Advancing dental implants: Bioactive and therapeutic modifications of zirconia.

Zirconium-based implants have gained popularity in the dental implant field owing to their corrosion resistance and biocompatibility, attributed to the formation of a native zirconia (ZrO2) film. However, enhanced bioactivity and local therapy from such implants are desirable to enable the earlier establishment and improved long-term maintenance of implant integration, especially in compromised patient conditions. As a result, surface modification of zirconium-based implants have been performed using various physical, chemical and biological techniques at the macro-, micro-, and nano-scales. In this extensive review, we discuss and detail the development of Zr implants covering the spectrum from past and present advancements to future perspectives, arriving at the next generation of highly bioactive and therapeutic nano-engineered Zr-based implants. The review provides in-depth knowledge of the bioactive/therapeutic value of surface modification of Zr implants in dental implant applications focusing on clinical translation.

Introduction

Biocompatible and corrosion-resistant zirconium (Zr) has gained popularity as a materialchoice for orthopaedic and dental implants [2,3]. In just two decades since ZrO2 (Zirconia) was introduced as a biomedical grade metal, around 600,000 femoral heads have been implanted worldwide, and the market for dental implants has increased by more than 12% per year [6]. While titanium (Ti) dental implants demonstrate excellent biocompatibility and have been the popular choice clinically, the following limitations associated with Ti have led to a search for an alternative material choice [7]:

Grey colour (reduced aesthetic outcomes) [8].

Development of hypersensitivity to Ti [9].

Accumulation of Ti particles in lymph nodes and organs [10].

Corrosion in the presence of fluoride or metal alloys in saliva [11].

Oxidation induced by bacterial biofilms in acidic conditions [12].

Initially used for fabricating crowns and abutments, Zirconia ceramics (superior biomechanical characteristics as compared to other ceramics like alumina) have become a popular choice for dental implants [13]. It has been established that micro-rough ZrO2 implants are equivalent to the ‘gold standard’ Ti micro-rough implants in terms of osseointegration capacity [14].

The following attributes are the key reasons for preferring ZrO2 over Ti as a dental implant material choice:

White, opaque colour

Reduced affinity to bacterial plaque, reduced inflammatory infiltrate and favourable soft-tissue integration [15], translating into reduced risk for peri-implant diseases

Reduced thermal conductivity, high flexural strength and high fracture toughness

In comparison to other metals, like stainless steel, CoCr alloys and Ti alloys, ZrO2 is non-magnetic [16], which means it does not interfere with standard diagnostic techniques, such as magnetic resonance imaging (MRI) [17].

Clinically, Zr-based implants have shown promising outcomes with low ion release, lower cytotoxicity, favourable biocompatibility, and good osseointegration capability compared to Ti [18].

Zr readily forms biocompatible ZrO2 upon exposure to oxygen and the bio-inertness and non-resorbable nature of this oxide layer makes Zr an ideal candidate for dental implants. For dental implants, the ivory colour of ZrO2, which resembles the natural tooth, makes it an aesthetic choice for dental restorations [19,20]. Various in vitro and in vivo investigations have established the biocompatibility and osteogenic potential of Zr-based implants [14]. However, in compromised patient conditions (inadequate bone quality/quantity, aged and diabetic patients), early establishment and long-term maintenance of osseointegration at the bone-implant interface, and soft-tissue integration at the transmucosal region of dental implants, may be inadequate [21].

Further, compromised conditions increase the possibility of bacterial infection and implant failure, requiring thorough decontamination and revision surgery [22,23]. Studies have shown that ZrO2 surfaces are associated with reduced bacterial accumulation and that the bacterial plaque grown on ZrO2 were less mature than Ti counterparts [24,25]. Additionally, ZrO2 implants demonstrate reduced amounts of inflammatory infiltrate (albeit further clinical evidence is warranted) and promote soft-tissue integration [7]. This evidence suggests that ZrO2 implants/abutments with low bacterial colonization potential and immunomodulatory properties may reduce the risk of peri-implant chronic inflammation associated diseases (such as mucositis and peri-implantitis). Hence, ZrO2 may be particularly favourable for use in compromised conditions that predispose to peri-implant disease, including diabetic and immuno-compromised patients (eg. post-radiation therapy).

To address such challenges, surface modification of Zr-based implants to enable enhanced bioactivity and local therapy has been proposed to counter poor implant integration and bacterial infection. Moreover, attainment of immuno-modulation can further support integration and control of infection [26]. However, with various modification strategies employed, involving physical, chemical, and biological enhancements spanning across the macro-, micro-, and nano-scales, there is a gap in understanding the bioactivity and therapeutic effectiveness of such implant surface modifications. In this review, we compare and contraste the current knowledge of Zr-implant modification for improved understanding towards the clinical translation of the next generation of highly bioactive and therapeutic Zr implants.

Surface modification strategies and bioactivity evaluation

To date, various strategies have been employed to alter the surface texture of Zr-based implants; for example, physical, chemical, electrochemical and bioactive treatments. This review addresses a knowledge gap regarding surface modification of Zr implants to provide an improved understanding of the various strategies, their optimization, and effectiveness to enable easy clinical translation. Fig. 1 summarizes the various topographical, chemical and bioactive modifications performed on Zr implants to impart unique characteristics.

Fig. 1

Surface modification of Zirconia implants. Schematic representation of various surface topographical, bioactive and chemical modifications and the nano-engineered topographies.

The interface between living tissue and the biomaterial surface has been studied extensively for Ti and its alloys [27] but remains poorly understood for Zr-based implants. Both the topography and chemistry of the implant surface influence early-stage cellular interaction and dictate the fate of the implant [28]. Studies have shown that ZrO2 reduces bacterial adhesion and biofilm formation [29], while thickening of the native ZrO2 film (varies between 2 and 5 nm) may improve the barrier effect against corrosion [30]. Both in vitro and in vivo studies have established that ZrO2 implants have superior osseointegrating abilities [31]. Despite such favourable outcomes, it is noteworthy that the long-term clinical results have not been appropriately explored, and controversy regarding the osseointegration ability of ZrO2 implants remain unaddressed [32].

In order to achieve successful long-term treatment outcomes for Zr-based orthopaedic and dental implants, modification to alter surface chemistry, topography, and bioactivity has been suggested, as widely applied for Ti implants [21,33]. Zr surface modification can influence cellular adhesion, proliferation, spreading morphology and differentiation of cells that interact with the implant at the implant-tissue interface [34]. For dental implants, such modifications can augment biocompatibility and integration (both hard- and soft-tissue) towards achieving favourable clinical outcomes and peri-implant stability.

Physical modifications

Physical or mechanical methods have been widely applied to fabricate rough or smooth implant surfaces via either subtraction or attrition processes. This serves the purpose of attaining desired surface topography towards bioactivity enhancements while also facilitating surface cleaning. Techniques like machining, polishing and grit-blasting have been applied towards the modification of Zr-based implants. Additionally, sputtering, plasma spraying, arc melting, physical vapour deposition, laser treatment and magnetron sputtering have also been performed to render Zr surfaces bioactive, as demonstrated in various investigations [[35], [36], [37], [38], [39], [40]].

Initial attempts at modifying Zr implants involved grit blasting (alumina particles 50–110 μm), which enabled augmented peri-implant osteogenesis and osseointegration compared to machined Ti [41]. Studies have, however, shown that grit blasting can reduce fatigue resistance of ZrO2 [42]. It has been suggested that the use of soft and round particles for grit blasting can reduce the formation of micro-cracks while still producing the desired roughness. Compared to machined, plasma-sprayed, and alumina blasted Ti, Zr sandblasted Ti implants significantly enhanced bone ingrowth, as demonstrated in sheep implantation in vivo [41].

Lasers have also been employed to modify the surface texture of ZrO2 implants [43,44]. This method reduces the water contact angle (making the surface more hydrophilic), thereby augmenting the implants' osteogenic potential [44]. For instance, continuous-wave Nd: YAG laser-treated Zr oxidation on Ti was reported to enhance cell-material interactions in vitro [40]. It is noteworthy that laser-oxidized Zr has two orders of magnitude reduced wear rates compared to as-deposited Zr, attributed to the high surface energy and wettability [40]. Laser-oxidized Zr contains both monoclinic (m) and tetragonal (t) oxides, and increasing the t-phase (enhanced surface energy and hydrophilicity) has been shown to promote osteoblast functions [40]. Further, femtosecond laser exposure has been used to generate micro-grooves on ZrO2, which increased the number of transverse collagen fibres and enhanced bone remodelling, compared with grit-blasted ZrO2 and micro rough sand blasted and acid-etched (SLA)-Ti [45].

