodily fluids (SBF) under close physiological conditions in

biomimetic techniques for CaP coating deposition [62,63]. CaP coatings made with

this approach can be seen in SEM images of a biomimetically synthesised

hydroxyapatite coating on a Mg-alloys at higher and lower magnification. The

technique offers an affordable option to coat a large number of specimens at once

with a uniform coating and is easy to set up and operate. When compared to CaP

coatings made using non-physiological pH, composition, and temperature, coatings

created in physiological circumstances are expected to produce CaP crystalline

structures that are more similar to bone minerals [64]. According to ongoing research

[65], altering the coating solution's pH and temperature can help some CaP phases

form more quickly than others. Furthermore, it was discovered that the geometry of

the specimen to be coated has an impact on the CaP phases that form on sample

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surface [66]. Although the concentration of the SBF used in the publications is

typically concentrated, it can range from 1 to 10 SBF [63,67-69]. Additionally,

immersion times fluctuate significantly between established protocols, and can be

somewhere around 2 h to 20 days [68,70-73]. While using a biomimetic technique, pH

and temperature appear to have been explored consistently, staying within

physiological range, to be specific at 37 °C and 7.4, respectively. Pre-treatment is

typically necessary in this process to affect the coating properties, the Mg reactivity,

and the coated specimen's surface reactivity. To achieve this, most studies utilise an

alkaline solution, such as NaOH, at various concentrations [74,75], or an acidic

solution, such as HCl [77] or HF [76], at various concentrations. In biomimetic norms,

however, many investigations have focused on temperature pre-medicines [76,78], or

acidic or soluble post-medicines [79]. Covering magnesium and similar alloys using

biomimetic technologies has shown to be a successful solution. Some researchers

have raised concerns about biomimetic coatings on non-biodegradable substrates like

titanium, however they haven't considered the influence of the SBF immersion period

on substrate degradation. Previous research has demonstrated that very lengthy

immersion in SBF can result in thick coating development if phosphate and calcium

ions are abundant [73]. Researchers discovered that a second immersion in a new

SBF solution is effective in these circumstances [73]. Numerous researchers have

found that it is beneficial to modify the biomimetic convention for biodegradable

specimens like Mg and its alloys, by shortening the period spent incubating in SBF,

hence minimising the possibility of coating disintegration [80]. Furthermore, as

indicated by horizontal CaP crystal binding across the coating, biomimetic CaP

coatings on Ti have been labelled as thick, comprehensive, homogenous, and non-

porous [81]. According to one of the studies [82], coating formation took place around

hydrogen bubbles that were developed on magnesium surface while submerge,

resulting in irregular and porous CaP coatings on Mg samples. The non-uniform and

porous CaP coating on Mg alloy, according to several studies, is caused by the

specimen's non-uniform shape [77] and the presence of Mg2+ ions, which prevent

crystal formation [83]. Porosity production on degradable metals like Mg is another

key disadvantage of a biomimetic coating method. Until now, no modification or

variation of this procedure has been successful in achieving strong cohesiveness

between the specimen and the CaP coating. For biomedical applications, such as

orthopaedic applications, it is critical to develop an extremely strong adhesive

covering with a long-life span and the ability to endure surgeries.

Biomimetically developed apatite coating on a pure Mg specimen to improve

corrosion protection were explored by one researcher [73]. In SBF, they used either

single-coated or double-coated substrates, with pure Mg as a control. Other

investigations found that coating Mg alloys AZ91D and AZ31 with this method before

doing in vitro immersion testing, electrochemical [84] and SBF solution [85] provided

similar corrosion resistance. Recent investigations on biomimetic CaP-coated pure

Mg, on the other hand, have shown corrosion protection and enhanced cell adhesion

[75,79]. The biomimetic method has most likely been studied as a viable strategy for

employing CaP on Mg specimens, as previously indicated. Given this, and as reported

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in the literature, further refinement of these methods is required if they are to be used

as a viable CaP coating approach on magnesium and its alloys.

