*3.2. Titanium Dioxide (TiO2)*

TiO<sup>2</sup> coatings are the most important materials in biomedical applications that are known for their antibacterial properties along with good mechanical properties. The applications of TiO<sup>2</sup> coating include drug delivery systems [155], orthopedic [156], and dental applications [157]. It also shows high catalytic activity, antibacterial activity, and long-term stability under photo and chemical corrosion [158]. TiO<sup>2</sup> promotes the formation of bone-like apatite or calcium phosphate on its surface. This property makes it a suitable candidate for reconstruction and bone replacement [159].

TiO<sup>2</sup> coated metallic substrates show better antibacterial properties. Gartner et al. [160] observed the same biocidal effect by applying TiO<sup>2</sup> coating on glass substrates by a sol-gel method. Photocatalytic activity of TiO<sup>2</sup> coating received much attention as a potential material for anti-bacterial coatings. The antibacterial effects of TiO<sup>2</sup> coating involve both a reduction in bacteria's viability and their destruction [161]. Park et al. [162] showed that the

1

antibacterial effect against *S. aureus* could be improved by adjusting the nucleation time of TiO<sup>2</sup> film during the deposition process. The antibacterial effect of TiO<sup>2</sup> was explained by the formation of reactive oxygen species. Apart from antibacterial properties, the antiviral properties of the TiO<sup>2</sup> coating are also studied [163]. Table 4 summarizes the use of TiO<sup>2</sup> and its composite coatings for bioimplant applications.

Yetim [164] prepared TiO<sup>2</sup> coating with different concentrations of Ag using the sol-gel process on the commercially pure titanium substrate. Electrochemical corrosion properties obtained from electrochemical impedance spectroscopy measurements and potentiodynamic polarization tests in simulated body fluid (SBF) suggest that Ag doped TiO<sup>2</sup> enhances corrosion resistance over that of the bare Ti substrates as well as undoped TiO<sup>2</sup> coated samples [165]. The silver doped TiO<sup>2</sup> (Ag/TiO2) nanocomposite coated glass substrate with varying Ag content synthesized by the sol-gel route showed antiviral properties against *E. coli*, enterovirus, and influenza A virus (H1N1) [166]. The highest level of photocatalytic degradation under irradiation with either visible or ultraviolet light was observed at an optimum Ag:TiO<sup>2</sup> weight ratio of 1:100. The antibacterial effectiveness was greater than 99.99% against *E. coli* and other infectious diseases after visible light illumination.

Sol-gel derived TiO2-PTFE nanocomposite coating on stainless steel substrates was prepared by Zhang et al. [164] and their bacterial adherence were tested against two pathogens, namely *S. aureus* and *E. coli*. The bacterial adhesion and bacterial growth studies were evaluated by fluorescence microscopy after 2 h, 6 h, 12 h, and 24 h of incubation (Figure 2a,b). The TiO2-PTFE coated substrate shows the lowest bacterial adhesion when compared with the uncoated substrate. The bacterial inhibition increases with the increasing TiO<sup>2</sup> concentration (Figure 2c,d). It is also observed that Gram-positive bacteria are less sensitive due to their cell wall thickness.

**Figure 2.** Effect of bacterial adhesion (**a**,**b**) and bacterial growth of *E. coli* and *S. aureus* pathogens on TiO<sup>2</sup> -PTFE coated and uncoated substrates [164]. TiO<sup>2</sup> -PTFE coated substrates exhibit lower bacterial adherence and a significant reduction in bacterial growth (**c**,**d**) as compared to uncoated substrates.


**Table 4.** Uses of TiO<sup>2</sup> and its composite coatings in bioimplant applications.

#### *3.3. Transition Metal Nitrides*

Earlier, transition metal nitrides and carbides were widely used to protect the metallic components against wear, tear, and corrosion, potentially offering high-temperature stability. Titanium nitride (TiN) coatings were used as decorative coatings in earlier days. In the last decade, nitride coatings for orthopedic implants were also proposed to protect the implants against wear and tear and to act as a diffusion barrier layer preventing the toxic ion release from the implant metal surfaces to the human body fluids [177–180]. The physical properties of TiN coated substrates show high scratch resistance, hardness, and low frictional coefficients. These properties make them a potential candidate for use as coatings on different metals used in arthroplasty. TiN-based coatings used for orthopedic applications show better biological properties as compared to other nitrides [181]. TiN coatings show better blood tolerability properties with a hemolysis percentage near zero [182]. TiAlN is another biocompatible nitride that has proven to be a promising alternative to TiN in biomedical applications despite its aluminum (Al) content [183].