Plasma treated Zr surfaces have been generated using plasma electrolytic oxidation (PEO), plasma immersion ion implantation (PIII), ion-assisted arc-plasma deposition and simple plasma spraying [38,[46], [47], [48]]. Ivanova et al. reported the use of ion-assisted arc-plasma deposition on Ti–Zr alloys. They reported an increase in nanohardness due to increased Zr content in the coating [48]. Briefly, plasma-modified bare Zr, Zr coated Ti, or Zr incorporated Ti alloys have demonstrated enhanced osteogenic potential, superior mechanical properties and corrosion resistance [47,49]. Recently, Liu et al. reported deposition of Zr-incorporated amorphous carbon gradient multilayer films on Ti alloys via magnetron sputtering toward enhancing bioactivity, as well as mechanical and bio-tribological properties [50]. Compared to bare amorphous carbon and Ti alloy, the Zr–C/Ti alloys significantly augmented wear resistance and osteoblast functions (viability, proliferation and adhesion) in vitro. Yuan et al. fabricated a barrier layer of ZrO2 nanofilm on Zn–Li alloys using atomic layer deposition (ALD) that enabled controlled biodegradation and augmented osseointegration abilities in vivo [51]. Noting that the release of metallic particles from the implant surfaces can trigger immuno-toxicity [52], the authors reported reduced ZrO2 accumulation in organs, which was attributed to nanocoating, and suggested the use of ZrO2 coating via ALD modification to control the corrosion of biodegradable metals and augment their biocompatibility. Various physical surface modifications of ZrO2 and Zr alloys are summarized in Table 1.

Table 1

Summary of surface modification of Zr/ZrO2 or Zr-alloyed implants via physical methods.

No.	Surface	Modification Strategy	Treatment	Dimensions	In-vivo/In-vitro studies	Bioactivity Evaluation/Conclusion	Ref.
1	Cp Zr	Mechanical polishing + femtosecond laser-assisted texturing	Smooth polished surface	D: 10 mmT: 2 mm	In-vitro HDFa (Human Dermal Fibroblasts-Adult) cells) and in-vivo (rat) model	Similar to Ti alloy, modified Zr bioactivity depends on both topography and chemistry	[35]
2	Zr plate	Particle blasting	Zr Macro -machined	D: 6 mmT: 1 mm	S. epidermidis	High roughness and hydrophilicity increase bacterial interaction with surfaces.	[89]
			Zr Micro - particle blasting
3	Zr oxide discs	Cell cultured directly on Zr oxide discs	Expression profiling by DNA microarray	–	Osteoblast-like cells (MG-63) in vitro	ZrO2 is able to modulate immunity, vesicular transport and cell cycle regulation	[3]
4	Zr	Physical treatment	Machined Ti	–	Sheep tibia mid-diaphysis cortical bone in vivo	Zr-SL implants showed significantly increased bone ingrowth and microhardness than Ti.	[90]
			Titanium oxide: plasma sprayed
			Alumina sandblasted
			Zr: sandblasting
5	ZrO2	wTo different glass layers: AP40	Ball milling	–	SBF	Both glass coatings substantially enhanced ZrO2 integration with bone cells	[91]
		RKKP
6	ZrO2	–	Magnetron sputtering deposition	–	E. coli and S. aureus	S. aureus adhesion was lower on ZrO2	[36]
7	Zr5Ti (5 Ti wt%)	HAP-ZrO2-Ag	Multiple electron beam drip melting + plasma laser deposition	D: 8 mmH: 8 mm	Pig tibia model in vivo	Significantly augmented osseointegration for HAP-ZrO2-Ag coated Zr45Ti	[5]
	Zr25Ti (25 Ti wt%)
	Zr45Ti (45 Ti wt%)
8	Zr Macro	–	Machined: only cleaned	T: 1 mmD: 6 mm	S. epidermidis	Machined samples showed reduced biofilm formation	[93]
	Zr Micro		Particle Blasting: airborne particle abrasion with 50–100 μm Al2O3 particles.
9	Ti–13Nb–13Zr	–	Plasma electrolytic oxidation (PEO)	–	SBF and human bone marrow-derived mesenchymal stem cells (hBMSCs) in vitro	Upregulated osteoblast activity	[38]
10	Ti–35Nb–7Zr–5Ta (β micro-structure)	–	Machined	–	–	Lowest hardness and elastic modulus (p < 0.05) and increased polarisation resistance relative to cpTi.	[94]
			Plasma electrolytic oxidation			Porous surface with increased roughness, surface free energy, hardness and electrochemical stability
11	Ti6Al4V	Ti–Nb–Zr–Si thin film metallic glass (TFMG)	Sputtering	–	Cytotoxicity test with L929 fibroblast cells	Superior corrosion resistance and electrochemical stability, non-cytotoxicity, better hemocompatibility	[95]
12	TiZr alloy	–	MAO treatment	–	MG-63 cells and SKF cells	Increase in cell viability and cell growth	[96]
13	Ti–25Nb–3Mo–3Zr–2Sn	Without the α phase	Surface mechanical attrition treatment (SMAT)	–	hFOB1.19	Enhanced the osteoblast response	[97]
		With α phase
14	(Y,Nb)-TZP/alumina	–	Cold isostatic pressed	–	HOS cells	Supports continuous cellular growth	[98]
15	HA- ZrO2 composite	Ti6Al4V alloy	Magnetic sputtering	–	–	The growth of bone tissue reduce its residual stress	[99]
16	Ti–35Nb–10Zr	HA thin film coating	Femtosecond laser texturing	–	MG 63 osteoblast-like cells in vitro	Significantly higher cell attachment and spreading	[100]
17	Ti–35Nb–2Ta–3Zr	–	Micro-arc oxidation	–	–	Excellent corrosion resistance and hydrophilicity	[101]
18	Zr and ZrO2 alloyed layers	316 L stainless steel	Plasma surface alloying + annealing	–	MC3T3-E1 preosteoblast cells in vitro	Significantly enhanced the wear resistance, improved adhesion and spreading	[102]
19	Zircaloy-2 alloy	Potassium hydroxide/sodium silicate electrolytes	Plasma electrolytic oxidation	–	–	Enhanced corrosion resistance	[103]
20	Ti15Zr alloy	–	Plasma electrolytic oxidation (PEO)	D: 15 mmT: 1 mm	Bacterial test:S. sanguinisProtein adsorption: Albumin	Improved albumin adsorption, reduced bacterial adhesion, improved hardness, roughness and corrosion resistance	[104]
21	Pure Zr	Ti	Continuous-wave Nd: YAG laser	T:7 μm	Human osteoblasts in vitro	Enhanced biocompatibility, excellent cell-material interactions	[40]
22	Zr incorporated amorphous carbon	Ti alloys	Magnetron sputtering	–	Immortalized calvarium osteoblast-like cells in vitro	Improved biocompatibility	[50]
23	Zr	Ti	Ion-Assisted Arc-Plasma Deposition in Vacuum	–	–	Reduced elastic modulus, enhanced elastic strain to failure and plastic deformation resistance	[105]
24	ZrO2	WE43 Mg	Liquid Phase Plasma Technique		SBF in vitro	Enhanced cell proliferation and differentiation	[106]
25	Zr	AZ91 Mg alloys	Plasma Immersion Ion Implantation (PIII)	T: 80 nm	MG-63 cells in vitro	Enhanced corrosion resistance, improved antimicrobial properties in vitro, promoted the bone formation	[47]
26	ZrO2	Zn-0.1 (wt.%) Li alloy	Atomic Layer Deposition (ALD)	–	Mouse osteoblast-like cells (MC3T3-E1) in vitro. Ten-week old male Sprague-Dawley (SD) rats in vivo	Improved cell adhesion and viability in vitro. Promoted osseointegration and controlled biodegradation behaviour in vivo.	[107]
27	Zr	-	Plasma electrolytic oxidation	–	Human osteosarcoma cells in vitro	Porosity increased with frequency, superior pitting/corrosion resistance, good apatite forming ability, and cell adhesion	[108]

Chemical modifications

Chemical immobilization or functionalization can further enhance the bioactivity of Zr implants, and as a result, acid-etching of implants has been widely explored and clinically applied. Further, dual topographical and chemical modifications and sol-gel methods have also been applied on Zr implants. For instance, grit-blasted Zr has been chemically treated with KOH, NaOH, and HF to enhance its bone-forming ability further. It has been reported that incorporation of fluoride on ZrO2 resulted in bone-implant contact (BIC) of 81% [53]. Acid etching has been widely explored for both Ti- and Zr-based implants. Studies have shown similar osseointegration for acid-etched Ti and ZrO2 implants, with no statistical difference observed [54]. It is noteworthy that dual micro- and nano-topography of acid-etched ZrO2 implants may have a synergistic effect on biocompatibility and osseointegration [55]. It is also well established that micro and nanoscale modification can mechanically stimulate cells, thereby altering cell motility, adhesion and shape. This, in turn, influences the early establishment of osseointegration on dental implants [56].