2.4. HYDROTHERMAL COATING

The uncoated Mg alloy's outer layer lacks adsorption sites associated with Ca-P

crystal formation. In the absence of a chelating specialist in hydrothermal coatings

method, high-temperature corrosion would produce a significant number of Mg2+

ions, leading to the creation of Mg(OH)2 and aggressive absorption of a Ca2+ ions.

Similarly, coating a Mg alloy with a very pure Ca-P coating is difficult, because the

binding strength between the specimen and the coating is poor, reducing the implant's

mechanical stability in the body. Several conjugates, polymers, and chemicals are

utilised as precursors or inducers to ramp up the formation of calcium phosphate in

attempt to optimise phase purity, bonding strength, and density of Ca-P coatings

[86-88]. EDTA is an organic compound that promotes HA nanocrystal nucleation and

growth by binding to divalent metallic cation and reacting with calcium ion. On the

other hand, MEA is an aminol that speeds up the synthesis of hydroxyapatite crystal

and increases the solvency of reactants in hydrothermal reactions [89,90]. Because it

contains -NH2, -COOH, and -SH, L-cysteine was discovered to get a decent capacity

to suppress metallic corrosion. Using a hydrothermal technique, one of the

researchers [91] created an L-cysteine CaP coatings on AZ31 which is bioinspired.

Addition of L-cysteine increased corrosion resistance, illustrated in Fig.3, which

depicts the Ca-PL-Cys-coating development mechanism. A hydrothermal technique

was used to coat the exterior layer of Mg alloy with HA, with in the centre, PDA acts

as a glue. PDA's catechol functional group is very sensitive to metal iron absorption,

which helps generate the HA coating by binding Ca2+ ions and subsequently drawing

PO43- to develop hydroxyapatite nuclei, allowing for cross-connection with

hydroxyapatite. The interactions of the polyhydroxyl unit of glucose with Ca2+ ions in

fluid solution may enable the formation of Ca-P crystals on the pure magnesium

surface [92].

Figure 3. Ca-P coating formation mechanism (a-c) and Ca-PL-Cys coating creation

technique (d-f) [92]

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According to one of the researchers [93], glucose can be used as a catalyst to

convert gluconic acid into a carboxyl group under hydrothermal treatment, creating a

negative charged magnesium surface that draws in Ca2+ ions and advancing the

creation of a corrosion protective Ca-P coatings while also significantly reducing the

Mg substrate's anodic dynamics in Hank's solution. The glucose-initiated compound

coating of Mg (OH)2 and Ca-P provoked by hydrothermal treatment of pure Mg is

depicted schematically in Fig. 3. Table 2 lists the characteristics, structure, and

composition of chelating agents utilised in the recent creation of Ca-P coatings on Mg

alloy surfaces using the hydrothermal process.

Table 2. Recent investigations on chelating compounds, related active groups, and coating

characteristics.

Chelating agent

Specimen

Conditio

ns of

hydroth

ermal

reaction

Coordina

tion ion

Active

group

Coating

compositio

Properties

of coating

Ref

Glucose pure Mg 500 mM

glucose,

250 mM

Ca(NO3

)2 • 4H2

250 mM

KH2PO4

120 ◦C,

24 h, pH

= 10

Ca2+ ion Carboxyl

group

transform

ed from

aldehyde

group

Mg(OH)2

DCPA

CDPA

Corrosion

protection is

enhanced by

a denser

coating

morphology

that uses

several large

single

particles as

its building

blocks.

L-cysteine

Mg alloy

AZ31

0.15 g/L

L-

cysteine,

0.25 M

of CaCl2

and

KH2PO4

, 60◦C,

30min

Ca2+ ion

sulfenyl)-

SH(amin

o)-NH2

(carboxyl

)- COOH

Ca(10-x)

(HPO4)x

(PO 4)(6-x)

(OH)2-x

MgHPO4 •

3H2O,

Ca10

(PO4)6(OH

The

morphology

of the

coating has

been greatly

enhanced in

terms of

uniformity

and

integrity.

Coating

thickness

twice that of

the control

group;

enhanced

corrosion

resistance.