Transition metal carbonitrides (TiCN, ZrCN) were found to increase the service life of orthopedic implants in terms of wear resistance in biological media [184–186]. Recently, quaternary carbonitrides-based coatings (TiAlCN, TiCrCN, TiNbCN, etc.) were found to show increased anticorrosive, mechanical, and tribological properties compared to ternary carbonitride-based coatings [187–189]. The tribological properties of these carbonitride coatings are very complex. However, the carbon-based carbonitride coatings show good biocompatibility, better wear resistance, and low friction [190]. Much attention has been paid to developing MeSiC-, MeSiCN-, and MeSiN- (where Me is a transition metal, and Si is an alloying element) based hard coatings [191–194]. These types of coatings show high thermal stability, a low frictional coefficient, excellent wear resistance, and good mechanical properties (hardness, Young's modulus). Moreover, in many investigations, TiSi-based carbide and carbonitride coatings proved to be a potential candidate for a metallic implant

which combines the mechanical, tribological, and anticorrosive properties of TiN and TiC with the biocompatibility behavior of SiC and SiCN [192,195–197]. implant which combines the mechanical, tribological, and anticorrosive properties of TiN and TiC with the biocompatibility behavior of SiC and SiCN [192,195–197].

high thermal stability, a low frictional coefficient, excellent wear resistance, and good mechanical properties (hardness, Young's modulus). Moreover, in many investigations, TiSibased carbide and carbonitride coatings proved to be a potential candidate for a metallic

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TiN coating shows plastic deformation at the coating/surface interfaces due to dissimilarities in the hardness of the substrate and coating [198]. Thus, TiN coating cannot accommodate the fracture and deformation that creates flakes, and defects in the coatings cause deterioration of the coatings from the substrate. Therefore, chromium nitride (CrN) and chromium carbonitride (CrCN) coatings are recommended, which act as a better diffusion barrier for ion release from the alloys. These coatings also exhibit higher toughness, higher cohesive strength, and lower wear debris than TiN coatings [199]. TiN coating shows plastic deformation at the coating/surface interfaces due to dissimilarities in the hardness of the substrate and coating [198]. Thus, TiN coating cannot accommodate the fracture and deformation that creates flakes, and defects in the coatings cause deterioration of the coatings from the substrate. Therefore, chromium nitride (CrN) and chromium carbonitride (CrCN) coatings are recommended, which act as a better diffusion barrier for ion release from the alloys. These coatings also exhibit higher toughness, higher cohesive strength, and lower wear debris than TiN coatings [199]. TiN and TiCuN coatings were prepared by the axial magnetic field enhanced arc ion

TiN and TiCuN coatings were prepared by the axial magnetic field enhanced arc ion plating (AMFE-AIP) technique, and the in vitro angiogenic response of human umbilical vein endothelial cells was studied by Liu et al. [200]. The TiCuN coating showed better antibacterial activity, and both coatings showed no cytotoxicity to human umbilical vein endothelial cells (HUVECs). TiCuN coatings promote early cell apoptosis, which is important for vascular tissue modeling (Figure 3). plating (AMFE-AIP) technique, and the in vitro angiogenic response of human umbilical vein endothelial cells was studied by Liu et al. [200]. The TiCuN coating showed better antibacterial activity, and both coatings showed no cytotoxicity to human umbilical vein endothelial cells (HUVECs). TiCuN coatings promote early cell apoptosis, which is important for vascular tissue modeling (Figure 3).

Figure 3. Apoptosis rate of TiN and TiCuN coatings tested for Day 1 and Day 3. Annexin V-FITC/PI double staining kit was used to evaluate the apoptosis rate of these coatings [200]. TiCuN coating promoted the early cell apoptosis rate more than TiN coating. \*: Denotes TiCuN coating superior performance **Figure 3.** Apoptosis rate of TiN and TiCuN coatings tested for Day 1 and Day 3. Annexin V-FITC/PI double staining kit was used to evaluate the apoptosis rate of these coatings [200]. TiCuN coating promoted the early cell apoptosis rate more than TiN coating. \*: Denotes TiCuN coating superior performance.

Transition metal oxynitrides have been considered as interesting materials due to their known mechanical properties, chemical stability, and corrosion resistance in simulated body fluid. Zirconium oxynitride (ZrON) and titanium oxynitride (TiON) based coatings were recently used in biomedical applications for their better corrosion resistance than TiN coating and their anti-fouling ability [201,202]. The magnetron sputtered ZrON and TiON coated 316L SS specimen show better hardness and wear resistance behavior than the uncoated substrate [203]. In addition, both coatings show better anti-fouling performance against Pseudomonas aeruginosa bacterial adhesion than uncoated substrates. The coated substrates also show better corrosion protection with or without the addition of hydrogen peroxide (H2O2) in artificial blood plasma (ABP) solution [203]. Surface modified coatings prepared from ternary nitrides such as TiZrN, TiCrN, and Transition metal oxynitrides have been considered as interesting materials due to their known mechanical properties, chemical stability, and corrosion resistance in simulated body fluid. Zirconium oxynitride (ZrON) and titanium oxynitride (TiON) based coatings were recently used in biomedical applications for their better corrosion resistance than TiN coating and their anti-fouling ability [201,202]. The magnetron sputtered ZrON and TiON coated 316L SS specimen show better hardness and wear resistance behavior than the uncoated substrate [203]. In addition, both coatings show better anti-fouling performance against *Pseudomonas aeruginosa* bacterial adhesion than uncoated substrates. The coated substrates also show better corrosion protection with or without the addition of hydrogen peroxide (H2O2) in artificial blood plasma (ABP) solution [203].

TiAlN gained considerable attention because they retain their physiochemical properties, Surface modified coatings prepared from ternary nitrides such as TiZrN, TiCrN, and TiAlN gained considerable attention because they retain their physiochemical properties, such as oxidation resistance, hardness, corrosion resistance, biocompatibility, and structural stability after implantation [204,205]. Magnetron sputtered TiZrN coated 316L SS substrates showed less bacterial adhesion, increased corrosion protection, and negligible human blood platelets activity than uncoated substrates [206]. Recent developments in binary, ternary, and quaternary systems of transition metal nitrides and carbide coatings are tabulated in Table 5.