Further, any differences in the physical and chemical characteristics of the implant (which are often interdependent and occur during implant surface modification) can significantly influence cellular responses (both host cells and pathogens). Another study reported a statistically significant reduction in three-species biofilm thickness on grit blasted and acid-etched ZrO2 (ZrO2-SLA) compared to Ti-SLA [57]. Further, beyond minor topographical differences, varied biofilm responses between modified Zr and Ti can also be attributed to material composition and hydrophilicity differences between metal (Ti) and ceramic (ZrO2).

Other investigations involving chemical treatment of ZrO2 include the use of various acids (HF, acetic and citric acids) [58], evaluation of the effect of concentration/time of HF treatment on ZrO2 [59], testing osteoblast functions in vitro [60], and their comparison with established Ti counterparts [61]. Hempel et al. studied the response of osteoblast-like SAOS-2 cells on sandblasted, sandblasted/etched ZrO2 and sandblasted/etched Ti and reported the pronounced effect of ZrO2 on cellular adhesion, proliferation and differentiation [62]. Interestingly, both ZrO2 modifications resulted in similar effects, and the difference with Ti was attributed to the difference in materials.

In terms of mechanical characteristics, a combination of heat and acid treatment can reduce ZrO2 flexural strength due to monoclinic phase transformation and low-temperature degradation conditions [63]. Further, mechanical properties like flexural strength and hardness (altered by chemical treatment) can influence clinical performance.

Recently, He et al. studied the cytotoxicity of HF-treated Ti and Zr implants in a mini pig maxilla model in vivo [64]. Ti/Zr release was quantified using inductively coupled plasma optical emission spectrometry (ICP-OES) and inductively coupled plasma mass spectrometry (ICP-MS). At the same time, a histological analysis was performed 12 weeks post-implantation. Interestingly, Ti particle release from Ti implants was two times higher than Zr release from ZrO2 implants, confirming reduced cytotoxicity and DNA damage from ZrO2. Further, the sol-gel method has been used to modify HF-etched Ti implants with ZrO2–SiO2 sol, which did not alter Ti biocompatibility, and offered corrosion protection [65]. Next, machined cp Ti discs were dip-coated with either TiO2 or ZrO2 nanocoating, which upregulated bone-implant contact and removal torque values [66]. Bioactivity assessments of various chemically modified ZrO2 implants are summarized in Table 2.

Table 2

Chemical modification of Zr/ZrO2 and Ti–Zr alloys towards bioactivity enhancement.

No.	Surface	Modification Strategy	Treatment	Dimensions	In-vivo/In-vitro studies	Bioactivity Evaluation/Conclusion	Ref.
1	Cylindrical low-pressure injection moulded ZrO2	Chemical treatment, acid-etched	Ti-SLA controls	Threaded implants with a 6-cornered shaftD: 4.1 mmL: 10 mm	Mini pig maxillae in vivo	Leached ZrO2-NPs showed lower cytotoxicity and DNA damage compared to Ti-NPs in human cells.No difference in osseointegration between ZrO2 implants and Ti‐SLA controls regarding peri‐implant bone density and BIC ratio.	[109,110]
2	3Y-TZP	Physical and chemical treatment	Micro-structured Ti	D: 5 mm	In-vitro Primary human bone cells (HBC) advanced osseointegration model	ZrO2 surface showed increased fibrinogen adsorption, platelet adhesion, activation, and thrombogenicity compared to Ti surfaces. Mineralization of HBC was significantly higher on ZrO2 but significantly lower compared to nanostructured Ti.	[111]
3	Zr	Treated by calcium phosphate slurry	–	–	SBF in vitro	Calcification of an osteoblast-like cell was enhanced on the treated surface	[112]
4	Zr	Varied concentrations of phosphate, silicate and KOH based electrolyte using a pulsed DC power source	Plasma electrolytic oxidation (PEO)	T: 6–11 μm	Human osteosarcoma cells in vitro	Cells firmly adhered and spread on all the oxide films. Silicon doped films showed higher surface roughness and wettability.	[113]
5	Ti–5Zr	–	Machined polished (MP)	D: 10 mmT: 2 mm	MC3T3-E1 osteoblasts in vitro	MP-DAE treatment improved mechanical properties, cell adhesion/proliferation, and corrosion resistance.	[114]
	Ti–10Zr		Machined polished + double acid etching (MP-DAE)
	Ti–15Zr
6	TiZr	–	Polished (P)	D: 4.39 mmT: 2 mm	Primary human gingival fibroblasts in vitro	Increased expression of fibrotic markers	[116]
			Polished hydride (PH)			Decreased expression of fibrotic markers
			Polished, HNO3/HF acid-etched and hydride (PEFH)			–
			Machined (M)			Cell alignment
			Machined hydrides (MH)			Increased initial cell attachment and the expression of genes necessary responsible for a collagen-rich ECM
			Machined, HNO3/HF acid-etched and hydride (MEFH)			–
			Machined and HCL/H2SO4 acid-etched (MES)			–
			Machined, HCL/H2SO4 acid-etched and hydride (MESH)			–
7	ZrO2	ORMOSILs (Organically Modified Silicate)	Sol-gel process	–	Human osteoblast cell line MG-63 (CRL-1427) in vitro	Supports cell adhesion and proliferation.	[117]
8	ZrO2	cpTi	Sol-gel Technique	–	In vivo in rat tibiae	Improved differentiation of rat MSCs into osteoblasts, increased bone-to-implant contact and removal torque values	[118]
9	ZrO2	Ti	Sol-gel	–	Human osteoblasts and artificial saliva in vitro	Reduced Ti susceptibility to corrosion	[119]
	ZrO2–SiO2

Electro-chemical modifications

Electrochemical techniques such as electrochemical anodization (EA) have been extensively applied to enable the fabrication of controlled metal oxide nanostructures, especially for Ti-based implants [21,67,68]. Briefly, EA involves immersion of target substrate (metal) as anode and non-target metal as a cathode in an electrolyte containing water and fluoride, connected via a DC power supply [69]. Under optimum conditions, controlled metal oxide nanostructures (metal oxide nanotubes or nanopores) are formed on the surface of the metal substrate (anode) [70]. Compared to alternate nano-engineering approaches, EA stands out due to its cost-effectiveness, scalability, and control over the physical/chemical characteristics of the fabricated nanostructures [71]. The same technique has been extended to fabricate ZrO2 nanotubes/nanopores on the surface of Zr implants [[72], [73], [74], [75], [76], [77], [78], [79]] (Table 3). Key findings from these studies include:

Table 3

Summary of electrochemical techniques utilized to achieve enhanced bioactivity from Zr/ZrO2 implants.

No.	Surface	Modification Strategy	Treatment	Dimensions	In-vivo/In-vitro studies	Bioactivity Evaluation/Conclusion	Ref.
1	Cp Zr cylinders	Mechanical polishing + anodization	As- received	L: 40–50 mmD: 1 mm	Twelve-week-old male WKAH/Hok rats in vivo	Anodized implants at (60 V) augments osseointegration	[120]
2	Cp Zr cylinders	Mechanical polishing + anodization	As- received	L: 40–50 mmD: 1 mm For in-vitro 1 cm2 plates	In-vivo: rat femur osteotomy model.In-vitro: RAW 264.7 for osteoclast differentiation, pre-myoblast C2C12-GFP and preosteoblastic MC3T3-E1 cells	Anodized Zr allows bone augmented cell adhesion and proliferation, affecting cytoskeleton alignment and permitting bone cell differentiation.Zr60 V enabled accelerated bone formation.	[121]
			Anodized:30 V (Zr30 V)
			Anodized:30 V (Zr60 V)
3	Zr	Two-step anodization	Zr flat	–	MC3T3-E1 mouse osteoblast cells in vitro	Increased cell adhesion, spreading, ALP activity and mineralization for ZrNT.	[83]
			Zr NT (Nanotubes)
4	Zr	Anodization	Zr NT	–	–	Reduced hydrophilicity with reducing diameters	[84]
5	ZrO2	Anodization	Annealed	–	SBF	Reduced corrosion resistance for annealed ZrO2	[80]
6	Zr-2.5Nb	–	EA	–	SBF	Enhanced corrosion resistance	[122]
7	TiO2–ZrO2–ZrTiO4 (20 V)	–	Anodization + annealing	D: 40 ± 12 nm	SaOS2 cells in vitro	40 nm diameter nanotubes had the highest percentage of cell adhesion	[123]
	TiO2–ZrO2–ZrTiO4 (25 V)			D: 59 ± 17 nm
	TiO2–ZrO2–ZrTiO4 (30 V)			D: 64 ± 23 nm
	TiO2–ZrO2–ZrTiO4 (35 V)			D: 82 ± 26 nm
8	ZrTi alloy	–	Electrodeposition + thermal treatment	–	MC3T3-E1 osteoblasts	Good viability decreased ROS level and a better cytoskeleton organization	[124]
9	ZrO2	TiO2	Anodic Plasma-Electrochemical Oxidation	–	Primary human osteoblast cells, bone sialoprotein (BSP) and osteocalcin (OC) in vitro	Upregulated proliferation and bone formation ability	[125]

ZrO2 nanotubes improve the stability of Zr, and the corrosion resistance can be enhanced upon annealing of the nanotubes [80].