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Polydopamine

(PDA)

Mg alloy

AZ31

5 M

NaOH

60 ◦C 3

h 2 mg/

dopamin

hydrochl

oride14

Ca(NO3)

2 8.4mM

NaH2PO

4 4mM

NaHCO

3,393K 4

Ca 2+ ion

Catechol

PDA/HAp

Increased

adhesion,

proliferation

, and

dispersion

osteoblasts;

Denser

coating

structure;

Lower

corrosion

rate

126

Deoxyribonuclei

c acid (DNA)

Mg alloy

AZ31

1g/L

DNA, 14

Ca(NO3)

2 8.4mM

NaH2PO

4 4mM,

NaHCO

3 150

◦C,4h

Ca 2+ ion

Base pair

and

deoxyribo

se double

helix

tricalcium

phosphate

(TCP),

dicalcium

phosphate

anhydrous

(DCPA)

and

calcium-

deﬁcient

hydroxyapa

tite

(CDHA)

By ﬁne-

tuning the

coating

grain,

researchers

can increase

bonding

with sample.

Improve the

resistance to

corrosion

127

Polyacrylic acid

(PAA)

Mg alloy

AZ31

1 M

NaOH

60 °C for

1h, ,14

Ca(NO3)

2 8.4mM

NaH2PO

4 4mM

NaHCO

3 ,90

◦C,4 h

Ca 2+ ion

COO-

Mg(H2PO4

)2; (Ca,

Mg)3

(PO4)2;

and HA

Good

corrosion

resistance;

denser

stereoscopic

blade

structure;

10.69 N

adhesive

force

128

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2.5. ALKALI-HEAT TREATED CONVERSION COATING

To remove oil and other impurities from Mg and Mg alloy alkali heat treatment is an

important step. Coatings are typically used as a pre-treatment to improve corrosion

resistance of adjacent layer to strengthen the bond between outer layer and the

sample. The sample was submerged in an extremely saturated solution of NaHCO3 -

MgCO3 having a pH of 9.3 for 24 hours before being heated at 500 °C for 10 hours

for alkali-heat treatments [94]. The coating in SBF has a high level of corrosion

resistance. Simultaneously, in an early cytotoxicity investigation, it shows no

cytotoxicity by cell development, with no morphological changes on cells or inhibitory

effect. One of the researchers [95] used alkali heat pre-treatment to create LbL

assembled coatings, which boosted the binding force between the sample and LbL

layer and improved corrosion protection. The benefits of alkali heat treatment coating

are its good biocompatibility and excellent corrosion resistance. In any event, the

coatings are exceedingly thiner and therefore could not be utilised as a topcoat, or

even as a primer.

3. MAO COATING

Plasma Electrolytic Oxidation is another name for micro arc oxidation (MAO)

(PEO), is a greener alternative to anodization [96,97]. MAO has lately been popular

Phytic acid (PA)

Mg alloy

AZ31

0.70 wt.

% phytic

acid

,4.90 wt.

Ca(NO3)

2 •4H2O

and 0.89

wt.%

P2O5 at

40 C for

40 min,

pH = 4.5

Ca 2+ ion

COO-

Phytic acid/

24.3 ± 1.7

MPa

bonding

strength;

good

corrosion

resistance

129

Ciproﬂoxacin

hydrochloride

(CIP)/

Polyacrylamide

hydrochloride

(PAH)

Mg alloy

AZ31

5 M

NaOH

30 min

14 mM

Ca(NO3)

2 , 8.4

NaH2PO

4 , 4 mM

NaHCO

3 , 150

◦C, 240

min

Ca 2+ ion COO- CIP/PAH/

HA;

composite

coating

Excellent

corrosion

protections;

CIP

discharge

that may be

controlled;

Antibacteria

l activity is

excellent,

and

cytocompati

bility is

adequate.

130

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for covering magnesium alloys with oxide-ceramic coatings [98]. In the PEO

technique, the cathode is a cylindrical SS container, and the anode is a Mg alloy. Due

to the higher temperature induced by the high voltage, an oxide coating developed on

the outer layer of the Mg alloy at first, and vigorous gas development could be seen.