**Table 5.** Recent work on binary, ternary, and quaternary systems of transition metal nitride and carbide coatings for implant applications.

#### *3.4. Carbon Based Coatings*

Carbon based materials are categorized under bioinert coatings. These coatings are used in load-bearing applications and wear components to improve elevated corrosion resistance, wear, and frictional effects [219]. Besides, carbon-based coatings show minimum protein adhesion and very good biocompatibility due to the hydrophobic nature of carboncoated surfaces. Three different types of carbon-based coatings are used for biomedical applications. They are (a) nanocrystalline diamond (NCD), (b) pyrolytic carbon (PyC), and (c) diamond-like carbon (DLC) [220]. Some of the coatings are commercially available, while others are under development.

Most of the PyC coatings in biomedical applications are found in the heart valves due to their thromboresistant qualities and biocompatibility [221]. Most of the artificial heart valves are lined with a thick PyC coating. PyC biocompatibility in heart valves is well established. PyC coatings have also been used in orthopedic applications [222]. By varying the process parameters of the PyC (such as temperature, surface area, gas flow rate, precursor) in the CVD process, a variety of the structures can be produced. The most interesting structure for biomedical applications is lamellar, isotropic, granular, and columnar [223–225]. PyC coated orthopedic implants are used to replace small joints such as wrist joints, knuckles, and arthroplasty of proximal interphalangeal joints [226].

Carbon coatings, including nanocrystalline diamond and DLC coating, show many remarkable biological properties and are considered as coatings for medical implants. NCD coatings deposited by the CVD process consist of sp<sup>3</sup> -hybridized carbon bonds and show grain sizes in the range of a few nanometers. NCD coatings generally show very low surface roughness and possess the properties of a diamond, such as hydrophobicity and excellent biocompatibility with blood [227,228]. This makes them an ideal coating choice for wear-resistant implant applications and cardiovascular devices. NCD coating can also be used as hard antibacterial coatings that reduce the risk of infections. The electrically active NCD coating surfaces can establish a chemical bond with the biomolecules in the surrounding environment. Medina et al. [229] observed that the NCD coating surfaces react with the cell wall or membrane of Gram-negative *P. aeruginosa* bacteria and establish a chemical bond that alters the bacteria morphology, hindering bacterial adhesion and colonization on the surface of the coating. The properties of NCD films are utilized in biosensing and neurochemical sensing applications [230]. and excellent biocompatibility with blood [227,228]. This makes them an ideal coating choice for wear-resistant implant applications and cardiovascular devices. NCD coating can also be used as hard antibacterial coatings that reduce the risk of infections. The electrically active NCD coating surfaces can establish a chemical bond with the biomolecules in the surrounding environment. Medina et al. [229] observed that the NCD coating surfaces react with the cell wall or membrane of Gram-negative P. aeruginosa bacteria and establish a chemical bond that alters the bacteria morphology, hindering bacterial adhesion and colonization on the surface of the coating. The properties of NCD films are uti-

joints, knuckles, and arthroplasty of proximal interphalangeal joints [226].

[223–225]. PyC coated orthopedic implants are used to replace small joints such as wrist

Carbon coatings, including nanocrystalline diamond and DLC coating, show many remarkable biological properties and are considered as coatings for medical implants. NCD coatings deposited by the CVD process consist of sp3-hybridized carbon bonds and show grain sizes in the range of a few nanometers. NCD coatings generally show very low surface roughness and possess the properties of a diamond, such as hydrophobicity

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More experimental studies have been reported on DLC based coatings, which are considered as the most promising materials for bioimplant applications [231–234]. Medical grade PEEK samples were coated with DLC using plasma immersion ion implantation and deposition (PIII & D) technique, and their in vitro cytocompatibility and osteogenesis studies were carried out by Mo et al. using human bone marrow mesenchymal stem cells (hBMSCs) [235]. DLC coated substrates show better surface coverage of cells and show high cell viability on the seventh day, which indicates better biocompatibility of DLC-PEEK coatings than PEEK coating (Figure 4). However, DLC suffers from residual stress arising from the substrate/coating thermal expansion mismatch and lattice misfit, which cause poor substrate adhesion and delamination of the coating from the substrate. Another major concern about DLC coatings is their instability in the aqueous environment, which promotes delamination of the coating [236]. To avoid this issue, it is recommended to use interlayers (called buffer layer) such as CrC, Ti, and Si3N<sup>4</sup> at the interface of the substrate and DLC coating [237]. Another approach is to dope DLC coating with N, F, Ag, Zr, or Ti to avoid a thermal expansion mismatch and residual stress [238]. lized in biosensing and neurochemical sensing applications [230]. More experimental studies have been reported on DLC based coatings, which are considered as the most promising materials for bioimplant applications [231–234]. Medical grade PEEK samples were coated with DLC using plasma immersion ion implantation and deposition (PIII & D) technique, and their in vitro cytocompatibility and osteogenesis studies were carried out by Mo et al. using human bone marrow mesenchymal stem cells (hBMSCs) [235]. DLC coated substrates show better surface coverage of cells and show high cell viability on the seventh day, which indicates better biocompatibility of DLC-PEEK coatings than PEEK coating (Figure 4). However, DLC suffers from residual stress arising from the substrate/coating thermal expansion mismatch and lattice misfit, which cause poor substrate adhesion and delamination of the coating from the substrate. Another major concern about DLC coatings is their instability in the aqueous environment, which promotes delamination of the coating [236]. To avoid this issue, it is recommended to use interlayers (called buffer layer) such as CrC, Ti, and Si3N4 at the interface of the substrate and DLC coating [237]. Another approach is to dope DLC coating with N, F, Ag, Zr, or Ti to avoid a thermal expansion mismatch and residual stress [238].