Attempts at optimizing anodization to understand the growth of nanotubes on Zr [73].

Fabrication of smooth and high aspect ratio nanotubes [77].

In 2017, Katunar et al. reported an extensive study focussed on in vitro and in vivo bioactivity evaluation of anodized Zr implants [81]. Cp Zr cylinders were anodized at 30 and 60 V for 60 min, followed by mechanical preparation and degreasing. In vitro culture of mouse myoblasts C2C12-GFP, osteoblastic MC3T3-E1 cells and the macrophage RAW 264.7 murine cells revealed increased cell spreading and osteoblastic and osteoclastic morphology. Further, in vivo implantation in a rat femur osteotomy model confirmed new bone formation around 60 V anodized Zr implants. In a similar study, 60 V anodized Zr implanted in a rat model in vivo obtained significantly enhanced cancellous bone volume and trabecular thickness, confirming the earlier onset of osseointegration around anodized implants [82].

By tuning anodization conditions, controlled nanotopographies can be fabricated on Zr implants. We have recently demonstrated the fabrication of nanopores and nanotubes on micro-rough Zr wires, mimicking clinically utilized Zr implants and demonstrating clinical translation of anodized nano-engineered Zr implants [78]. We have also shown that dual micro-nano structures can be fabricated by conserving the underlying micro-rough topography of Zr implants and superimposing nanotopography [79]. Frandsen et al. compared the bioactivity of ZrO2 nanotube modified Zr implants and bare Zr implants using osteoblasts in vitro and reported enhanced initial adhesion and spreading on nanotubes, along with highly organized cytoskeleton, increased ALP activity and mineralization [83]. Further, annealing ZrO2 nanotubes increased corrosion resistance compared to bare Zr and as formed ZrO2 nanotubes [80]. It is worth noting that the hydrophilicity of ZrO2 nanotube surfaces increases with reduced diameters and annealing (which results in surface cracks), while hydrophilicity reduces with ageing (when nanotubes were exposed to air for 105 days) [84]. These were attributed to the balance of capillary force and reduction in hydroxyl content. Compared with anodized Ti with TiO2 nanotubes, similar results are seen with respect to annealing (increased hydrophilicity) and ageing (reduced hydrophilicity) [85]. However, for TiO2 nanotubes, hydrophilicity decreases with reducing diameters [86]. It is notable that nanotube diameter can be increased via increasing the voltage, current or time of anodization [87,88]; however, we have recently reported that for fabrication of ZrO2 nanotubes/nanopores on Zr, these approaches can result in cracks on the anodic film (when anodizing curved Zr substrates) [78] and excessive growth rates (which can result in nanograss like structures) [79].

Bioactive coatings

Bioactive coatings on ZrO2 have been developed to augment osteoblast functions, induce hydroxyapatite formation, contribute to osteogenesis, and achieve antibacterial properties [[92], [115]]. Numerous studies have reported bioactivity enhancements toward augmented osseointegration/soft-tissue integration on Zr implants [4,129,[142], [143], [144], [145], [146]] involving modifications with calcium phosphate (CaP), hydroxyapatite (HA), and various biopolymers and biomacromolecules [142,[146], [147], [148], [149]]. Bioactivity enhancements using hydroxyapatite and calcium are presented in Table 4.

Table 4

Use of hydroxyapatite and calcium to augment bioactivity of Zr/ZrO2 implants.

Study	Surface	Modification Strategy	Treatments	Dimensions	In-vivo/In-vitro studies	Bioactivity Evaluation/Conclusion	Ref.
1	Zr	HAP-based bioceramic	Single-step Plasma Electrolytic Oxidation (PEO)	–	SBF in vitro	Enhanced bioactivity and reduced microbial adhesion	[126]
2	ZrO2	Ca-doped	Wet synthesis and self-assembly	–	Saos-2 human osteoblastic cells in vitro	Increased stability and enhanced osteoblast activity.	[127]
3	Zr	CaO partially stabilized ZrO2 (Ca-PSZ) coating covered with HA	Micro-arc oxidation (MAO) and hydrothermal treatment (HT)	NanorodsD:50 nmL:450 nm	–	Good hydrophilicity, excellent apatite-inducing ability	[128]
4	ZrO2	CaP decomposed from HAP during sintering	Chemical co-precipitation method	–	Rat osteoblast cells in vitro	Augmented tensile and binding strength. Enhanced proliferation and ALP activity.	[129]
5	ZrO2	Laminin-5	Argon plasma	–	Epithelial cells	Enhanced cell adhesion	[130]
6	Zr	HAP-based plasma electrolytic oxide (PEO)	Single-step plasma electrolytic oxidation	–	In vitro: simulated body fluid (SBF), MTT assay and bacterial adhesion	PEO/Zr surface significantly improved bioactivity under SBF.Reduced bacterial adhesion on PEO/Zr	[131]
7	TiZr	–	Powder metallurgy followed by alkali-heat treatment and Ca-deposition	–	Osteoblast-like cells (SaOS2)	TiZr alloys exhibited excellent cytocompatibility and satisfactory bioactivity	[132]
8	ZrTi alloy	HA/TiO2	Sol-gel method	–	SBF	Good bone-like apatite forming	[133]
9	Ti–13Nb–13Zr alloy	Incorporation of Ca ions	Electropolishing + plasma electrolytic oxidation (PEO)	–	SBF, hBMSC	Enhanced bioactivity	[134]
10	ZrO2/HA composite film	Zr	Plasma electrolytic oxidation coupled with electrophoretic deposition process in a single step	–	SBF,Human osteosarcoma cells	Human osteosarcoma cells could attach, adhere and propagate well	[135]
11	Ti–13Nb–13Zr (TNZ)	Anodization	Plasma electrolytic oxidation coupled with a sol-gel process	–	Osteoblast-like MG-63	Enhanced surface roughness and cytocompatibility	[136]
		+ Adsorbed collagen
		+ Adsorbed lactoferrin
12	Ti–3Zr–2Sn–3Mo–25Nb	HA coating	Micro-arc oxidation (MAO)	–	Left proximal femoral medullary canal of beagles in vivo	Significantly promoted bone ingrowth and the mechanical performance of the bone-implant interface	[137]
13	Ti–35Nb–xZr alloy	HA coating	Electron beam-physical vapour deposition	–	–	Ti–35Nb–10Zr alloy showed higher corrosion potential. HA/Ti–35Nb–10Zr alloy showed high polarisation resistance by crystallization.	[138]
14	Ca doping ZrO2	NiTi alloys	Cathodic plasma electrolytic deposition (CPED) technology	–	SBF in vitro	Enhanced corrosion resistance, excellent apatite-inducing ability, enhanced bioactivity.	[139]
15	ZrO2–Y2O3	Mg–Ca	AtmosphericPlasma Jet Technique	–	MTT cell viability	Higher polarisation resistanceImproved cell adhesion and viability	[140]
	ZrO2–CaO	Mg–Ca–Zr
16	Ca with Zr coating	–	Plasma spray technique	–	SBF, Osteoblast-like MG63 cells in vitro	Cytocompatibility and enhanced cell growth and proliferation	[141]

NPs, nanoparticles; BIC, bone-implant contact; HA/HAP, hydroxyapatite; SBF, simulated body fluid.

Calcium phosphate (CaP) coatings

CaP is a critical mineral component of bone, and incorporation of CaP into implants can lead to the rapid establishment of bone-implant integration [149,150]. CaP coatings enhance calcium deposition and protein adhesion on Zr-based implants, improving surface bioactivity [129,151]. Stefanic et al. reported a stabilized beta-tricalcium phosphate (β-TCP) layer on ZrO2 implants via chemical deposition and hydrothermal treatment (900 °C) [151]. Such β-TCP coatings enhanced the in vitro apatite deposition in simulated body fluid (SBF) solution and promoted serum protein adhesion to ZrO2 substrates [151]. In another study, Quan et al. synthesized and coated a ZrO2–CaP composite onto ZrO2 substrates via chemical co-precipitation and significantly enhanced the in vitro expression of alkaline phosphatase (ALP), interleukin-6 (IL-6), and transforming growth factor-β (TGF-β) in murine calvaria osteoblasts [129].

