Abstract
Titanium (Ti) and its alloys are widely used in orthopedic and dental implants owing to their high biocompatibility with tissues, low toxicity, and excellent mechanical properties, such as high strength, fatigue strength, and corrosion resistance. Total hip arthroplasty (THA) is predicted to rise from1.8 million in 2015 to 2.8 million in 2050, and the demand for Ti-based THA is also increasing. The biocompatibility of Ti originates from the several-nanometer-thick oxide layer present on its surface, which inhibits the redox reactions. The oxide forms spontaneously on the surface upon exposure to air and stays in thermodynamic equilibrium; however, it is easily disrupted by the interfacial shear stress owing to the low wear resistance of Ti. Ti exposed to corrosive body fluids elutes metal ions, generating wear debris in the biological fluids and tissues. This causes injury and disease, incites allergies, and promotes the formation of granulomas and even carcinomas. Furthermore, poor osseointegration due to poor adhesion with adjacent bone causes the loosening of the implant-bone interface and slows the healing process. To overcome these drawbacks of implant Ti materials, surface modifications using biocompatible TiO2 are expected for imparting biofunctions such as osseointegration, antivirus activity, and tribocorrosion. Although various methods have been studied for the fabrication of TiO2 on Ti alloys, anodic oxidation has attracted considerable attention owing to its advantages. This review aims to provide a comprehensive, evidence-based overview of current studies on the osseointegration, antimicrobial properties, and cytotoxicity of surface-modified implant Ti alloys, in addition to a brief introduction to different metallic biomaterials.
Introduction
An increasing number of surgeries are being performed to replace dysfunctional hard tissues, such as bones and teeth, with biomedical devices, such as artificial bones, hip joints, teeth, bone screws, and bone plates. 1 The requirements for biomaterials for hard tissue instrumentation can be broadly divided into biological and mechanical conditions. 2 The biological conditions include (a-1) no toxicity or allergic reactions and chemical stability, (a-2) strong compatibility with biological tissues, (a-3) no carcinogenicity or antigenicity, (a-4) no blood coagulation or hemolysis, (a-5) no metabolic disorders, (a-6) no deterioration or decomposition in the body, (a-7) no extraction, and (a-8) no adsorption or precipitation. The first three criteria ((a-1), (a-2), and (a-3)) are influenced by the chemical properties of the materials employed. The other conditions have minimal to no effect on these properties. In contrast, the mechanical conditions include (b-1) excellent static strength (tensile, compressive, bending, shear, etc.), (b-2) adequate elastic modulus and hardness, (b-3) excellent fatigue resistance, (b-4) good wear resistance, and (b-5) lubricity. Equations (b-1), (b-2), and (b-3) pertain to the properties of the bulk material, while (b-4) and (b-5) are related to the material's surface. Metallic materials are commonly used in clinical practice for medical devices, such as dental, spinal, hip, and knee implants owing to their mechanical properties and durability under load-bearing conditions.3–5 Conditions (a-1), (a-2), and (a-3) are influenced by the chemical properties of the elements in the device material. In contrast, the other conditions are either not significantly affected by the material elements or are unrelated. The high strength and toughness of metals offer advantages for orthopedic implant devices, particularly during hip arthroplasty (THA) and primary total knee arthroplasty (TKA).6,7 Concerning orthopedics, metallic materials such as nickel-plated steel and 302 stainless steels (Fe-Ni-Cr) were used to fix fractures in the nineteenth century, followed by 316 stainless steels (Fe-Ni-Cr-Mo) and Co-Cr alloys for artificial hip joints. Currently, Ti and its alloys, known as excellent biocompatibility, are increasingly utilized in medical applications, including stents and artificial joints. 8 316 stainless steel, an alloy steel with Ni, Cr, and Mo, is particularly valued for its corrosion resistance in non-oxidizing environments and is widely employed in implants such as bone screws, artificial hip joints, and cardiovascular stents. 9 Alloying elements can lead to significant issues when released into bodily fluids due to corrosion.10,11 Co based alloys are commonly used as biomaterials in artificial hip and knee joints because they exhibit corrosion resistance due to their passivation effects as well as wear resistance due to their high hardness. 12 Ti and its alloys are widely employed in biometallic devices due to their superior biocompatibility and mechanical properties. 13 Dental metal materials have a long history; Au was used for dentures and fillings even before the Common Era. Dental amalgams, which consist of Hg, Ag, Cu, Sn, and In, are commonly utilized as casting materials for dental restorations because they can be easily replaced. 14 Amalgams have been replaced by precious metals, Ni-Cr alloys, Co-Cr alloys, and Ti alloys owing to the toxicity of Hg. 15 While superelastic Ni-Ti alloys are widely used as archwires for orthodontic and endodontic treatments,16,17 wires made from Ti alloys with low cytotoxicity have also been introduced. 18 Mg alloys have been developed aggressively owing to their biodegradation properties in physiological environments 19 because the degradation products have no significant adverse effects on the surrounding tissues. However, their low corrosion resistance in chloride-rich body fluids causes the release of metallic ions into surrounding tissues. 20 Zn alloys, such as Mg alloys, are used as degradable metallic materials in orthopedic devices, 21 and their corrosion rates are significantly lower than those of Mg-Al-Zn alloys. 22
Figure 1 shows a photograph of an artificial hip prosthesis with a Co-Cr head and Ti-6Al-4V stem removed from a patient who had worn it for over ten years. The Co-Cr head was clearly corroded, and EPMA analysis detected the corrosion product as Cr oxide. The release of metal particulate due to tribocorrosion damage is called "Trunnionosis," a significant clinical problem. 23 Human body fluids (approximately 0.9% saline solution containing Na+ and Cl−) have a neutral pH 24 ; however, secretions lower the pH to 3–4 due to the infiltration and activation of inflammatory cells in the tissue. 25 In such an environment, metals can dissolve and, in extreme cases, lead to periarticular osteolysis and bone resorption, causing implants to dislodge from their anchorages. 26 Metal ions released from the implant can travel through the bloodstream and cause biological reactions such as toxicity and metal hypersensitivity.27,28 In addition, eluted metal ions combine with proteins in the body to produce proteins (allergens) not originally present. 29 Human immune systems recognize them as foreign substances and send white blood cells to kill them to protect the body by causing inflammation, resulting in the symptoms of “metal allergies”. 30 Nakayama et al. reported the anodic polarization curves of various metals, including biometallic materials, in rabbit and Ringer's solutions. 