Figure 4. Represents in vitro cytocompatibility of DLC coated PEEK substrates, (a) surface morphology of hBMSC cultured on PEEK, DLC-PEEK, and NH2-DLC-PEEK substrates for 1 day, and the enlarged cells are shown in pseudo-color, (b) cell viability for 1 day, and (c) proliferation of hBMSCs after culturing samples for 1 d, 3 d, and 7 days [235]. \*\* denotes p < 0.01 **Figure 4.** Represents in vitro cytocompatibility of DLC coated PEEK substrates, (**a**) surface morphology of hBMSC cultured on PEEK, DLC-PEEK, and NH2-DLC-PEEK substrates for 1 day, and the enlarged cells are shown in pseudo-color, (**b**) cell viability for 1 day, and (**c**) proliferation of hBMSCs after culturing samples for 1 d, 3 d, and 7 days [235]. \*\* denotes *p* < 0.01.

The properties of DLC such as chemical inertness, surface smoothness, and hydrophobicity are important for providing better compatibility with blood, reducing platelet activation in contact with the blood, which could trigger thrombosis. DLC can act as a protective coating under the conditions of the human blood environment, which limits the release of nickel ions from metallic implants such as SS 316L. Several studies suggest that DLC coating prepared by various routes is biocompatible and does not induce any inflammation reaction both under in vivo and in vitro conditions [235,239]. Because of these remarkable features, DLC coatings found various applications as coatings in many implant devices such as cardiovascular stents, heart valves, surgery needles, medical wires, contact lenses, etc. DLC coatings can also be used as protective coating in knee replacement because of

their high corrosion resistance, hardness, and low wear rate. Generally, DLC films are used to reduce the frictional coefficient and offer better wear resistance [238]. Carbon based coatings and their significance in biomedical field are summarized in Table 6.


**Table 6.** Different carbon coatings and their properties.

#### *3.5. Calcium Phosphates*

Calcium phosphate (CaP) ceramics are widely used as implants since they have a chemical composition similar to the inorganic composition of the bone. By controlling the surface properties such as roughness and porosity of CaP, one can regulate the biomineral formation and cell/protein adhesion. Bioactivity properties are varied depending on the type of calcium phosphates (HAp, tricalcium phosphate (TCP)) because of the differences in crystallinity, solubility, stability, ion release, and mechanical properties. At first, CaP coatings were deposited through the vapor phase process, but in recent years, biomimetic and solution-based methods were developed. Each synthesis approach has its own intrinsic properties, but in general, CaP based coatings are promising to improve implant longevity and biocompatibility. Many studies have been focused on the development of CaP ceramic coatings on metallic substrates to achieve the biological properties identical to a bulk and to enhance the implant durability and fixation [250–253].

Presently, atmospheric plasma spraying (APS) is currently employed to develop CaP coating on implant surfaces [2]. The CaP phases in the coatings exhibit higher solubility in an aqueous medium than HAp which is desirable for activating bone formation. However, faster dissolution reduces the stability and can cause loosening of the implant. A highly crystalline HAp phase dissolves in human physiological conditions at a lower rate which provides long-term stability of the implants. Thus, for the development of implants with required properties, one must control the purity and crystallinity of the coatings. CaP coatings with a denser microstructure lower the risk of delamination of the coating during in vivo tests with human body fluids. Coating surface roughness affects its dissolution and bone apposition and growth. Porous surfaces may enhance cell attachment or formation of the extra-cellular matrix, but the accumulation of macropores at the coating/substrate interface weakens the coating adhesion [254].

CaP in the form of HAp is widely used in implant applications due to its superior biological response. The HAp composition is Ca10(PO4)6(OH)<sup>2</sup> (Ca/P = 1.67), which resembles the chemical composition of hard tissues such as bone and teeth [255]. Hence, HAp is considered as a primary candidate material due to its exceptional biological properties such as excellent biocompatibility, osteoconductivity, osteoinductivity, and bioactivity [256]. HAp coatings release calcium and phosphate ions and regulate the activation of osteoclasts and osteoblasts, facilitating bone regeneration [257]. The use of HAp ceramics enhances the regeneration of bones, improves osteoconductivity for bone growth, and promotes mineralization through ion release control and encapsulating growth factors. HAp ceramic coating enhances bone apposition in orthopedic implants through the formation of an extremely thin bonding layer with the existing bone. Due to such tissue bonding characteristics, Hap-based ceramics are considered as bioactive-based coatings. The continuous effort to improve the durability of the HAp ceramic coatings has led to development of high-quality HAp coatings and the development of Hap-based composite coatings.