It is noteworthy that CaP coatings have weak bonding strength with Zr/ZrO2 substrates, especially those obtained from the physical deposition method. Consequently, multiple attempts have been made to reinforce its adhesive strength on implants, including co-coating with HA, laser treatment (before CaP), and hydrothermal sintering (after CaP) [142,149,152]. Besides these techniques, the multiple-layer composite coating is an alternative to obtain stable CaP coatings, as reported by Bao et al. [153]. In this method, a thin TiO2 film is deposited and immobilized on Zr substrate by sol-gel technique and annealing, followed by the sequential deposition and superimposition of octacalcium phosphate (OCP) composite onto the TiO2 film in Si-OCP solutions [153]. Such sequential deposition can generate a consistent Si-doped OCP coating layer without cracks/defects, and the reliable mechanical stability of such coating has been confirmed by ageing tests [153]. Further, Stefanic et al. confirmed the formation of a stable CaP coating via two-step biomimetic deposition: an amorphous OCP layer was initially precipitated on ZrO2 substrate and transformed into an apatite phase that served as a template for superimposing the final OCP coatings [154]. Such technique is scalable and reproducible and allows synthesizing a CaP layer that is strongly integrated onto Zr substrates [154].

Hydroxyapatite (HA)

With a similar crystalline structure to dental enamel, dentin, and alveolar bone, hydroxyapatite (HA) has been used to augment osteogenic potential at bone defect sites [144,155]. Compared with other biological coatings on Zr/ZrO2, HA-based coatings have been widely applied in various in vitro and in vivo investigations to achieve bioactivity enhancements [5,147,156,157]. Applying thermal treatment to Zn-doped hydroxyapatite (ZnHA)-coated ZrO2 substrate (1200 °C for 2 h, after coating) may convert part of the ZnHA coating into other crystalline forms, such as β-calcium phosphate (β-TCP), calcium oxide phosphate, and calcium zirconium oxide, to stabilize the coating layer [156]. Such modified ZnHA coatings have favourable mechanical properties and are firmly adhered to the underlying ZrO2 substrate. ZnHA has shown improvements with respect to osteogenic potential by enhancing MC3T3-E1 osteoprogenitor cell proliferation and spreading morphology in vitro [156]. Cho et al. reported significantly enhanced expression of ALP, alizarin red, and bone marker genes in MC3T3-E1 cells in vitro, using aerosol deposited HA coating [147]. Histological findings from in vivo studies have reported osteogenic enhancements of HA-coated Zr/ZrO2 [5,157].

Various strategies have been employed to modify Zr/ZrO2 with HA. After coating different Zr alloys with HA via plasma laser deposition, Trinca et al. implanted HA-Zr implants in the tibial crest of minipigs and assessed bone formation at different distances from the implants [5]. Briefly, superior new bone formation was observed around HA-coated Zr implant surfaces within one month, with numerous infiltrating cuboid-shaped osteoblasts [5]. Moreover, there were fewer infiltrating macrophages around HA-coated implants than the non-coated Zr counterparts (Fig. 2) [5]. In another in vivo study, nano-crystallized HA-coated ZrO2 implants were implanted into the jaws of Beagle dogs, and histological findings confirmed significantly enhanced new bone formation around coated implants (33 ± 14%) after six weeks in comparison with the non-coated controls (21 ± 11%) [157].

Fig. 2

ZrTi alloy implant modified with hydroxyapatite and silver. Histology of bone tissue at implant surfaces post-implantation in pig tibiae after one month. Bone tissue around ZrTi alloys coated with a composite of hydroxyapatite-zirconia-silver layer (HAP-ZrO2-Ag; A-F) and uncoated ZrTi alloys (control group; G, H). (A, C, E, G) Bone area adjacent to the implant (<2500 μm); (B, D, F, H) bone area 2500–6000 μm to the implant surfaces. For all coated ZrTi implants, the newly generated bone (yellow arrows) was evident after one-month healing with numerous cuboid-shaped osteoblasts infiltration (red arrows and green arrows). Less bone formation was observed around non-coated ZrTi surface, with numerous macrophages infiltration (black arrows). Adapted with permission from Ref. [5].

It is noteworthy that HA modified Zr/ZrO2 implants may cause detachment/delamination of HA coatings when facing high torque at the implant-bone interface during implant placement. Thus, many studies have focused on enhancing HA coatings' bonding stability on Zr/ZrO2 substrates [142,143,145]. One strategy is ceramic slurry infiltration treatment before coating, forming a porous layer on ZrO2 substrate, as reported by Miao et al. [155]. Such a porous layer increases the HA-ZrO2 contact area resulting in higher interfacial bonding strength [155]. An alternative approach involves coating combined yttria-stabilized ZrO2 (Y-TZP) powder with HA crystals on ZrO2 substrates, resulting in significantly increased adhesion strength on ZrO2 surface compared to bare HA coatings [145]. Such composite coatings enhanced human osteoblasts (HOBs) proliferation and ALP secretion in vitro [145]. Further, applying a 2-step freeze-drying treatment (−40 °C for 4 h at 89 μBar, followed by room temperature drying for 24 h at 23 μBar) resulted in a stable HA coating layer on microporous ZrO2 substrates with favourable compressive and flexural strength [143].

It is noteworthy that HA in the bulk form possesses inadequate mechanical characteristics, including weak bending stress and low fatigue resistance [158]. Further, the mismatched thermal expansion between HA and Zr and decomposition of HA during sintering can limit the mechanical bonding strength of HA coatings on Zr substrates [155]. This can result in mechanical brittleness and weakness of the HA-coated Zr implants, compromising the safe implantation and functioning of HA modified ZrO2 implants. To address this, Faria et al. fabricated a ZrO2-HA/TCP composite coating using gas compacting and sintering techniques to produce superior mechanical stability with enhanced hydrophilicity [142]. Table 5 summarizes the bioactivity enhancements on yttria-stabilized zirconia (YSZ) implants using Silica, HA, TiO2 and Si3N4 particles.

Table 5

Surface modification of Yttria-stabilized Zirconia (YSZ).

No.	YSZ/Calcia Coating	Modification Strategy	Dimensions	In-vivo/In-vitro studies	Bioactivity Evaluation/Conclusion	Ref.
1	YSZ with AISI 316-L	Pulsed Electron Deposition	–	–	Working gas pressure strongly affected the surface properties of YSZ films.	[159]
2	YSZ with silica coatings	Soft lithography and sol-gel	–	Human osteoblast-like cells in vitro	Biocompatibility, early alignment of osteoblast-like cells	[160]
3	YSZ with HA coating on Zirconia discs	Wet powder spraying (WPS)	–	SBF, human osteoblast cells (HOB) in vitro	Good mechanical strength, excellent interfacial bonding and bioactivity	[145]
4	YSZ with reinforced TiO2	Plasma spray technique	–	SBF in vitro	Excellent mechanical stability, highly effective in generating apatite	[161]
6	HA	Sol-gel, dip coating + calcination	Particle size: ∼30 nm	SBF Ringer's solution, In vitro Osteoblasts from calvaria of neonatal (<2 days old) Sprague–Dawley rats	Enhanced corrosion resistance	[162]
7	YSZ with 3% of yttria coating with Si3N4 particles	Laser cladding	–	SaOS-2 human osteosarcoma cell line in vitro	Improved cellular adhesion and bone tissue formation, with higher degrees of maturity and overall better quality	[163]

HA, hydroxyapatite.

Dopamine and poly-dopamine

Dopamine and poly-dopamine (PD) can aid in cell-material adhesion and interaction, and their use as coatings has been proposed to augment bioactivity of Zr/ZrO2 surface [148]. Dopamine coating can improve cell adhesion by influencing cell filopodia and enhancing protein adsorption on modified Zr/ZrO2 implants [148,164,165]. Compared with HA and CaP coatings, dopamine coating on Zr/ZrO2 can be achieved via physical deposition (immersion) in dopamine hydrochloride solution [164], although to obtain evenly distributed dopamine coatings, it is necessary to maintain constant stirring and temperature stabilization (37–50 °C) of the dopamine hydrochloride solution [164,165]. ZrO2 substrates modified with 3,4-dihydroxy-l-phenylalanine (l-DOPA) coatings exhibit enhanced protein adhesion capacity and increased MG-63 human osteoblastic cell proliferation and cell spreading in vitro compared with uncoated ZrO2 substrates [166]. Further, increased proliferation of human gingival fibroblasts (hGFs) on PD coated ZrO2 in vitro, with enhanced expression of fibronectin, integrin β1, and secretion of collagen 1 has been reported [167]. Clearly, PD coated Zr implants can be used towards soft-tissue integration and osseointegration to ensure the long-term success of a dental implant system [167].