31 In both solutions, Ti and its alloys exhibited better corrosion resistance than stainless steel (SUS) and CoCrMo alloys. Both Ni and SUS showed a rapid current increase, suggesting a limit to the corrosion resistance owing to the passive layer. Interestingly, the corrosive environment in rabbits was more severe than that in Ringer's solution, indicating that corrosion was more advanced in vivo than in vitro environment. In addition to the dissolved metal ions, the adverse effects of metal debris generated by sliding and fretting wear on the body must be considered. 32 In an artificial hip prosthesis, the relative reciprocating motion between components causes wear damage, generating wear debris and corrosive products in addition to metal ion elution.33,34 Wear debris is engulfed by macrophages, 35 and dying macrophages release enzymes and metabolites that cause severe acidification (a decrease in pH) in the surrounding tissue. 36 Figure 2 shows the potentiodynamic polarization curves in a simulated body fluid, Hanks’ solution, of (a) Ti and (b) TiNbSn alloy (Ti-33.6Nb-4Sn, TNS) anodized for 5 and 30 min. Untreated TNS shows higher passivation current density than untreated Ti, however, the anodized TNS exhibits a remarkable decrease in passivation current density compared to Ti. This suggests that surface modification is effective method to impart corrosion resistance to Ti and its alloys. 37
The hip prosthesis removed from the patient: (a) overall longitudinal section of the CoCr head, (b) appearance of the head, (c) high-magnification section of the severely damaged area of the head, (d) line analysis of the dotted line in (c), (e) appearance of the Ti6Al4V stem neck, (f) high-magnification section of the neck, (g) line analysis of the dotted line in (f). The compositional variation caused by the unevenness of the screw is evident. In the CoCr head, profile (d) indicates that the bright areas in (c) contain a significant amount of oxides, while the grey areas are rich in Co and contain minimal oxygen. Conversely, for the Ti6Al4V stem, profile (f) reveals that the bright areas are abundant in Cr and Mo oxides, whereas the dark areas are primarily composed of Ti and V oxides
Potentiodynamic polarization curves of untreated samples and anodic oxides on (a) pure Ti (grade 1) and (b) TiNbSn alloy (Ti-34.52 wt.% Nb-4.11 wt.% Sn, TNS) for 5- and 30- min in a simulated body fluid, Hanks’ solution. Anodization was conducted galvanostatically at a constant current density of 50 mA/cm2 for 5 or 30 min in an electrolyte containing 50 mM sodium tartrate and 0.7 M H2O2. Reproduced with permission from reference. 37
Metals, especially transition metals, are present in trace amounts in the body and are the active centers of enzyme functions such as oxygen adsorption/desorption, oxidation/reduction, and hydrolysis. 38 These metals are also essential for advanced physiological functions because they promote biological reactions by binding to organic molecules. However, it is important to note that there are negative aspects that depend on content. Okazaki et al. investigated the viability of two cell types—fibroblast cell line L929 and osteoblast cell line MC3T3-E1—when various concentrations of metal ions were added to the culture medium. 39 Increasing concentrations of Ni, Co, V, and Al ions significantly decreased cell viability. Ni possibly penetrates the cell nucleus and produces active oxygen, which damages DNA and causes cancer. 40 Al and V are also highly cytotoxic to healthy individuals, 41 although the toxicity of Al remains controversial. 42 Pentavalent NaVO3 or tetravalent VOSO4 induce apoptosis and enter cells through facilitated diffusion or ionic channels, depending on their oxidative state. 43 In contrast, Ti, Zr, Sn, Nb, and Ta barely affect the viability of mouse-fibroblast-derived L929 or mouse-osteogenic MC3T3-E1 cells.
Metallic biomaterials must be corrosion- and wear-resistant to prevent metal ion elution and debris generation from metal implant materials. 44 Furthermore, the lack of osseointegration of metallic implant materials causes disuse bone atrophy due to stress shielding, as observed in artificial hip joints.45,46 Implant materials must also possess antibacterial properties. 47 In orthopedic surgery using implants, implant-related infections are a major complication, and severe infections may require revision or amputation, significantly affecting the patient's quality of life. 48 While antibacterial implants such as Ag, Cu, iodine, and vancomycin have been studied to reduce implant-related infections, concerns remain regarding the side effects caused by harmful substances eluting into the bloodstream. 49 Recently, antibiotic-resistant bacteria have been reported, corresponding to increased demand for effective antibacterial implants against these resistant bacteria. 50 All the above issues are related to the surface of the implant materials, and we believe that surface modification is essential to address them. To suppress metal ion elution into bodily fluids and promote osteoblast formation on metal surfaces, coatings with hydroxyapatite (HAp), a bone component, have been investigated. 51 However, HAp has poor adhesion to metals 52 and is absorbed by alkaline solutions 53 ; moreover, its porous structure reduces its corrosion resistance. 54 Surface modification of metal implants can impart biofunctions, such as osseointegration, antiviral activity, and tribocorrosion properties, and is expected to promote safe and reliable implant treatment. Bioactive ceramic and hybrid coatings on biomedical metal implants can enhance both the biological and mechanical properties of these implants.55,56 The bond between metal and ceramic often exhibits low adhesive strength, which raises concerns about the durability of ceramic coatings. Under load-bearing conditions in the human body, these coatings are at risk of peeling off, resulting in a loss of functionality as observed with HAp coatings. Therefore, careful consideration of the combination of ceramics and metals is essential for developing biomaterials that effectively possess biofunctions.
Surface modification
Implants must excel in biocompatibility, possess a low elastic modulus, and exhibit high wear resistance. The biocompatibility enables faster osseointegration, and the antibacterial properties are clinically important. The objective of the surface modification for the Ti implants is to further improve their biocompatibility, which Ti inherently possesses to some extent. Some surface modifications are also carried out to improve wear resistance and impart bactericidal activity. This chapter describes an overview of surface modification for the Ti implants with an accent on the practicable technologies besides the anodization; it will be separately summarized in the next chapter. The process and the objective for the surface treatment methods for implant materials including the anodization are summarized in Table 1. The performances to be improved by each treatment are also described in the table.
Table 1 Summary of titanium implant surface modification techniques. The respective processing steps, intended objectives (such as enhancing surface roughness or oxide film formation), targeted performance improvements—mainly osseointegration—and unique characteristics of each method are described. These surface treatments aim to optimize the biological performance and early fixation of titanium implants.