Highly porous or highly crystalline HAp coating shows poor adhesion to the substrate. Sankar et al. [258] studied the corrosion behavior of HAp coatings prepared by electrophoretic deposition (EPD) and the pulsed laser deposition (PLD) method. The corrosion results suggest that the HAp coatings show lower corrosion protection than the coatings prepared by the PLD method due to the formation of denser and pore-free coating [258]. Corrosion protection can also be enhanced by the addition of antimicrobial dopants. For example, Yugeswaran et al. [259] prepared HAp-TiO<sup>2</sup> nanocomposite coatings by APS. The coating shows better corrosion performance in SBF medium than HAp coating without dopants due to its high compactness and the presence of TiO<sup>2</sup> [259]. Silver (Ag) containing HAp coatings prepared by Trujillo et al. [260] show better antibacterial activity than HAp coating alone against *P. aeruginosa* and *S. epidermidis* pathogens due to the antibacterial activity of Ag. The antimicrobial activity of the Ag-doped HAp composite against *E. coli* and *S. aureus* was tested by Lett et al. [261]. The results indicated that the Ag-doped HAp composite has better inhibition of bacterial growth and shows a stronger ability against *S. aureus* bacteria to fight against toxic responses (Figure 5). The absence of Ag in the composite results in lower antibacterial activity of HAp composites. The variation in antibacterial activity was attributed to a thinner cell wall response of *S. aureus* (Figure 5b) to Ag ions than *E. coli* (Figure 5a) [261]. Gels 2022, 8, x FOR PEER REVIEW 19 of 38

Figure 5. Demonstration of antimicrobial activity of HAp and Ag doped HAp composites against E. coli (Gram-negative) and S. aureus (Gram-positive) bacteria [261]. The photograph shows that Ag-doped HAp inhibits S. aureus bacteria more effectively than E. coli. **Figure 5.** Demonstration of antimicrobial activity of HAp and Ag doped HAp composites against *E. coli* (Gram-negative) (**a**) and *S. aureus* (Gram-positive) (**b**) bacteria [261]. The photograph shows that Ag-doped HAp inhibits *S. aureus* bacteria more effectively than *E. coli*.

In biomedical implants, the major challenge for the performance of implants is bacterial invasion. During surgical operation, the bacteria may enter the surface of the implants through surgical equipment or cross contamination which form a biofilm. Once the surrounding implant is infected, the infection causes implant loosening. To overcome this issue, antimicrobial agents are used as dopants in ceramics, protecting implant material from bacterial invasion and improving their durability. Zinc doped HAp composites prepared by the sol-gel route and annealed at different temperatures (500 °C and 700 °C) In biomedical implants, the major challenge for the performance of implants is bacterial invasion. During surgical operation, the bacteria may enter the surface of the implants through surgical equipment or cross contamination which form a biofilm. Once the surrounding implant is infected, the infection causes implant loosening. To overcome this issue, antimicrobial agents are used as dopants in ceramics, protecting implant material from bacterial invasion and improving their durability. Zinc doped HAp composites pre-

show higher antimicrobial activity against C. albicans fungal cells and S. aureus bacteria

F- doped hydroxyapatite and studied the properties of the coating. The addition of the dopant improves the structural stability of the HAp lattice and promotes osteogenic cell differentiation. Moreover, the addition of F- ions potentially arrests the formation of S. aureus. Dopants such as Cu, Zn, Mg, Ag added to HAp enhance antibacterial activity and decrease the toxic effects towards the human body cells [264–267]. For example, Mg-

Multiple doping of ions into HAp coatings was also attempted to improve their struc-

The differences in the thermal expansion coefficient of HAp and metallic alloys result in residual thermal stress. The stress accumulation increases with the increase in the coating thickness, which promotes cracking or delamination of the coating. For a thicker coating, the outer layer may detach from the implant, whereas a thin HAp coating can prematurely resorb during bone regeneration. Various HAp composites and their biological

the osteoblast cells [270]

layer [272]

[271]

Better antibacterial efficacy against Escherichia coli and Staphylococcus aureus bacteria than HAp coating

properties are summarized in Table 7.

on Ti surface Biomimetic method

doped HAp shows better osteoblast cell adhesion than pure HAp [268].

Table 7. Hydroxyapatite and its composites' coatings for implant applications.

substrate Precipitation method Favorable regeneration of crystalline Si-HA

substrate Solvothermal method Excellent apatite-forming ability [269]

S. No. Coatings Deposition Method Significance Ref.

Fe doped HAp on Si substrate Co-precipitation method Promote better proliferation and adhesion of

[262].

HAp nanowire coating on glass

Ce doped HAp/collagen coating

Si substituted HAp coating on Ti

pared by the sol-gel route and annealed at different temperatures (500 ◦C and 700 ◦C) show higher antimicrobial activity against *C. albicans* fungal cells and *S. aureus* bacteria [262].

Multiple doping of ions into HAp coatings was also attempted to improve their structural stability, partial dissolution, and biocompatibility. Wang et al. [263] prepared Sr and F− doped hydroxyapatite and studied the properties of the coating. The addition of the dopant improves the structural stability of the HAp lattice and promotes osteogenic cell differentiation. Moreover, the addition of F− ions potentially arrests the formation of *S. aureus*. Dopants such as Cu, Zn, Mg, Ag added to HAp enhance antibacterial activity and decrease the toxic effects towards the human body cells [264–267]. For example, Mg-doped HAp shows better osteoblast cell adhesion than pure HAp [268].