Biomacromolecular coatings

Biomacromolecules such as Arginylglycylaspartic acid (RGD), a minimal recognition sequence within fibronectin that can promote cell interactions, holds great promise in augmenting the bioactivity of conventional implants [168,169]. Since RGD coatings must be performed under mild conditions to prevent protein/peptide denaturation, sequential pre-treatments are required prior to its coating; these include acid etching, plasma treatment, and salinization [169]. As Fernandez-Garcia et al. reported, RGD-functionalized ZrO2 implants increased murine osteoblasts MC3T3-E1 adhesion rates in vitro [169]. An alternative protocol for immobilizing RGD on ZrO2 substrate is covalent bonding, which involves sequential immersion in acid/alkaline solutions to form hydroxyl groups leading to covalent bonding and strengthening of the RGD coating layer [170]. Such RGD-coated ZrO2 surfaces have been reported to enhance proliferation, adhesion, and differentiation of MG-63 osteoblasts in vitro [170].

Protein and cytokine coating of implant surfaces is problematic due to potential denaturation during the immobilization process [168]. Thus, hydrothermal treatment, which can effectively immobilize CaP/HA coatings, is unsuitable for coating sensitive biomacromolecules [146]. Moreover, to obtain a stable biomacromolecule coat, Zr/ZrO2 substrates must be pre-treated [146]. Aiming to enhance the efficiency of fibronectin (FN) coating on ZrO2 implants, Rubinstein et al. utilized ion beam assisted deposition (IBAD) to create a nanostructured ZrO2 surface (nano-peaks with negatively charged patches) [146]. The nanopeaks obtained by IBAD were ultra-hydrophilic and had enhanced FN adhesion capacity, thus promoting cell adhesion on the FN-coated ZrO2 implant surface [146].

Bone morphogenetic protein-2 (BMP-2) and growth and differentiation factor-5 (GDF-5) have also been immobilized on ZrO2 by applying multiple hydrogel loading treatments [4]. Briefly, the ZrO2 surface was initially functionalized by 2-aminoethyl methacrylate (AEMA)-conjugated HA (HA-AEMA) and then immersed in a hyaluronic hydrogel that contained BMP-2 or GDF-5 [4]. The BMP-2 and GDF-5 loaded ZrO2 surface showed significantly enhanced alkaline phosphatase (ALP) activity from MG-63 osteoblasts at day 7 in a dose-dependent manner. Similarly, Alizarin Red staining showed increased calcium deposition in vitro from MG-63 cells on BMP-2, and GDF-5 coated Zr implants. mRNA expression for ALP and osteopontin in MG-63 cells in vitro was enhanced on GDF-5 and BMP-2 functionalized ZrO2 surface, further indicating enhanced osteogenic potential (Fig. 3) [4].

Fig. 3

Protein incorporated zirconia implants. Zr-1/Zr-3: Non-coated Zr surface; Zr-4/Zr-5: Bone morphogenetic protein-2 (BMP-2) coated Zr; Zr-6/Zr-7: Growth differentiation factor-5 (GDF-5) coated Zr surface. (A) BMP-2 and GDF-5 coatings augmented in vitro alkaline phosphatase (ALP) activity levels of MG-63 osteoblasts at day 7 and 14 on Zr surface. (B) Alizarin red staining showing enhanced calcium deposition from MG-63 osteoblasts on BMP-2 and GDF-5 coated Zr surfaces. (C) Increased mRNA expression of ALP and osteopontin (OPN) from MG-63 cells on BMP-2 and GDF-5 coated Zr surface. Adapted with permission from Ref. [4].

Ultraviolet irradiation

Previous studies have established that UV-irradiated Ti surfaces exhibit increased bioactivity and osteoconductive properties [171]. The same strategy had been extended to ZrO2 surfaces. In a pioneering study, Att et al. evaluated the effect of UV light exposure on ZrO2 [1]. They employed a 15 W bactericidal lamp (250–360 nm) as the UV light source for a period of 48 h and observed an increase in hydrophilicity on the ZrO2 surface (Fig. 4a). Enhanced cellular attachment, spread, and proliferation of bone marrow cells were also observed for ZrO2 surfaces exposed to UV over this time. Although ALP activity and mineralization was significantly higher in UV treated samples, a significant difference in gene expression for osteogenic markers between treated and untreated surfaces was not achieved. This led to the conclusion that enhanced ALP activity was due to the higher number of cells attached to ZrO2 surfaces upon UV treatment, indicating that UV treatment improved cell attachment and proliferation.

Fig. 4

Ultraviolet (UV) irradiated ZrO2 surfaces. (A) Photograph of water droplet on untreated and UV treated ZrO2 surface indicating a shift from hydrophobic to hydrophilic upon UV exposure. (B) Water contact angle of ZrO2 surface at various treatment times. (C) Initial spread and cytoskeleton of osteoblasts (3 h post cell seeding on treated and untreated surface). (D) ALP activity. (E) Osteogenic marker gene expression. Adapted with permission from Ref. [1].

In another study, Tuna et al. evaluated the effect of UV treatment on two types of biomedical grade Zr, Zr1 [(ZrO2 85.7 wt%; Al2 O3 8.3 wt%; Y2O3 4.3 wt%; La2 O3 1.7 wt%] and Zr2 [ZrO2 93 wt%; Y2O3 5 wt%; HfO2 1.9 wt%; Al2O 0.1 wt%] [172]. Smooth (m) and rough (r) ZrO2 (roughened by sandblasting) were exposed to a 15 W bactericidal lamp (250–360 nm) for 48 h. Contact angle analysis showed a significant shift of surface properties from hydrophobic to hydrophilic upon UV treatment. UV treatment also reduced the amount of carbon present on the surface. Further, the culture of osteoblasts in vitro revealed accelerated attachment and enhanced spreading on the UV treated surfaces. However, no significant difference was observed in ALP activity due to surface treatment. This observation is in contrast to the earlier reported by Att et al. [1]. The efficacy of UV treated Zr cylindrical implants in an in vivo rat femoral model was investigated [173] and showed extensive bone formation around UV treated implants after two weeks of implantation, compared to untreated implants. Additionally, enhanced osteogenesis, increased peri-implant bone formation, and better implant-bone integration in the absence of fibrous tissue between the implant surface and the bone was observed after four weeks of implantation.

Alloyed Zr implants

Conventionally, Zr is alloyed with Ti for improving the mechanical properties of Ti. However, in recent years Zr based alloys have gained importance in implant development attributed to their ability to form an intrinsic bone-like apatite layer on their surfaces upon implantation [174,175]. In certain situations like implantation in a narrow edentulous alveolar bone ridge or the replacement of a single tooth in a narrow gap, small diameter implants of enhanced mechanical/tensile strength are required, and in these cases, the mechanical strength of pure metals like Ti and Zr are insufficient [176,177]. The use of Zr-based alloys, such as TiZr, addresses this problem as they have higher mechanical strength making them ideal candidates for conditions requiring enhanced mechanical properties [177].

Titanium zirconium (TiZr) alloys

The enhanced mechanical properties of TiZr alloy, compared to Ti and Zr alone, make them a suitable candidate for small diameter dental/medical implants, especially in high loading settings. Chen et al. evaluated the effect of alkali heat treatment (AH) followed by soaking in stimulated body fluid (SBF) to coat calcium phosphate on the TiZr surface (AH-SBF) [178]. Samples were initially immersed in 10 M NaOH, then heated to 600 °C for 1 h, then finally immersed in SBF at 37 °C for 30 days. It is important to note that the alloy readily forms metal oxide passive film on the surface, providing corrosion resistance. However, upon alkali treatment, a porous sodium titanate and zirconate hydrogel is formed on the surface of the alloy with a porosity of ∼500 nm. Upon immersion in SBF, small cubic particles of CaP deposit on the surface of the material and gradually form a uniform layer of CaP on the porous surface. This apatite mimicking CaP can be used to augment integration with the surrounding bone upon implantation. Elucidation of the oxide layer formed on the surface and the role of Ti and Zr in the formation of sodium titanate and zirconate hydrogel leading to CaP deposition can be obtained via in-depth surface chemistry analysis. Further investigations on the mechanical stability of CaP coatings are also needed.

In 2011, Chen et al. further evaluated the in vitro response of human osteoblast-like cells (SaOS2) towards the AH-SBF modified TiZr surfaces [132]. Enhanced cellular alignment with multiple extended filopodial extensions was observed on the modified surfaces. These results provided evidence that modified TiZr surfaces can be used to orchestrate osteogenesis at the bone-implant interface. Subsequently, Grigorescu et al. fabricated nanotubes on the surface of TiZr alloy using a 2-step anodization process using ethylene glycol electrolyte with 15 vol% H2O and 0.2 M NH4F [179]. The first step of anodization was carried out for 2 h, followed by removing the formed oxide layer by sonication in de-ionized water, with the second step of anodization carried out for 1 h in the same electrolyte. The anodization at voltages 15 V, 30 V and 45 V, yielded nanotubes of diameter 30–40 nm, 50–60 nm and 80–100 nm, respectively, after the first anodization, which then reduced to 20–30 nm for 15 V, 35–40 nm for 30 V, and 40–70 nm for 45 V upon second anodization. An increase in surface hydrophilicity was observed with decreasing nanotube diameters.