The implant's surface, which is shaped from cast or rolled Ti through machining, is covered with uneven deformation layers. Thick oxide films will form if heat treatment is applied to the machined implant. Therefore, as the first step of surface modification, these layers and films should be removed mechanically to obtain an acceptable surface for the following modifications. Abrasive blasting is an effective surface-cleaning technique to remove the deformation layer or the thick oxide films. Sands or the grit particles of silicon-carbide, alumina, zirconia, HAp, or titania (TiO2) are projected to the surface to be cleaned at elevated speed. 57 The abrasive blasting increases the roughness of the treated surface and activates it as prone to forming a strong chemical bonding with the subsequently prepared layers in other methods. 58 However, it is difficult to completely remove the residual particles that have been embedded on the treated surface with usual post-treatment methods 59 like ultrasonication in a solvent, because Ti is soft with low elastic modulus; thus, the modification with the abrasive blasting is always followed by a chemical and an electrochemical treatment to remove the residual particles 57 from the treated surface for biomedical application. The abrasive blasting with heavier particles promises to bring the compressive residual stresses on the treated surfaces and improve the wear resistance. However, it seems no result has been published for Ti to support this assumption. 60 The Ti-6Al-4V sandblasted with alumina, followed by nitric-acid passivation, showed a high biocompatibility with no cytotoxic effects. 61 The osteoconductivity of the Pure Ti and the Ti-6Al-4V implant that was blasted with alumina particles followed by etching in acid (AB/AE) or treating with the microblasting technique were evaluated by in vivo experiments. The osseointegration was higher in the microblasted and the AB/AE implant than in the as-machined and sterilized one. 62 Chemical treatments are roughly categorized into acid, alkali, passivation, and hydrogen peroxide treatments. After the acid treatment, significantly thin (below 10 nm) oxide layers can be formed on Ti surfaces. 58 However, the chief objective of the acid treatment is the etching of the mechanically treated surface, aiming at the removal of the oxide layer as well as the contaminants, to establish a clean and uniform surface over the Ti implants. The solution composed of HNO3 and 1–3% of HF has been recommended as a standard for acid etching. 63 A mixture of HCl and H2SO4 solutions at elevated temperatures is employed. 64 The surface modification with the sandblasting followed by then acid etching (SLA) in an HCl solution has beneficial effects on the biocompatibility and bone formation around the Ti implant. Although there is a slight decrease in the roughness only on the acid-etching treated surface, it develops more uniform small micro pits and sharp-edged peaks on the blasted surface. 65
The crystalline titanium oxide film with needle-like microstructures of incorporating phosphate-ion, Ti2O(PO4)2(H2O)2 could be coated on pure Ti by a hydrothermal treatment in the dilute H3PO4 solutions at 400°C with a following heat treatment 66 ; the Ti implants coated this film shows significantly higher removal torque forces compared with untreated machined ones. In alkali treatment, Ti implants are modified by soaking in highly concentrated NaOH or KOH solutions for several hours to form the modified layers composed of sodium titanate gel and then crystalize by a heat treatment at 400–800°C. 67 The passivation treatment is the method of soaking the Ti implants in strongly oxidizing acid solutions and forms oxide surface films. Thermal oxidation techniques, the method of heating the Ti implants in an O2-containing atmosphere or soaking in boiling water, are also categorized as the passivation treatment. 58 Hydrogen peroxide reacts with a Ti substrate and forms a TiOOH gel layer. 68 Hydrogen peroxide treatment is a method to form an anatase porous film on the Ti implants, over which a calcium phosphate layer is readily developed. 69 It is also employed the method of immersing the Ti substrates in the dilute HCl solution with 5 to 25% of H2O2 to form the gel layer and heat-treating at 400°C to crystalize.69,70 The sol-gel process is a method to form an oxide layer on a material by adhering a sol-gel solution with the dipping method and following thermal treatment (Figure 3). For the surface modification of the Ti implants, a TiO2 layer, 71 an HAp layer 72 and their duplex layers, the TiO2/HAp coating73,74 are fabricated with the sol-gel process. A composite coating that encapsulates the HAp particulates into the TiO2 matrix has also been developed. 75
Schematic representations of the sol-gel dip coating process. (1) A sol-gel solution layer deposits on the substrate in the process of vertical substrate withdrawing from the solution. (2) The deposited layer is gelatinized through its chemical reactions during the drying. (3) The gel layer is transformed into an oxide film by heating at high temperatures
Li and others found a bonelike HA layer was induced on the Gel-derived TiO2 coating formed on the pure Ti substrate by its immersing in a simulated body fluid (SBF). 71 They estimated the bonelike apatite was induced by the nucleation with the calcium ions, and the phosphate groups that were caught at the coating surface were negatively charged and rich in hydroxyl groups, respectively. The HAp coatings synthesized by the sol-gel method are typically bioactive 72 but have poor adhesion strength to the substrate. On the other hand, TiO2 coatings strongly adhere to Ti, but their biocompatibility is limited. 63 The TiO2/HAp coatings are bioactive due to the presence of hydroxyl groups on the surface, promoting calcium and phosphate precipitation, thereby improving the interactions with osteoblastic cells. The TiO2/HAp coatings on the pure Ti 73 and the Ti-6Al-4V substrate 74 had been reported.
Plasma spraying is a thermal spraying technique, squirting molten material over the substrate and forming its coating (Figure 4); the particles of material to be coated are melted inside a spray gun by loading into a plasma jet at high temperatures up to several ten-thousands degree Celsius. 76 Plasma spraying is used to fabricate the HA or metal oxide coatings for the surface modification of Ti implants. Plasma spraying has an advantage, enabling the emergence of objects of complex shapes at high rates; however, the bond strength of the coated layer between the coating and the substrate is relatively poor because of the mismatch of their thermal expansion coefficients. 58 In order to overcome this disadvantage, a bond coating is often applied between the HAp coating and the Ti substrate, as the TiO2/HAp duplex coating for the sol-gel process. Kurzweg and others have reported that the TiO2/ZrO2 (73/27 mol%) and TiO2 bond coats increase the peel strength against the Ti-6Al-4V substrate by 50% and 100%, respectively. 77 The alloying modification with a high-power laser beam has been used for surface modification of Ti and its alloys.58,78 The modification using carbon or nitrogen as an alloying additive promises to improve the wear resistance of Ti by forming the TiN or the TiC layer through the reaction with the additive and the substrate. Walker and others reported the Ti materials that were subjected to laser surface alloying with carbon and nitrogen enjoyed a high hardness of up to 650 Hv and over 1000 Hv, respectively. 79
Schematic cross-sectional representation of a plasma spray gun. The spray stream with the molten ceramic powder melted by the heat of the plasma flame is sprayed over the material and forms the ceramic coating on it. The plasma flame is generated from the plasma gas, typically Ar heated by an electric arc between the axially aligned cathode and anode
Glow discharge plasma treatment is an old method that exposes the material to a low-pressure glow discharge to obtain a specific surface with some expected properties. 58 Akhavan and others used ion-assisted plasma polymerization to activate the Ti substrate and fabricated Ag nanoparticle-functionalized surface based on wet-chemistry methods. 80 They have reported that the fabricated surface showed a good antimicrobial function. Ion implantation is a promising surface modification that enables the improvement of the wear resistance and bone conductivity of the Ti implants without fear of the delamination of the treated layer.72,81 Ca ion and P ion implantation are useful in the improvement of the biocompatibility of the Ti implants.81–83 While the implantation of Ag and Cu adds antibacterial properties to the treated implants. 84 C ion implantation is regarded as effective in increasing the wear resistance of the implants.85,86
Anodization