The differences in the thermal expansion coefficient of HAp and metallic alloys result in residual thermal stress. The stress accumulation increases with the increase in the coating thickness, which promotes cracking or delamination of the coating. For a thicker coating, the outer layer may detach from the implant, whereas a thin HAp coating can prematurely resorb during bone regeneration. Various HAp composites and their biological properties are summarized in Table 7.

**Table 7.** Hydroxyapatite and its composites' coatings for implant applications.


#### *3.6. Zirconia*

Zirconia (ZrO2) is a ceramic material that can withstand high temperatures as well as higher stresses. It has widespread applications in dental implants and in the coatings on metallic implants to increase their corrosion resistance [280]. ZrO<sup>2</sup> ceramics offer many advantages, including mechanical strength, chemical stability, biocompatibility, good aesthetics, and better wear resistance. Zirconia stabilized with yttria (YSZ) has been used as a dental implant due to its excellent mechanical strength and fracture toughness [281]. YSZ coatings show better hardness and scratch resistance than HAp coating [282]. Gobi Saravanan et al. [283] observed that the YSZ coated Ti substrates show improved hemocompatibility, activating blood platelets with pseudopods. In addition to that, superior in vitro biomineralization behavior was observed and documented through the weight gain on YSZ coating.

Zirconia stabilized with different weight fractions (0, 4, 10 wt. %) of yttria yields different phases (monoclinic, tetragonal, and cubic): zirconia ceramics with tailored mechanical

properties and biocompatibility can be thus prepared. Attempts were made to deposit different phases of zirconia (*m*-ZrO2, *t*-ZrO2, and *c*-ZrO2) with the use of electron beam physical vapor deposition (EBPVD) [284]. All the coatings show lower surface roughness than coating prepared through the APS method and reduce pathogen bacterial invasion. Particularly, *t*-ZrO<sup>2</sup> shows superior hardness over the other two zirconia phases. All the allotropes show better blood plasma protein adhesion and enhanced resistance to corrosion in comparison to uncoated medical grade stainless steel substrates in ABP solution.

Antibacterial activity of ZrO<sup>2</sup> coating can be enhanced by the addition of Ag. Ag-ZrO<sup>2</sup> composite coatings were prepared by Pradhaban et al. [285]. The results suggest that the coating shows antimicrobial activity against *E. coli*. Santos et al. [286] prepared glass ceramic composites with different concentrations of ZrO<sup>2</sup> particles (0–50 vol. %) and carried out a ball-on-plate tribology test. ZrO<sup>2</sup> glass ceramic composite (30 vol. % of ZrO2) shows optimal wear properties (coefficient of friction is 0.3) and is recommended for load-bearing applications. Bermi et al. [287] deposited YSZ coating through pulsed plasma deposition, and the tribological behavior of the coating in both dry and wet conditions was tested. YSZ coating deposited on a Ti6Al4V alloy ball sliding against the UHMWPE disk shows a reduction in wear rate (17% and 4% in dry and lubricated conditions) than uncoated alloy substrate.

Kaliaraj et al. [288] prepared zirconia coatings on a 316L SS substrate by electron beam physical vapor deposition (EBPVD), and a bacterial adhesion study with *P. aeruginosa* was carried out. Epifluorescence microscopy analysis of live/dead cells after incubation of 1, 2, 3, and 4 days showed a drastic reduction in bacterial adhesion on ZrO<sup>2</sup> coatings, along with retardation in biofilm formation (Figure 6). This observation was attributed to the decrease in surface roughness obtained through coating deposition and the surface chemistry of ZrO<sup>2</sup> that inhibits bacterial adhesion. Electrochemical impedance corrosion results show that ZrO<sup>2</sup> exhibited superior corrosion resistance in the presence of H2O<sup>2</sup> in an artificial blood plasma electrolyte solution. [288]. *Gels* **2022**, *8*, x FOR PEER REVIEW 22 of 38

**Figure 6.** Epifluorescence microscopy analysis of P. aeruginosa bacterial invasion on 316L SS (**a**–**d**) and ZrO2 film (**e**–**h**) after 1, 2, 3, and 4 days incubation [288]. The used acridine orange staining shows orange color for live cells and green color for dead cells. The reduction in bacterial adhesion was seen on ZrO2 coated substrate compared to uncoated 316L SS. **Figure 6.** Epifluorescence microscopy analysis of P. aeruginosa bacterial invasion on 316L SS (**a**–**d**) and ZrO<sup>2</sup> film (**e**–**h**) after 1, 2, 3, and 4 days incubation [288]. The used acridine orange staining shows orange color for live cells and green color for dead cells. The reduction in bacterial adhesion was seen on ZrO<sup>2</sup> coated substrate compared to uncoated 316L SS.

Hench pioneered bioactive materials research and revolutionized the fields of bioactive materials and ceramics with his discovery of bioactive glass (45S5 composition), commercially known as Bioglass [289]. In the wake of Bioglass, various compositions and com-

of their amorphous or semi-crystalline nature, they often fail as an implant material due to their poor mechanical strength. To overcome the shortage in mechanical properties, bioactive glasses are often composited with various metal oxides such as TiO2, Al2O3, ZrO2, and 2-D materials such as graphene and its derivatives (graphene oxide and reduced graphene oxide) [290]. These composites were reported to improve the corrosion resistance, antibacterial activity, and angiogenic properties of bioactive glass coatings without losing the bioactivity [291]. Similar to many ceramic materials, bioactive glasses can also be prepared in the form of particles of nano and micron size, as mesoporous par-