The voltage then continues to rise, although at a slower rate, and the oxide film is

broken at the weaker portion, with the thickness of the coating gradually increasing.

Due to exceptionally powerful discharges on the surface of the specimen [99],

discharge channels emerged, and the electrolyte fused into these channels. A

combination of high voltage and high temperature in the discharge channel converts a

few of the Mg on the specimen external layer and an electrolyte within channels to

plasma, which is subsequently transformed to plasma through a plasma chemical

process. Metallic ions generated by the magnesium alloy are expelled and moved

away from it, whereas oxygen ion moved in reverse direction. The oxide is

subsequently deposited on the outer layer [100] as a result of an interaction between

magnesium metallic ions and oxygen ions. The applied electric field pushes anions

from the MAO electrolyte, such as SiO32- or PO43-, toward the anode, where they

reacting with Mg2+ions from the discharge channels of Mg specimen [101]. A

ceramics-like covering with better corrosion resistance, bonding strength and wear

resistance is formed by MAO process. Fig. 4 depicts the MAO coatings on Mg alloy

production procedure and structure. At the start of the MAO process, a dense and thin

coating is applied to magnesium alloy. Because of this, the MAO coatings have a two-

layered microstructure with a thick inside layer and a porous outside layer [102].

Figure 4. Schematic example of MAO coating preparation and structure on Mg alloy [96].

The two-layered structure of MAO coatings can be seen using scanning electron

microscopy (SEM). Thermal strains induced by quick solidification of molten oxide in a

gradually cooling electrolyte form small pores, whereas miniature fractures are formed

by thermal stresses caused by quick solidification of molten oxide in a gradually

cooled electrolyte [103]. MAO coating act as physical barriers, successfully isolating

the Mg sample from hostile environments and lowering its corrosion resistance [104].

The corrosive Cl- iron penetrates the exterior layer through micropores during the

initial phase of submersion in physiological environment. Simultaneously, the coating's

main ingredient (MgO) reacts with H2O to form Mg (OH)2 as per Eq. (1) [105]. The

hydrated substance partly fills the pores in the coating, whereas the rest settles on the

outer layer [106]. Notwithstanding its volatile chemical characteristics, Mg (OH)2 could

be converted to dissolvable MgCl2 through Cl-as per Eq. (2), releasing hydroxyl ions

[107]. After then, as per Eq. (3) [105], a Ca-P layer structure arises as the OH-

interacts with various chemicals and ions. During this period, the MAO coating's

breakdown is slowed by the outside porous layer. The corrosion zone expands as the

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submersion duration rises, and the pores get deeper and larger, facilitating electrolyte

infusion into the inward dense layer [108]. The innermost thick layer is useful in

stopping aggressive medium from reaching and touching magnesium substrate, and if

this layer is removed, corrosive medium would penetrate and touch the magnesium

substrate. Because of its high inherent corrosion characteristics in aqueous solution,

Mg corrosion occurs quickly when the electrolyte meets the substrate [109]. During

this time, the MAO coatings' corrosion resistance is primarily dependent on the interior

dense layer. The MAO coating's two-layered structure, as illustrated above, can

greatly increase the corrosion protection of Mg alloys. Figure 3 depicts the

biodegradation process of the MAO coating on Mg alloy.

Figure 5. MAO coating on Mg alloy biodegradation process [96]

MgO + H2OMg (OH)2 (1)

Mg (OH)2 + 2Cl- =MgCl2 + 2OH (2)

Ca2+ + HPO2-4 + OH- = Ca - Player + H2O (3)

3.1. ANTIBACTERIAL MAO COATINGS ON MG ALLOYS

Mg's antibacterial capabilities have been revealed in a number of previous studies

[110,110]. The results demonstrated that Mg was effective towards methicillin-resistant

Staphylococcus aureus when it was injected into rats with implant-related infection