Anodization, for Ti implants, generally refers to a surface modification technique to form the oxide layer on a metal surface using an electrochemical method under a constant current or voltage. The pulsed current or voltage, of which amplitude to the anodic side is greater than the cathodic one, is also employed. The anodization is usually performed in the electrochemical cell with a couple of electrodes immersed in an electrolyte; the couple is composed of the substrate to be surface modified, and inert materials like platinum and stainless steel. The former and the later electrodes are called the working and the counter electrodes. The working electrode is oxidized in an electrolyte with the constant or the pulsed anodic current supplied from the outer power source through the counter electrode. The potential difference between the working and the counter electrode, which is called the (cell) voltage, and the current intensity through them are usually monitored during the anodization. The accurate electrode potential of working electrode could be measured using a reference electrode set in the electrochemical cell; the potential difference to the working electrode is measured with a potentiometer, or a voltmeter of high input impedance (Figure 5). 87 The main advantage of anodic oxide films on Ti alloys is the enhanced adhesion and wear resistance. 58 The oxide layer on the surface plays a major role in the success of osseointegration. Thicker and rougher oxide coatings encourage osseointegration to occur reliably and quickly, at least over the shorter term. 88 The result of in vivo experiments for the electropolished Ti implants, which have a smooth surface with a thin oxide, shows significantly lower bone growth around the implants than the other groups in the early phase of the experiments. 89 The oxide coating also has a positive effect on passivating the metal so that corrosion is inhibited and the release of Ti ions is minimized.58,88
Experimental setup for anodization; (1) thermocouple, (2) Ti alloy anode, (3) Pt cathode, (4) Luggin capillary, and (5) Ag/AgCl reference electrode. The alloy anode is anodized with a constant current from the DC power supply in the electrolytic bath. The supplied current, the cell voltage, and the electrode potential against the reference electrode are measured by the ammeter, the voltmeter 1 and 2, respectively. Reproduced with permission from reference. 87
The titanium oxide films forming by the anodization would continuously grow as long as the electric field is strong enough to drive the ions through the oxide film. The final thickness of the oxide film is approximately proportional to the applied potential. At the lower applied potentials, the film thickness increases with the potential level. However, at higher potentials over 100–150 V, the oxide film breaks down electrostatically without a change in thickness, accompanying the increased gas evolution and frequent sparkles from the surface.58,90 The anodization being carried at the potential of this vigorous gas evolution with sparkling is called the Micro-arc oxidation (MAO) method. It is also termed plasma electrolytic oxidation (PEO) 91 or anodic spark deposition (ASD), with the essentially same meanings. After the MAO treatment, porous coatings with high hardness, wear resistance, and adhesion could be produced. Such porous coatings improve the biocompatibility of Ti implants.58,92 The surface morphology, thickness, chemical composition, crystal structure, and photocatalysis properties of the Ti oxide film formed by MAO vary with the applied potential in combining the specific materials and the electrolyte. The oxide film by MAO is usually porous due to the evolution of gas and the spark. The crystal structure of the TiO2 film is a mixture of anatase and rutile. However, there is a tendency for the rutile ratio to increase with the applied voltage. 93 The temperature of the electrolyte would also affect the morphology of the formed film. Habazaki and others obtained a high wear resistance coating on Ti-15V-3Al-3Cr-3Sn alloy by the PEO (MAO) method with the bipolarly pulse current at 200 mA/cm2 on the anodic side in a trisodium phosphate electrolyte containing 0.15 M K2Al2O4. 94 The coating at a low solution temperature of 278K showed the highest wear resistance in dry sliding tests among the investigated temperatures. The authors surmised the high wear resistance of the coating came from the presence of α-Al2O3, which must be derived from the composition of the electrolyte, as well as its less porosity. Li and others anodized Ti-35Nb-4Sn alloy by the MAO method with the bipolarly pulse current at 80 mA/cm2 on the anodic side in a sodium metasilicate and/or a trisodium phosphate electrolyte. 95 The MAO coatings formed in the mixed silicate and phosphate electrolyte exhibited improved wear resistance as well as firm adhesion to the substrate. It was estimated that the silicate oxides in the outer layer of the coating, which were deposited from the silicate components of the electrolyte, contributed to the improvement in wear resistance. Kubota and others anodized Ti-35Nb-4Sn alloy with a constant current of 50 mA/cm2 in a sodium tartrate electrolyte containing 0.7M H2O2. 96 The anodic oxidation proceeded with spark discharge under a voltage over 300 V. The anodized material showed superior fretting tribocorrosion resistance in Hank's solution. The apparatus used for the tribocorrosion tests shows in Figure 6. They concluded this resistance property came from the hard and strongly-adhered rutile TiO2 layer formed on the material by the high voltage anodization with spark discharging.96,97
Schematic illustration of the tribocorrosion tests. The SiC ball being pressed against the working electrode (the sample) immersed in Hank's solution, under a normal load moves reciprocally on its surfaces. The electrode potential or the electrolytic current is monitored with the reference and the counter electrode by using the measuring devices. Reproduced with permission from reference. 96
Implant-associated infection is regarded as a major clinical challenge in implant materials. Cardoso and others anodized Ti-30Nb-5Mo by the MAO process at a voltage of 300V in a calcium acetate monohydrate base electrolyte containing 1.5 to 3.5 mM AgNO3. 98 They obtained a coating contained Ag beside the alloying elements (Ti, Nb, and Mo) and other incorporated elements during the anodizing process (Ca, P, Mg, O, and C). They estimated the AgNO3 was decomposed into metallic Ag with disproportionation at a high temperature of the MAO process. The Ag-containing coating showed substantial bactericidal properties to some kinds of microorganisms. It was noncytotoxic and displayed improved osseointegration capabilities. Muhaffel and others anodized Ti-6Al-4V alloy by the MAO process with the bipolarly voltage pulse of +490 V/-60 V in a calcium acetate hydrate and disodium hydrogen phosphate anhydrous electrolyte containing 0.1 and 0.4 g/L AgNO3. 99 The multi-layer coatings consisting mainly of inner TiO2 layers and outer HAp layers were formed and favored precipitation of Ag nanoparticles on the HAp layer. The coatings exhibited superior antibacterial activity against E. coli strains. Zhang and others anodized pure-Ti by the MAO treatment with the voltage pulse of 390 V in a calcium acetate and zinc acetate base electrolyte, dispersing 6 g/L Ag nanoparticles. 100 The TiO2 coating co-doped with Zn2+ and Ag nanoparticles was fabricated. The coatings showed high antimicrobial efficacy. Hu and others anodized pure-Ti by the PEO treatment with the pulsed anodic current at 165 mA/cm2 in a calcium acetate monohydrate base electrolyte with zinc acetate. 101 The Zn-incorporated TiO2 coatings were formed, and they were highly effective in inhibiting both S. aureus and E. coli strain adhesion. They also exhibited good biocompatibility in vitro tests.
There are some studies that aim to improve the HAp formation on the Ti implants with the anodization in the electrolyte containing Ca and P as the compositions. Santos-Coquillat and others anodized pure-Ti by the PEO method with the bipolarly voltage plus of +490 V/-60 V in a positive-to-negative ratio in a calcium acetate electrolyte with sodium dihydrogen phosphate. 102 The electrophoresis can be used as an auxiliary method for MAO. Nie and others carried out MAO treatment on a Ti-6Al-4V alloy using a phosphate salt solution as an electrolyte and followed by conducting electrophoresis using a HAp powder aqueous suspension as an electrolyte. 103 Gorejová and other electrochemical deposited the HAp inside the pores of the TiO2 coatings that were previously formed by a separate modification. 104 The coating was carried out in the solution consisting of NH4H2PO4 and Ca(NO3)2 with a constant current of 5 mA or a constant potential of 0.8 V.
Stress shielding
In orthopedic surgery, the number of joint replacement surgeries has increased in recent years and is highly important. Globally, over 1 million total hip arthroplasties (THA) are performed annually. 105 This trend is expected to continue, especially among individuals under the age of 65.106,107 This rise in younger and more active patients seeking THA is due to the fact that they have extensive hip degeneration that cannot be adequately addressed through osteotomy, and they desire a return to full activity. While THA has demonstrated positive clinical outcomes, one of its most significant complications is aseptic loosening, particularly of the femoral stem—a concern that is especially critical in younger patients.105,108 Aseptic loosening of the femoral stem is a critical concern in THA, particularly among younger patients.105,108 In fact, a younger age at the time of surgery is associated with a higher risk of revision. For example, the lifetime risk of revision is significantly greater for men aged 50–54 years compared to those aged 70–74 years (29.6% vs. 7.7%). 109 Therefore, efforts to prevent aseptic loosening are crucial for improving the long-term success of THA procedures.