In this section, various types of coating technologies that can be used for the coating of bioactive glasses and their composites on different types of metals, alloys, and certain specific surfaces are discussed. One of the most simple and economical coating processes is the sol-gel dip-coating process. However, the coatings are often porous because of the solvent evaporation leading to poor corrosion resistance and mechanical properties. Nevertheless, this problem can be solved by incorporating metal oxides such as B2O3 as reported by Pinki Dey et al. [293]. According to their report, by replacing the silica weight percentage in the 45S5 system by 1% to 5 wt. %, they were able to decrease the porosity in the particles. Thermal spray coating, an industrial coating process, can also be employed for bioactive glass coating preparation. This process involves the coating of bioactive glasses as fine droplets or as plasma and sprayed over metal surfaces. Porous and non-

ticles, fibers, 3D scaffolds or monoliths, and thin films or coatings [292].

*3.7. Bioactive Glass Coatings* 

#### *3.7. Bioactive Glass Coatings*

Hench pioneered bioactive materials research and revolutionized the fields of bioactive materials and ceramics with his discovery of bioactive glass (45S5 composition), commercially known as Bioglass [289]. In the wake of Bioglass, various compositions and composites of bioactive glasses or silicates prepared both by melt quench and sol-gel techniques were investigated. Although bioactive glasses exhibit excellent bioactivity, because of their amorphous or semi-crystalline nature, they often fail as an implant material due to their poor mechanical strength. To overcome the shortage in mechanical properties, bioactive glasses are often composited with various metal oxides such as TiO2, Al2O3, ZrO2, and 2-D materials such as graphene and its derivatives (graphene oxide and reduced graphene oxide) [290]. These composites were reported to improve the corrosion resistance, antibacterial activity, and angiogenic properties of bioactive glass coatings without losing the bioactivity [291]. Similar to many ceramic materials, bioactive glasses can also be prepared in the form of particles of nano and micron size, as mesoporous particles, fibers, 3D scaffolds or monoliths, and thin films or coatings [292].

In this section, various types of coating technologies that can be used for the coating of bioactive glasses and their composites on different types of metals, alloys, and certain specific surfaces are discussed. One of the most simple and economical coating processes is the sol-gel dip-coating process. However, the coatings are often porous because of the solvent evaporation leading to poor corrosion resistance and mechanical properties. Nevertheless, this problem can be solved by incorporating metal oxides such as B2O<sup>3</sup> as reported by Pinki Dey et al. [293]. According to their report, by replacing the silica weight percentage in the 45S5 system by 1% to 5 wt. %, they were able to decrease the porosity in the particles. Thermal spray coating, an industrial coating process, can also be employed for bioactive glass coating preparation. This process involves the coating of bioactive glasses as fine droplets or as plasma and sprayed over metal surfaces. Porous and non-porous layers with varying coating thicknesses can be achieved by the thermal spray process by tuning the deposition parameters such as velocity, size of the droplets, and temperature of the substrates [294].

Bioactive glasses can also be coated by physical deposition techniques such as radiofrequency magnetron sputtering (RF-MS) and pulsed laser deposition. In a recent study conducted by Qaisar Nawaz et al. [295], silver nanoclusters embedded in a silica matrix were deposited over the PEEK/BG layer using RF co-sputtering. They report a uniform 100 nm of the Ag-SiO<sup>2</sup> layer that showed slower and sustained release of silver ions compared to the electrophoretically deposited coating. Although the physical deposition techniques are very robust and highly reproducible, their shortcoming is often the expensive experimental setup and precursors when compared to wet chemical sol-gel coating techniques. On the other hand, electrophoretic deposition (EPD) combines both the advantages and disadvantages of sol-gel coating and physical deposition methods. EPD is both a versatile and cost-effective method for coating ceramic materials on conducting surfaces.

Ashokraja et al. [296] reported bioactivity in simulated body fluid (SBF) and reactive oxygen production using the XTT assay for reduced graphene oxide (rGO), sol-gel derived bioactive glass rods (BGNR) followed by different methods for developing composites of rGO and BGNR such as under constant stirring (COL), under constant sonication (SOL), and with a simultaneous reduction in graphene oxide-BGNR composites (RED). In their study, they report the role of pH changes in the sol-gel process facilitating one-dimensional rodshaped bioactive glass formation, and their immersion studies exhibited a 50-micron thick HAp layer on the seventh day for rGO/BG composites [297]. Their work also reports that the different methods employed to prepare the composites influence the HCA formation, antibacterial efficacy, hemocompatibility, and cell proliferation as shown in Figure 7.

Figure 7. Schematics for HCA formation, antibacterial activity, hemocompatibility, and cell proliferation of bioactive glass rods (BGNR) and their composites with rGO (COL, SON, and RED) [297]. Figure also shows the bioactive behavior of the BGNR-rGO composites. It is noticed that the RED composites showed better HCA layer formation, cell proliferation, and hemocompatibility. **Figure 7.** Schematics for HCA formation, antibacterial activity, hemocompatibility, and cell proliferation of bioactive glass rods (BGNR) and their composites with rGO (COL, SON, and RED) [297]. Figure also shows the bioactive behavior of the BGNR-rGO composites. It is noticed that the RED composites showed better HCA layer formation, cell proliferation, and hemocompatibility.

porous layers with varying coating thicknesses can be achieved by the thermal spray process by tuning the deposition parameters such as velocity, size of the droplets, and tem-