It is noteworthy that hydrophilic surfaces demonstrate enhanced bacterial anti-adhesive properties [180]. Further, it is established that nanotubular structures, owing to their high surface area and enhanced protein adhesion, augment osteogenesis [181,182]. However, further assessment of antibacterial efficacy and biofilm formation using a polymicrobial system may provide further insight into how these techniques can be applied clinically.

Charles et al. evaluated surface modification of TiZr alloy using neodymium-doped yttrium aluminium garnet (Nd-YAG) laser [183]. Nd-YAG laser at a wavelength of 1064 nm was moved over samples in a linear motion with 8 W power, 300 mJ/pulse energy, and 50-kHz pulse frequency. The roughness of the TiZr surface shifted from 0.03 μm to 0.06 μm upon laser treatment, indicating that lasers roughened the alloy surface. Contact angle analysis confirmed that the laser-treated surface was hydrophilic compared to the untreated surface. Further, surface modification enhanced the adhesion of human osteoblasts in vitro. Of particular interest is that the lower roughness of the unmodified surface led to the reduced cellular attachment that further hindered cell-cell interaction. In the laser modified surfaces, focal adhesion areas with dendritic projections and filapodial extensions were observed. Increased mineralization was also observed on the modified surface, indicating the osseointegration potential of laser-treated TiZr surfaces.

Plasma electrolytic oxidation (PEO) has been utilized to form a thick porous coating on Ti surfaces to enhance osteogenesis [184]. The same strategy was employed by Cordeiro et al. to modify the surface of a TiZr alloy and evaluate its effect on protein adsorption and bacterial adhesion [185]. Samples were oxidized in an electrolyte containing calcium acetate and glycerophosphate disodium at 290 V and 250 Hz. The PEO-treated surface exhibited a porous morphology with higher surface roughness and hydrophilicity than untreated surfaces and subsequently demonstrated a two-fold increase in protein adsorption. Further, a reduction in the number of colony-forming units of Streptococcus sanguinis indicated that PEO modification limits bacterial adhesion on TiZr surfaces. Additionally, PEO treatment has been reported to facilitate the incorporation of Ca and P present in the electrolyte onto the porous surface in an atomic ratio comparable to hydroxyapatite [184]. Coatings that mimic natural bone structure have great potential to augment implant-bone integration.

Zirconium Niobium (Zr–1Nb) alloy

Zr–1Nb alloys have higher corrosion and mechanical resistance than Zr, making them suitable candidates for bone and dental implants. Kim et al. evaluated the effect of polishing with abrasive paper (#100, #600, #2400) followed by NaOH treatment of Zr–1Nb alloy surfaces [186]. In addition, the treated surfaces were immersed in simulated body fluid (SBF) to evaluate the rate of surface apatite deposition. It is important to note that the effect of surface morphology on apatite deposition is critical in understanding the in vivo bio-mineralization of the implant. An increase in apatite deposition was observed with an increase in surface roughness, with substrates polished using #100 abrasive paper showing higher deposition than #2400 paper polished substrates. Other studies have shown that NaOH-treated Ti–6Al–4V surfaces enhance apatite formation [187]. When Kim et al. compared the effect of NaOH treatment on polished Zr–1Nb surfaces, they did not observe any change in apatite deposition. Therefore, it was concluded that the ZrO2 layer on the alloy surface aids better nucleation of apatite crystals than the TiO2 layer on Ti implants. Consequently, even without NaOH treatment, Zr alloy surfaces have significant potential for bio-mineralization.

Commercial Zr/ZrO2 implants

Three types of Zr/ZrO2 substrates are commonly utilized to fabricate commercial implant/abutment: yttria-stabilized ZrO2 (Y-TZP), alumina-toughened ZrO2 (ATZ), and hot isostatic pressed (HIP) ZrO2 (Table 6) [188]. Y-TZP is fabricated via sintering the composite of ZrO2 containing 3 mol% yttria, under 1300–1500 °C to yield a tetragonal crystallized Y-TZP. Y-TZP is favoured for its outstanding resistance against corrosion and low thermal degradation (LTD, or ZrO2 ageing), and thus utilized by commercial implant companies including Straumann®, Camlog®, Nobel® and ICX® [189,190]. Fabrication of ATZ involves combining 20 wt% Al2O3 with 80 wt% TZ-3Y composite (ZrO2 with 3%Y2O3), pressured in 50 MPa and sintered at 1400 °C for 2 h [191]. ATZ is also corrosion-resistant, with slightly enhanced bioactivity than pure Ti, and is used by Swiss Dental Solutions® (SDS), Nobel® and Zircon Vision® [98]. HIP involves sintering that compresses and densify ZrO2 without altering its chemical compositions, starting at 300 °C and 110 MPa, and continuously increasing to 1200 °C and 205 MPa for 2 h [168,189]. Previous studies have confirmed suitable chemical stability and mechanical strength of HIP-treated ZrO2, used by various implant manufacturers, including Bredent® and Z-systems® [168,189].

Table 6

Summary of commercially utilized zirconia implant types.

	Y-TZP	ATZ	HIP
Full form	Yttria-stabilized ZrO2	Alumina-toughened ZrO2	Hot isostatic pressed ZrO2
Fabrication	Sintering ZrO2 composite containing 3 mol% yttria under 1300–1500 °C	20 wt% Al2O3 + 80 wt% TZ-3Y composite, sintered in 50 MPa, 1400 °C for 2 h	Sintering ZrO2 at 1200 °C, 205 MPa for 2 h
Characteristics	Distinguished corrosion resistance and anti-ageing property	Corrosion-resistant	Pure ZrO2 without chemical changes.Chemically and mechanically stable
Used by Companies	Straumann®, Camlog®, and ICX®	Swiss Dental Solutions® (SDS), Nobel® and Zircon Vision®	Bredent® and Z-systems®
Crystal phase	ZrO2 Tetragonal	Al2O3 Rhombohedral + ZrO2 Tetragonal	ZrO2 Tetragonal
References	[189,190,192]	[98,191]	[168,189]

Various surface modifications have been performed on commercial ZrO2 implants to augment tissue integration, as summarized in Table 7 [149,151,153,155,189]. As reported by Kohal et al., the air-borne particle abrasion method effectively enables micro-roughened topographies on the ZrO2 substrate, which are favourable for osseointegration [193]. Hence, sandblasting and acid etching (SLA), commonly applied to fabricate Ti implants, has also been utilized on Zr/ZrO2 implants fabrication (e.g. ZLA® surfaces by Straumann®) [194]. However, it is also reported that the airborne particles abrasion during SLA treatment could alter the crystalline phase of ZrO2 substrates and compromise resistance against low thermal degradation (LTD), undermining long-term stability [189,193]. Besides sandblasting, other techniques to create micro-roughness to augment osseointegration abilities include milling, sintering and ceramic injection moulding (CIM) [189]. Incorporating bioactive coatings (e.g. hydroxyapatite, dopamine) have also been utilized to modify ZrO2 implants [149,151,153,155]. However, limited studies have investigated the long-term surface stability and in vitro/in vivo biosafety, which represents a significant research gap towards the clinical translation of such bioactive ZrO2 implants.

Table 7

Surface characteristics and modification strategies of clinically used zirconia implants.

Company	Implant Name	Material	Modification	Characteristics	Ref
Straumann®	PURE ceramic®	Y-TZP	ZrO2 sandblast & acid etch (ZLA)	Macroscale and microscale roughened surface	[190]
Bredent®	WhiteSKY®	HIP ZrO2	Sandblasting	One-piece implants only, microscale roughened surface	[200]
Camlog®	CERALOG®	Y-TZP	CIM (Ceramic Injection Moulding)	Microscale roughened surface	[201]
Nobel®	NobelPearl®/ZiUnite®	ATZ	Sandblast, acid etch & hydrophilic treatment	Lower plaque accumulation and enhanced soft-tissue integration	[202]
Z-systems®	Z5c/Z5m/Z5s®	HIP ZrO2	Sandblast and laser treatment	Predictable osseointegration	[203]
Swiss Dental Solutions®	SDS 1.0/1.1/2.2®	ATZ	Microporous surface by sandblasting	Microroughened surface	–
Zircon Vision®	ZV3 series®	ATZ	Milling & sintering	High surface roughness; osseointegration	[204]
ICX®	ICX-Active-White®	Y-TZP	N/A	Microroughened surface	–

To date, various studies have established the clinical reliability of commercial ZrO2 implants, including favourable implant survival rate (ISR) and restricted marginal bone loss (MBL) [[195], [196], [197], [198]]. As Pieralli et al. reported in 2014, among 12 clinical studies published between 2010 and 2015, the overall one-year ISR of commercial ZrO2 implants was 96%, with an average MBL of 0.79 mm after one year [195]. Similarly, Roehling et al. identified 11 clinical studies on commercial ZrO2 implants and showed an average one-year ISR of 94.64% with an MBL of 0.78 mm [196]. These findings are comparable to conventional Ti Implants. Besides predictable osseointegration, the long-term aesthetic outcomes around white-coloured ZrO2 implants/abutments were also favourable [198]. As Naveau et al. reported, peri-implant mucosa discolouration and gingival recession were alleviated around ZrO2 implants compared with the greyish-coloured Ti counterparts, supporting the notion of the superior long-term aesthetics of ZrO2 implant restorations [198]. Such favourable outcomes are attributed to the tooth-like appearance and the chemical stability of ZrO2 implant surfaces [193,199].