Stress shielding is a factor that contributes to aseptic loosening, especially when it occurs in combination with other risk factors such as polyethylene wear and imprecise installation of the prosthesis. 110 In efforts to reduce polyethylene wear, highly cross-linked polyethylene materials have been developed, leading to improvements in this aspect. 111 However, even with advancements in polyethylene wear reduction, stress shielding remains a concern, particularly in cementless hip prostheses. Radiographic images of a patient after hip replacement surgery resulting in stress shielding are presented (Figure 7). Addressing the issue of stress shielding in cementless hip prostheses is an ongoing challenge that requires further improvements and innovations to enhance the long-term performance and durability of these prosthetic devices. Stress shielding results from the mismatch in Young's moduli between bone and implant materials, leading to abnormal load transfer. This reduces the mechanical stress on the bone, causing bone weakening, resorption, and in some cases cortical hypertrophy (CH).110,112 Excessive stiffness of the stem has been identified as one of the causes of stress shielding. Therefore, there is a need for improved femoral stem materials that combine a low elastic modulus with strength. 113
Representative bone atrophy due to stress shielding following total hip arthroplasty. The arrowheads indicate areas of bone atrophy caused by stress shielding
Several efforts are being made to control stress shielding. The design and material choices for the femoral component of hip prostheses have evolved significantly to address this phenomenon, which can lead to bone resorption due to altered load distribution. Various redesign efforts, surface treatments, and innovative biomaterials have been introduced to promote more physiological bone loading. The first topic concerns the implementation of short stems. Short stem hip prostheses, typically less than 120 mm in length, aim to enable conservative surgery while improving load transfer to the metaphyseal area. Bieger et al. observed a more uniform load distribution in the proximal medial portion of the femur with short-stem prostheses compared to conventional designs.114,115 Their findings indicate a significant difference in stress variation, suggesting that short stems may help in reducing stress shielding. However, the reduced stem length may compromise implant stability, necessitating careful consideration of soft tissue integrity for successful implantation. Short stems have already been applied clinically, but future studies are needed to determine whether short stems prevent stress shielding. There are other efforts to resolve stress shielding in stem design. Hollow stem designs have been explored to modulate the stiffness of the implant, thus influencing the load transferred to the bone. Research by Gross and Abel demonstrated that increasing the internal diameter of hollow stems resulted in increased load application to the bone, thus potentially reducing proximal stress shielding. 116 This design approach allows customization of stem stiffness, which can be beneficial in minimizing unwanted stress concentrations. The Ribbed anatomic cementless femoral stem has shown promising results in retaining bone mineral density at certain areas of the femur. Wu et al. reported that the ribbed stem resulted in restrained osteolysis and more evenly distributed stress in the proximal femur, suggesting a reduced stress shielding effect compared to traditional smooth stems. 117
Surface treatments like sandblasting and HAp coatings have been employed to improve the integration of the prosthetic stem with the bone.118,119 HAp coatings, in particular, encourage bone growth on the stem's surface, accelerating implant fixation and potentially reducing the release of metallic wear particles. 120 To enhance bone ingrowth, prosthetic stems have been designed with porous surfaces that allow bone cells to populate and secure the implant. Porous Ti stems, featuring optimized porosity and pore sizes, have demonstrated reduced stiffness and improved mimicry of bone stress distribution, thereby significantly reducing stress shielding and promoting bone preservation.121,122 These stems are particularly beneficial in preserving bone mass and minimizing osteolysis. However, there are concerns about the strength of porous metallic materials, so ensuring strength for clinical application is an important issue. Ideally, Young's modulus of an implant material should be similar to that of bone while maintaining sufficient strength. To address the elasticity mismatch between the stem and the bone, the Robert Mathys isoelastic stem was created as a low-elasticity option to prevent stress shielding.123,124 However, previous attempts to develop femoral stems made of materials with lower Young's moduli aimed to approximate the modulus of bone but faced challenges due to poor strength, making their results unacceptable for clinical use.
A material should possess biocompatibility and the strength required to fix the femur securely in order to be suitable for use in femoral stems. Ti alloys, particularly Ti-6Al-4V alloys, are commonly used for orthopedic implants because they offer adequate biocompatibility and corrosion resistance. However, Young's modulus of the Ti-6Al-4V alloy is approximately 110 GPa, which differs significantly from the stiffness of bone. This disparity in stiffness is a consideration when designing and selecting implant materials for hip prostheses. Hanada et al. have made significant advancements in the development of a novel TNS alloy. 125 Beta Ti alloys, such as TNS alloy, have been developed to offer lower Young's modulus and higher dynamic stiffness, improving the stress distribution between the implant and the bone tissue. Young's modulus of TNS alloys is approximately 40 GPa. This alloy exhibits a low Young's modulus while maintaining the same tensile strength as Ti-6Al-4V. Additionally, it possesses functionally graded characteristics, allowing for an adjustable Young's modulus through heat treatment. The safety and biocompatibility of TNS alloys have also been verified, and it has been confirmed that they have the same biocompatibility as the conventional metal Ti-6Al-4V alloy. 126 The finite element simulation indicated that the stress transmitted to the proximal medial region of the bone immediately after surgery was 18% higher with the TNS alloy stem compared to the Ti-6Al-4V alloy stem, suggesting a higher load at the proximal end with the TNS alloy stem. 127 The stresses in this region were consistently higher for the TNS alloy stem than for the traditional Ti alloy stem. Furthermore, preclinical studies in rabbits have shown that intramedullary nails or fracture treatment plates with TNS alloys promote bone healing in tibial fracture models compared to Ti-6Al-4V alloy or pure Ti devices.128–132 The results of these studies suggest that the low Young's modulus of the TNS alloy eliminates the stress imbalance between bone and fracture treatment materials and promotes bone healing through adequate load distribution.
Constructing upon this innovative material, a novel cementless femoral stem has been created using the TNS alloy. Figure 8 shows the change in modulus of elasticity of the hip stem resulting from heat treatment. This femoral stem incorporates Young's modulus gradient properties that can be adjusted through heat treatment. Notably, short-term clinical outcomes of TNS stems have shown promising results, including a reduced incidence of stress shielding.133,134 This suggests that the use of TNS alloy in femoral stems may contribute to improved clinical outcomes and potentially address the issue of stress shielding in hip prostheses. The authors noted that although the stress-shielding phenomenon did not completely prevent proximal femoral bone loss, the gradual variation in Young's modulus of the TNS stem was beneficial for bone preservation and provided adequate mechanical strength. Overall, continuous innovation in prosthetic design, surface treatment, and material selection is critical in overcoming the challenges of stress shielding in hip prostheses.