Bioactive glasses can also be coated by physical deposition techniques such as radiofrequency magnetron sputtering (RF-MS) and pulsed laser deposition. In a recent study conducted by Qaisar Nawaz et al. [295], silver nanoclusters embedded in a silica matrix were deposited over the PEEK/BG layer using RF co-sputtering. They report a uniform 100 nm of the Ag-SiO2 layer that showed slower and sustained release of silver ions compared to the electrophoretically deposited coating. Although the physical deposition techniques are very robust and highly reproducible, their shortcoming is often the expensive experimental setup and precursors when compared to wet chemical sol-gel coating techniques. On the other hand, electrophoretic deposition (EPD) combines both the advantages and disadvantages of sol-gel coating and physical deposition methods. EPD is both a versatile and cost-effective method for coating ceramic materials on conducting

Ashokraja et al. [296] reported bioactivity in simulated body fluid (SBF) and reactive oxygen production using the XTT assay for reduced graphene oxide (rGO), sol-gel derived bioactive glass rods (BGNR) followed by different methods for developing composites of rGO and BGNR such as under constant stirring (COL), under constant sonication (SOL), and with a simultaneous reduction in graphene oxide-BGNR composites (RED). In their study, they report the role of pH changes in the sol-gel process facilitating one-dimensional rod-shaped bioactive glass formation, and their immersion studies exhibited a 50-micron thick HAp layer on the seventh day for rGO/BG composites [297]. Their work also reports that the different methods employed to prepare the composites influence the HCA formation, antibacterial efficacy, hemocompatibility, and cell proliferation as shown in Figure

perature of the substrates [294].

surfaces.

7.

A recent comparative study reported results between pure BG and rGO/BG thin films deposited over the anodized surface of titanium by EPD. The deposited bioactive coatings (both pure and composites) were 2 µm thick and exhibited very good HAp formation in simulated body fluids along with super hydrophilicity in pure bioactive glass coatings A recent comparative study reported results between pure BG and rGO/BG thin films deposited over the anodized surface of titanium by EPD. The deposited bioactive coatings (both pure and composites) were 2 µm thick and exhibited very good HAp formation in simulated body fluids along with super hydrophilicity in pure bioactive glass coatings [298]. Table 8 summarizes a brief list of bioactive glass coatings, their compositions, coating processes: important features are elucidated.


**Table 8.** Composition, the substrate used, coating process, and their salient features of bioactive glasses.


#### **Table 8.** *Cont.*

#### **4. Summary and Future Directions**

This paper reviews different biomaterials and explains their significant characteristics that influence their bioactivity. Bioimplant manufacturing involves an integrated process of selection of materials, design, fabrication, and surface modification through micro/nano texturing or coating application. Engineering of native metals by converting them into alloys yields desired properties and provides flexibility in designing the needs as per implant requirements. For a long-term application of bioimplants, surface characteristics and their biological functions are considered as key factors. Engineering the surface of the biomaterials by applying suitable coatings provides flexibility in tailoring the properties as per the requirements.

Bioceramic coatings hold great potential by tailoring the biological properties that suit our needs: the choice of the coating depends on the interaction between the cells with the coatings and substrates that are being used. Coatings on metallic implants are invaluable due to their functionality, biocompatibility, durability, and stability. Bioactive coatings are used to enhance the biological fixation between the bone and metallic implant despite their poor tribological and mechanical properties. Hence, they are often improved by developing composites with materials that possesses good mechanical strength. These improved coatings can be also used for durable load-bearing implants. All these properties lead to a better clinical success rate in long-term use in comparison to uncoated metallic implants. The bioactive ceramic coated biodegradable implants provide synergistic properties of both the implants and coating. Thus, these coatings find applications in cardiovascular stents, heart valves, orthopedic applications, tissue engineering, drug delivery, and biosensors. The current trends of ceramic coatings coated metallic implants are more focused on orthopedic applications.

Feasibility studies on complex structures, designing, fabrication of metallic alloys to form complex shapes without losing mechanical properties and surface integrity are a challenging task and should be attempted. The degradation mechanism of coatings on metallic implants changes in the human body environment. Moreover, lattice mismatch and the accumulation of residual stress cause degradation of the implant after implantation. Thus, there is a need to develop mathematical models for the prediction of degradation mechanisms. Another approach to reducing the residual stress is to deposit a functionally graded multi-layered or nanocomposite coating with multifunctional properties.

**Author Contributions:** Conceptualization, K.K.A.M., A.R.C. and A.D.; validation, A.P. and D.G.; writing—original draft preparation, K.K.A.M., A.R.C. and A.D.; writing—review and editing, A.P. and D.G.; supervision, A.P. and D.G.; project administration, D.G. and A.P.; funding acquisition, D.G. and A.P. All authors have read and agreed to the published version of the manuscript.

**Funding:** This project received funding from the European Union's Horizon 2020 research and innovation program under grant agreement No 739566. This work was also created in the frame of the project Centre for Functional and Surface Functionalized Glass (CEGLASS); ITMS code is 313011R453, operational program research and innovation, co-funded from the European Regional Development Fund.

**Institutional Review Board Statement:** Not Applicable.

**Informed Consent Statement:** Not Applicable.

**Data Availability Statement:** Not Applicable.

**Conflicts of Interest:** The authors declare no conflict of interest.

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