Challenges to clinical translation of zirconia implants

Wear and corrosion

Dental implants are subjected to continuous force during both the implant surgery and the masticatory process throughout the lifetime of the implant [205,206]. These forces can severely affect the stability of the implant surface and its modification, and can result in delamination and ion/particle leaching [168,206]. In addition to general wear and potential fretting corrosions developed during long-term functioning, these issues highlight the need to investigate and further understand the stability of ZrO2 implants to ensure their longevity and biosafety [168].

Compared with other implant materials, ZrO2 has reliable physical and electrochemical stability geared toward the favourable long-term performance of dental implants under the constant physical and chemical corrosive environment within the oral cavity [207]. To this end, Tsumita et al. have reported that the ZrO2 abutment-implant interface can bear repeated loads without any delamination, with a capacity against fatigue for ZrO2 abutments similar to Ti counterparts [207]. Corne et al. compared the stability of Ti, Ti alloy, Y-TZP ZrO2, and Y-TZP coated Ti alloys after 16 h of fretting corrosion in simulated human gingiva, under constant contact pressure (100 MPa) [208]. The results showed enhanced anti-fretting capacity from Y-TZP substrate compared to other groups, and a coating of Y-TZP layer on the Ti surface could also significantly enhance its anti-fretting capacity [208]. It is worth noting that ZrO2 is also widely coated onto other biomedical materials to reduce their electrochemical corrosion[119,209,210].

ZrO2 is known to exist in three prominent crystalline phases: monoclinic, tetragonal, and cubic [205,211]. Tetragonal is the main phase of commercial ZrO2 implants formed under a hydrothermal sintering process with reliable mechanical strength, but exhibits reduced stability compared to the monoclinic phase at room temperature. Thus the tetragonal structure of the ZrO2 implant is slowly transformed into a monoclinic phase at room temperature, known as low-temperature degradation (LTD) [212]. LTD can be accelerated with water/moisture that creates cracks on the ZrO2 surface and compromises its mechanical strength [213]. As such, LTD is a significant challenge towards the long-term stability of ZrO2 implants. Various studies have tried to inhibit the monoclinic transformation of ZrO2 implants and maintain their tetragonal phase [212,214]. One option to stabilize the tetragonal phase ZrO2 is to apply yttria to form the Y-TZP composite, that significantly enhances the physical stability in a humid atmosphere [214]. Zhang et al. reported that adding 3–5% yttria to ZrO2 substrate effectively inhibits monoclinic phase translation and stabilizes the ZrO2 composite against water-induced corrosion [215]. Besides Y-TZP, an alternate strategy to resist water-induced ageing is incorporating Al into the ZrO2 substrate to fabricate an ATZ composite [216]. Spies et al. reported that ATZ composites could bear multiple dynamic loading and hydrothermal/water treatment cycles without crack formation or delamination [216]. Furthermore, tolerating multiple loadings under hydrothermal/humid circumstances, most of the ATZ component remained in the corrosion-resistant tetragonal phase, indicating its reliability for long-term functioning within the oral cavity [216].

Cytotoxicity concerns

Despite the widespread use of Zr-based alloys as orthopaedic/dental implants, mechanical wear and tear can lead to the leaching of ZrO2 nanoparticles (NPs) into the surrounding tissue. Accumulation of NPs can lead to cytotoxicity and even acute organ failure [217]. Ye et al. investigated the effect of ZrO2 NPs on the cellular properties of mouse osteoblast cells in vitro [217]. Low concentrations of ZrO2 NPs (0–80 μg/mL) were non-toxic, however, at higher concentrations (100–150 μg/mL) ZrO2 NPs reduced cell viability by 50%. Exposure to the higher concentration of NPs over time led to changes in cellular morphology, an increase in the number of apoptotic and necrotic cells, elevated reactive oxygen species (ROS) levels, and reduced mineralization and osteogenic marker expression, indicating a cytotoxicity effect caused by prolonged exposure and bioaccumulation of ZrO2 NPs.

In another study, He et al. evaluated the potential toxicity effects of Zr implants in mini pig maxillae in vivo [109]. Briefly, threaded Ti and ZrO2 implants were inserted in the maxilla's edentulous parts, and no dental superstructure or loading was performed on the implants. The levels of Ti and Zr ions present within the tissues after 12 weeks of implantation were determined by inductively coupled plasma optical emission spectrometry (ICP-OES) and inductively coupled plasma mass spectrometry (ICP-MS). The findings revealed 1.67 ± 0.42 mg/kg-bone weight of Ti and 0.59 ± 0.13 mg/kg-bone weight of Zr in bone slices adjacent to the Ti and ZrO2 implants, respectively. This confirmed that the amount of Zr leaching from implants was low in comparison to their Ti counterparts.

Further, the spatial distribution of isotopes near the implants showed a higher intensity of 47Ti and 90Zr isotopes close to the screw thread tip, indicating that the chances of wear and ion release may be higher in areas of stress. Histological analysis of bone marrow also revealed Ti particles and traces of bone marrow fibrosis in tissues near the Ti implants, which was congruent with previous reports for Ti implants [218]. In contrast, no Zr particles were found in the bone marrow. However, minor bone marrow fibrosis was observed. Overall, the study showed that the leaching of ions and particles are lower in ZrO2 implants. Importantly, these studies were carried out in a mechanically unloaded condition, i.e., not subjected to masticatory forces. Hence, studies incorporating mechanical loading and more accurately resembling physiological conditions would be of value. The authors also compared the cytotoxicity of ZrO2 NPs and ZrO2 microparticles (MPs) in human periodontal ligament cells in vitro. EC50 of ZrO2-NPs and ZrO2-MPs was found to be 13.96 mg/mL and 80.99 mg/mL, respectively, indicating that NPs were more toxic than MPs. However, the in vivo studies showed that the Zr content near the implant was only 0.75 mg/kg bone, which is 18,613 times lower than the EC50 of ZrO2-NPs, implying that the cytotoxic effect of ZrO2 implants is low. Nevertheless, evaluation of Zr implants and their long term cytocompatibility requires further investigation.

Conclusions

While titanium may be the popular material choice for dental implants, zirconia based dental implants and abutments are receiving increased attention, which can be attributed to their unique characteristics, including white colour (improved esthetics), reduced bacterial affinity, high flexural strength and high fracture toughness, while maintaining the same osseointegration capacity as titanium. Zirconium and zirconia-based implants hold great promise as contemporary dental implants. In order to achieve enhanced bioactivity and therapeutic potential, various physical, chemical, electrochemical and biological enhancements have been performed, which demonstrate favourable outcomes in vitro and in vivo. Further, CaP, hydroxyapatite, polydopamine and other biomolecular coatings have enabled enhanced osteogenesis on zirconia implants. Also discussed in this review are the modified alloyed and clinically used Zr implants towards achieving favourable bioactivity performances. An ideal zirconia implant surface modification would preserve the micro-roughness that to date remains a clinically preferred ‘gold standard’, while superimposing nanotopography to further enhance bioactivity and enable ease of further biomolecule or therapeutic modification. Despite the progress made, significant research gaps remain, including mechanical stability and local cytotoxicity concerns. The next generation of zirconia implants will be nano-engineered with controlled bioactivity to accelerate implant integration, even in compromised patient conditions.

Declaration of competing interest

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

CRediT authorship contribution statement

Divya Chopra: Conceptualization, Investigation, Methodology, Formal analysis, Writing – original draft, Visualization. Anjana Jayasree: Conceptualization, Writing – review & editing. Tianqi Guo: Conceptualization, Writing – review & editing. Karan Gulati: Conceptualization, Methodology, Writing – review & editing. Sašo Ivanovski: Validation, Writing – review & editing, Supervision, Project administration.