A photograph of a clinically applied TNS hip prosthesis (A) and a schematic model of the temperature gradient and vickers hardness changes resulting from heat treatment of the TNS hip prosthesis (B). The numbers indicate the temperature gradient when the tip is heated to 673K, and the color map shows the changes in Vickers hardness. Reproduced with permission from reference. 133
Osseointegration
Osseointegration is the process by which bone and metal integrate directly, leading to both structural and functional unity between the bone and an implant's surface. 135 The term was first introduced by Per-Ingvar Brånemark in the early 1960s, based on his findings from macroscopic and microscopic evaluations of rabbit bone after Ti screw insertion. 136 He observed dense cortical bone in direct contact with the implant, without any soft tissue layer, and bone tissue extending into microgaps within the Ti surface. These early insights laid the foundation for subsequent research and clinical applications. In Ti and Ti alloys used for orthopedic implants, anodic oxidation is used to obtain osseointegration.137,138
Anodic oxidation is an electrochemical technique that creates oxide or insoluble layers on metal surfaces to enhance their corrosion and wear resistance. The metal is dissolved as a cation in an electrolyte solution in the anodic oxidation process. It then reacts with oxygen ions that originate from the electrolysis of water, forming a metal oxide layer on the surface of the metal. One common industrial application of anodic oxidation is the process known as anodized alumite, which is applied to aluminum surfaces.139,140 During anodized treatment, columnar pores are arranged in a manner parallel to the direction of oxide growth. These pores create a distinctive honeycomb-like pattern when viewed from the direction of oxide growth. The development of the oxide layer is directly proportional to the electric current applied. Simultaneously, some of the oxide may dissolve at the base of the cell, while the remaining undissolved portion grows as the cell wall. This process results in the formation of a porous structure. 141 In vivo, studies focusing on the osteoconductivity of Ti dioxide (TiO2) nanotubes have revealed that Ti alloys coated with TiO2 nanotubes exhibit higher levels of osteoconductivity compared to Ti alloy materials that have undergone blasting or acid treatment. 142 This enhanced osteoconductivity is attributed to several factors. The first is the formation of anatase TiO2. The presence of anatase TiO2 formed coherently with HAp contributes to improved osteoconductivity. 143 This crystalline structure is advantageous for osteointegration. The effect of TiO2 nanotubes on promoting osteoblast differentiation has been reported. 144 The presence of fluorine in TiO2, which is introduced by hydrofluoric acid in the electrolytic bath during nanotube formation, promotes osteoblast differentiation. As a result, interfacial bone formation is promoted. The pore size of TiO2 nanotubes has been reported to influence bone conduction performance. 145 Specific interactions between these nanotubes and integrins, which are membrane proteins present on the cell surface and molecules that bind to the cell cytoskeleton, contribute to improved bone cell response and bone formation. The structures that grow perpendicular to the substrate, such as TiO2 nanotubes, have been shown to promote cellular uptake, an important process for cell growth. 146 Furthermore, studies have demonstrated that coating TiO2 nanotubes with anti-sclerostin antibodies can further enhance osteoblast differentiation and activation. 147 This coating inhibits the secretion of sclerostin, a molecule that suppresses osteoblast activity. Consequently, increased alkaline phosphatase activity promotes osteoblast function and bone formation. While the exact biocompatibility mechanism of TiO2 nanotubes is not fully established, they are regarded as attractive surface structures that enhance the osteoconductivity of Ti alloy implants. These findings hold promise for improving the integration of orthopedic implants with the surrounding bone tissue, ultimately leading to better clinical outcomes. 142 In a previous study focusing on the biocompatibility of anodized TiO2 without a pore structure, a different approach was used to enhance surface roughness. Instead of creating nanotube pores, a high electric field was applied in a weak acid to increase the surface roughness of the Ti material. In this study conducted by Sul et al., anodic oxide layers were produced on pure Ti surfaces. 148 These layers had varying levels of surface roughness, which were systematically controlled by changing the applied potential during the anodization process. In subsequent in vivo tests, it was observed that this anodized TiO2 layer promoted the formation of new bone. 149 Furthermore, the study found that surfaces with greater surface roughness, such as those produced at higher electric potentials, were associated with increased bone formation. This suggests that surface roughness is crucial in promoting bone growth and integration with Ti materials, even without nanotube pores.
The process of anodic oxidation of Ti and its alloy substrate has been investigated in a study. A similar experiment was conducted on the TNS alloy substrate. The TNS alloy was anodized in a 1 M acetic acid electrolytic bath at 200 V for 30 min. Afterward, it was immersed in 25 mL of Hanks’ solution maintained at 36.5 °C for seven days. The specimens were washed with distilled water and dried in a dry incubator for 24 h. The results showed that HAp was observed on the surface of the anodized alloy that had undergone hot water treatment but not on the surface of the anodized alloy without hot water treatment.150,151 Surface analysis confirmed that the observed microstructure was HAp, indicating that bioactivity could be achieved through the hot water treatment of the anodized layer on a TNS substrate. This suggests that the hot water treatment enhances the bioactive properties of the anodized layer on both Ti and TNS alloy materials. To gain further insights, implanted specimens with and without hot water treatment were prepared from the distal parts of the rabbit femurs using a focused ion beam. These specimens were then subjected to microstructure observation and energy-dispersive X-ray spectroscopy (EDX) analysis using transmission electron microscopy (TEM) near the implant-bone interface. 152 The EDX analysis revealed the presence of calcium (Ca) and phosphorus (P) in the anodic oxide layer. These elements are constituents of bone tissue. The high adhesion strength observed in the anodized specimens can be attributed to the penetration of Ca and P, which are elements of bone, into the TiO2 layer of the implant. This suggests that the presence of Ca and P in the anodic oxide layer enhances the adhesion between the implant and bone tissue. Previous research has explored the biocompatibility mechanisms of anodized TiO2, with differences noted between nanotube structures and surface roughening.152–154 While a consistent mechanism has yet to be established for nanotubes, it has been proposed to involve both the porous structure and the composition of TiO2.152,153 In contrast, surface roughening is generally associated with enhanced osteoblast adsorption due to the roughened TiO2 surface, although variations exist among different studies.154,155
The mechanism of HAp formation through hot water treatment is somewhat analogous to surface roughening. Kokubo et al. proposed that hot water treatment increases hydroxyl adsorption and promotes crystallographic coherency between rutile TiO2 and HAp. 156 X-ray photoelectron spectroscopy (XPS) angle-resolved measurements of anodized TNS alloys prepared in a 1 M acetic acid electrolyte have been conducted.150,151 After hot water treatment, XPS analysis revealed a decrease in the fraction of hydroxyl groups (O 1s) and an increase in the oxide fraction compared to the pre-treatment state. For niobium (Nb) oxide (Nb2O5) and tin oxide (SnO2), hot water treatment increased the amount of Nb and Sn but did not significantly affect the number of hydroxyl groups. However, thin-layer X-ray diffraction profiles of TNS alloy showed an increase in the anatase 101 diffraction intensity after hot water treatment, and after in vitro testing, HAp diffraction was observed. Rutile phase diffraction was not detected, indicating that hot water treatment contributed to bone conductivity improvement through mechanisms other than those mentioned earlier. The experiments involving anodic oxidation utilized a change in the electrolytic bath from acetic acid (a weak acid) to sulfuric acid (a strong acid) to induce the formation of a porous structure through the generation of oxygen molecules.157,158 The anodic oxide produced in a 1 M acetic acid aqueous solution displayed anatase TiO2, whereas the one formed in a 1 M sulfuric acid aqueous solution showed rutile TiO2. TEM cross-sectional observations indicated that the acetic acid–anodized layer was about 370 nm thick, while the sulfuric acid–anodized layer was significantly thicker (around 7.7 μm) and exhibited a greater density of internal pores. In vitro tests were conducted on anodized TNS alloy, prepared using a sulfuric acid electrolyte with or without hot water treatment.150,157 In vivo tests demonstrated that the pull-out strength of the hot-water-treated anodized TNS alloy (prepared using acetic acid) surpassed that of its untreated counterpart. 150 Additionally, bone-bonding evaluations showed that sulfuric acid–anodized TNS alloys exhibited even higher pull-out strength compared to untreated rods. Notably, the high bone affinity of sulfuric acid–anodized specimens was maintained irrespective of subsequent thermal or hydrothermal treatments. 157 Non-decalcified tissue images obtained from untreated and anodized rods after six weeks of implantation in a rabbit femur showed robust bone formation around the specimens in the anodized group (Figure 9).150,157 Mapping by EDX at the bone-TNS alloy interface indicated that P and Ca permeated into TiO2, particularly in the pores (Figure 10).151,158 This suggests that the pores in TiO2 are open and connect the surface to the interior, allowing P and Ca to dissolve solidly within TiO2.