(MRSA). One of the researchers used pure Mg, glass slides, and polyurethane stents

to grow Escherichia coli (E. coli). After 16 hours of coculturing with pure Mg, the

bacterial cell density was at its lowest, indicating that Mg has a very powerful

antibacterial impact. Mg has also been shown to have outstanding antibacterial

properties in the past [112]. The alkaline rise caused by Mg's breakdown in solution is

responsible for its antibacterial effect [113]. It was observed that AZ31 and pure Mg

along with silicon and fluorine coatings created by chemical conversion technique lost

their antibacterial activities because of its dense coating on the surface [114]. Coating

Mg alloy with antibacterial coatings, in this way, could be a potential methodology for

increasing antibacterial effects and corrosion resistance of the magnesium alloy

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concurrently. For development of antibacterial coating on the metal surface MAO

technique is effective, economical and extensively used method [115].

The MAO approach produces coatings that are largely made up of substrate

oxides, but they also contain electrolyte components [116]. As a result, altering the

electrolyte composition can efficiently vary the composition of the MAO coating [117].

Furthermore, significant antimicrobial activity is known for Zn, Cu, and Ag, and these

antibacterial metallic components kill bacteria by denatureing proteins, damaging

membranes, and generating oxidative stress [118-121]. Antibacterial coatings can be

made using this metal complex or by merely putting micro/nanoparticles of these

metals to MAO electrolyte [122].

4. BIOCOMPATIBILITY OF CONVERSION COATINGS

CaP type coatings have good biocompatibility, according to the cell survival test

findings shown in Fig. 6, which can be attributed to their DCPD content.

Figure 6. After 1, 3, and 5 days of incubation, cell feasibility of all treated specimen and

bare AZ31 [123].

Because of its outstanding biocompatibility and biodegradability [125], DCPD is an

essential biomaterial in the field of bone cement [123,124], and its degradation

products can supply vital calcium and phosphorus supplies for bone tissue

regeneration. The effects of CaP, ZnP, and MgP coating types on the precipitation of

Ca3(PO4)2 and Mg3(PO4)2 have not been thoroughly investigated.

According to the different morphologies of all conversion coating after a 15-day

soaking, there are more precipitations on CaP and MgP type conversion coatings than

on ZnP type conversion coatings. The divergence of phosphates, magnesium, and

calcium precipitation on diverse types of coatings can be explained by two factors:

1. CaP type and MgP type coatings have a higher solubility than ZnP type coatings.

As a result, the alkalization of the Hanks' solution near the magnesium complex

sample and the precipitation of phosphates magnesium and calcium is accelerated by

the breakdown of conversion coatings.

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2. The CaP and MgP coatings have the same composition as the phosphate

precipitation from Hanks' solution, that minimises the principal factor driving

precipitation nucleation and development and, as a result, speeds up precipitation

during the immersion test [123].

5. CONCLUSION

In vivo corrosion of biomedical magnesium can be controlled by alloying

magnesium with by alloying it with biocompatible elements, by changing surface

microstructure and applying protective coating to isolate implant and body fluid.

Because of the lower cost and easiness in operation conversion coating can be

generally utilized in biomedical fields. Chemical conversion coatings are created by

interaction of precipitation and chemical dissolution. Fluoride conversion coating can

give safety during the initial phases of implantation due to its thin coating, however it

must be employed for long-term applications. Fluoride leak during Mg implants

corrosion and its cytotoxicity for tissue are unknown, and it would have a negative

impact on bone, necessitating additional investigation. Alkali heat treatment coating

has good biocompatibility and excellent corrosion resistance; however, coatings are

extremely thin and cannot be used as topcoat.

It was observed that Zn, Cu and Ag have excellent antimicrobial capacity, and

these antibacterial metallic components kill microorganisms through protein denature,

membrane destroy and oxidative stress. MAO is another technique used for bioactive

and antibacterial coating on magnesium and its alloys. Antibacterial coatings can be

made by mixing micro/nanoparticles of these metals with MAO electrolyte.

ACKNOWLEDGEMENTS

For the research, authoring, and/or publication of this article, the author(s) received

no financial funding.

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