Histological images showing new bone formation around TNS rods. Representative histological images of untreated and anodized TNS alloy rods implanted in rabbit femurs are displayed. Panels (a, c) show low-magnification images, while panels (b, d) present high-magnification images of the boxed areas. Mature lamellar bone (indicated by arrowheads) was observed in the anodized TNS group. (a, b) show images with an anodized rod; (c, d) show images with an untreated rod. Reproduced with permission from reference. 157
A transmission electron microscopy (TEM) image of the titanium oxide layer at the interface between the rod and bone in Anodized TNS alloy. The TEM image shows the titanium oxide layer at the rod-bone interface, along with element mapping for Ti, Nb, O, Ca, and P in the same region. Submicron-sized pores, indicated by arrows, were observed. Mapping revealed segregation of Ca and P, with these elements incorporated into the pores. Reproduced with permission from reference. 157
In summary, the porous TiO2 coating on TNS alloys has demonstrated promising osteoconductivity. This effect results from surface structure and chemical interactions with substances in body fluids. The combination of a highly porous TiO2 structure and its chemical interactions with ions from body fluids contributes to the improved osteoconductivity of TNS alloys. 159 On the other hand, the usefulness of anodic oxidation has also been reported for other metals. Magnesium-based biodegradable materials are a new generation of orthopedic implant materials intended to have the same mechanical properties as bone. Anodization is being considered as a solution to magnesium's challenge of rapid corrosion in vivo. 160 It has been reported that anodic oxidation protects the metal surface, improves surface porosity, and improves biocompatibility. The provision of a stable anodic oxide film with controllable in vivo corrosion of magnesium is a future challenge.161,162
Tantalum (Ta) has excellent mechanical and chemical properties that make it a promising biomaterial. 163 On the other hand, Ta exhibits bioinert properties and requires surface treatment to improve biocompatibility. 164 Nanoscale surface modification of Ta surfaces can improve their bioactivity and promote successful osseointegration. 165 Anodization stands out as a promising nanoscale surface modification technique for its simplicity, versatility, and affordability. 166 There have been reports on the bone affinity enhancement effect of Ta after anodic oxidation, but there are no reports on its clinical application, which is expected to be developed in the future.163,167 The knowledge of anodic oxidation of Ti and Ti alloys has been accumulated compared to that of other metals, and efforts will continue to focus on the probability and clinical application of anodic oxidation methods for Ti and Ti alloys due to their biocompatibility, corrosion resistance, and versatility.
Application of titanium oxide to biomaterials
Clark et al. proposed that the lack of surface chemical differences between metal implants and bone tissue may lead to delayed wound healing and failure of tissue formation in the biological environment. 168 This observation indicates the need for a biocompatible substance between the metallic implant material and natural bone tissue. The requirements for the intermediate substance layer are as follows: (1) the osteoblast nuclei must be easily adsorbed onto the layer, (2) the layer must adhere to the substrate and have fracture toughness, (3) the layer must be made of a safe material, (4) the layer must prevent the elution of metal ions from the substrate, and (5) the layer must be abrasion-resistant. Bones are composed of collagen (20%), water (10%), and hydroxyapatite (Ca10(PO4)6(OH)2 (70%) containing a large amount of calcium phosphate. Both Ca and P oxides combined with Ti to form CaTiO3, Ca4Ti3O10, Ca3Ti2O7, and Ti5P6O25 and TiP2O5, respectively. CaTiO3 and Ti5P6O25 are in equilibrium with TiO2 as shown in the phase equilibrium diagram (Figure 11), suggesting that TiO2 becomes an intermediate layer between the metal implants and bone tissue. TiO2 is considered a photocatalyst and is practically applied in numerous fields, such as the quantitative decomposition or sterilization of chemical substances such as endocrine disruptors, injurious volatile compounds, and organic pollutants in water and air.169–171 The requirements for TiO2 as a biomaterial are biocompatibility and antibacterial characteristics, both of which can be achieved by coating the biomaterial with TiO2. TiO2 is an n-type semiconductor with a bandgap energy of 3.2 eV for anatase and 3.0 eV for rutile, which generates electrons and holes upon illumination with ultraviolet (UV) light, corresponding to the bandgap energy. 172 Figure 12 shows the photocatalytic reaction of TiO2 and its associated formulas. The photogenerated charge carriers reduce O or oxidize the water molecules in the atmosphere to yield reactive oxygen species (ROS), such as hydroxyl radicals (•OH), superoxide anions (•O2−), and hydrogen peroxide (H2O2) degrading organic substances173,174 unless charge carrier recombination occurs at lattice defects. 175 The oxidizing power of hydroxyl radicals is second to that of fluorine and more than twice that of chlorine as shown in Table 2. ROS radicals decompose the cell wall and lipopolysaccharides in the outer membrane of Gram-negative bacteria, disrupting the internal cytoplasmic membrane.176–179 The presence of foreign bodies has been reported to reduce the phagocytic and bactericidal functions of white blood cells, thereby reducing the number of bacteria required to cause an infection to less than 1/100,000. 180 Implant-related infections are a major complication of orthopedic surgery, which often uses implants. In particular, periprosthetic joint infections can lead to persistent pain, large bone defects, additional surgery, extended hospital stays, and increased medical expenses. Severe infections may require revision or amputation of the artificial joint, significantly affecting the patient's quality of life.47,181,182 To date, several studies have reported the antiviral functions of TiO2 has been reported in several studies.183–186 Antiviral TiO2 coated on Ti implant materials suppresses viral infection through UV light exposure during surgery. In addition, a water droplet spreads completely over the entire surface of TiO2 under UV illumination, with the contact angle approaching zero,157,187 called superhydrophilicity. Sakai et al. proposed that superhydrophilicity is due to a newly formed oxygen defect formed after the surface oxygen traps the two holes generated by the photocatalysis reaction. 188 Hydrophilic surfaces are expected to activate biochemical reactions, such as cell integration and proliferation, that occur in the body's fluid environment because the surface of TiO2 is covered by OH− ions that react with the mineral constituents (Ca2+ and (PO4)3–) on the bone surface. For example, hydrophilic TiO2 reacts with chitosan to form a stable composite, 189 enhances bone regeneration and osseointegration around implants, 190 and promotes biochemical reactions resulting in increased cell proliferation and osseointegration. 191 Recent research has demonstrated that TiO2 fabricated on Ti-based implants significantly enhances osseointegration.192–194 Tengvall and Lundström proposed a model for the Ti-tissue interface, 195 in which Ti3+ ions formed by Ti4+ reduction facilitate the chemical bonding and interdiffusion of ions H2PO4− and Cl− in the absorbed protein, glycosaminoglycans, inorganic ions and other macromolecules.
Equilibrium diagram of TiO2 (tP6, P42/mnm, a=0.45924 nm, c=0.29575 nm) with (a) CaO (cF8, Fm-3m, a=0.48032 nm) and (b) P2O5 (hR84, R3c, a=b=1.03035 nm, c=1.35102 nm). TiO2 undergoes a eutectic reaction with mN CaTiO3 at 1417°C and with Ti5P6O25 at 1165°C. As a result, TiO2 coexists in two phases with CaTiO3 or Ti5P6O25 below the eutectic temperature
Schematic illustration of the photocatalytic reaction of anatase-structured TiO2 semiconductor and the related reaction equation. Upon UV light illumination corresponding to the band gap energy (1), the excited electrons and holes generated in equation (2) react with oxygen in equations (3) to (7) and water in equation (8), respectively, to generate hydroxyl radicals. When the charge carriers recombine at lattice defects, they generate heat in equation (9)
Table 2 representative oxidizing agents and their oxidation potentials relative to the standard hydrogen electrode and chlorine. The middle column displays the oxidation potential of various substances, while the right column indicates their relative oxidizing power, with the oxidizing strength of chlorine set to 1 as a reference. The hydroxyl radical ranks second on the list, showing an oxidizing power approximately twice that of chlorine.
Ti and its alloys have been used in orthopedic and dental surgeries because of their biocompatibility with tissues13,195 and corrosion resistance in bodily fluids. These properties are endowed by a spontaneously formed stable TiO2 passive layer with a thickness of several nanometers on the surface. The thin layer consisted of amorphous TiO2 and small amounts of sub-oxides (TiO and Ti2O3) close to the metal-oxide interface. 196 This TiO2 is in thermodynamic equilibrium with Ti and its alloys 197 and bonds strongly with the substrate. 198 The bone tissue on an osseointegrated Ti implant comes in contact with TiO2 and not Ti metal; however, the spontaneously formed oxide contains abundant defects and is too thin to endure external stress. 199 Even if a new oxide layer evolves on its surface after disruption, the rate of repassivation to prevent corrosion is very slow, 200 and the surface is exposed to corrosive body fluids under incessant stress, 201 suggesting that the spontaneously formed thin TiO2 is insufficient as a biomaterial. Various methods have been applied to form durable, thick, and stable TiO2 on Ti-based biomaterials, including hydrothermal treatment,191,194 sol-gel processes, 202 thermal chemical vapor deposition, 203 thermal spraying, 204 sputtering, 205 anodization,206–208 and thermal oxidation. 209 The quality of TiO2 varies depending on the desired biomaterial function. To achieve antiviral properties, the recombination of photogenerated charge carriers should be suppressed to increase the quantum efficiency of the photocatalyst. Therefore, reducing the number of lattice defects (oxygen vacancies, Ti3+ atoms adjacent to O vacancies, interstitial atoms, and interfaces) that become recombination sites for TiO2 is crucial. 210 In other words, preparing TiO2 with few defects and a high crystallinity 211 is necessary to improve its photoinduced characteristics. By contrast, biocompatibility requires a chemical composition with few impurities. Because Ca and P, the main constituent elements of the bone, form a solid solution in TiO2, the chemical composition of the oxide is important. 212 TiO2 is a safe substance used in food additives and cosmetics and has a favorable effect on the differentiation and generation of osteoblasts. The anodization of Ti and its alloys not only forms a biofunctional TiO2 but also establishes a strong bond with the substrate due to thermodynamic equilibrium. By adjusting the electrochemical conditions during the anodization, it is possible to enhance the photoinduced properties by increasing the crystallinity of the oxide or by incorporating dopants into the oxide.
TiO2 nanotubes (TNT) prepared by anodization in fluorine-contained electrolyte exhibits biocompatibility and cell viability, which is varied with the geometrical microstructure (cell length and diameter) of the oxide.213,214 Baishya et al. demonstrated that cell-compatible ZnO2 coating on TNT promotes proliferation of osteoblastic cells due to shading impurity F-based species. 215 Similarly, Zn or Sr doping to TNT by hydrothermal treatment enhanced early and faster differentiation of the adhered adipose derived stem cells into osteoblasts. 216 TNT is consistently formed in Ti alloys when the alloy is single-phase; however, this uniformity is not achieved in multi-phase alloys. For example, pore diameters of the nanotubes on top of the β-phase are smaller than those on top of the α-phase. 217 The mechanical properties of Ti alloys coated with nanotubes are crucial for practical applications. Tribocorrosion tests conducted on the nanotubular surface of pure Ti revealed that the open circuit potential (OCP) showed insufficient recovery after the tests were completed. 218 Ogawa et al. reported that Ti etched with sulfuric acid at 120 °C for 75 s showed increased rates of attachment, spreading, proliferation, and differentiation of rat bone marrow-derived osteoblasts when stored in a container for 4 weeks and then exposed to ultraviolet (UV) light for 48 h. This enhancement was attributed to the removal of hydrocarbons from the titanium dioxide (TiO2) surface, the thickening of the oxidized layer, and improved integration between the bone and Ti. 219 Although the effect of hydrocarbons on bone integration remains to be elucidated, this study suggests that an uncontaminated, thin TiO2 surface is involved in osteoconductivity. Kokubo et al. demonstrated that treatment of anodized Ti with hot water (80 °C, 48 h) resulted in the formation of numerous Ti-OH groups on the surface, which promoted HAp formation in SBF solution. 220 The mechanism can be explained by two factors. First, the coherent crystal structure between rutile-structured TiO2 and HAp facilitates bioactive bonding with living bone. Second, there is an electrostatic interaction between the negatively charged surfaces, which results from hydration, and the cations present in the SBF solution. 221 Both acid etching and anodization encourage the formation of highly crystalline oxides due to the heat generated during these processes. This improves the photoactivity and superhydrophilicity of pure TiO2 by reducing lattice defects. The photoactive physical function and superhydrophilic chemical properties associated with hydroxyl groups serve as an intermediate layer between Ti-based metal implants and bone tissue, providing antibacterial and osteoconductive benefits.
Mechanical properties
The primary reasons for using metals in biomaterials are their high strength and corrosion resistance in load-bearing environments of the human body. Dental and orthopedic materials must be strong enough to withstand the forces of biting/mastication and the stress imposed on the body, respectively. Considering dental materials are typically exposed to corrosive environments, 222 precious metals with lower ionization tendencies (such as Au, Pd, and Ag) are used as dental materials. The average pH of saliva and body fluids, such as blood and lymph, is weakly acidic at 6.8 and slightly alkaline at 7.40 ± 0.05, respectively, but varies depending on the environment. For example, after drinking a carbonated beverage (pH 2.45), the pH fell below 5.5 and 4 for a greater proportion of the time on the buccal surface of the upper molar in patients with erosion than in the controls (p<0.05). 223 Furthermore, when wear debris is engulfed by macrophages, the biological response to wear debris in total joint replacements and enzymes and metabolites released by dying macrophages cause significant acidification (decrease in pH) of the surrounding tissue. 224 The number of H+ ions increase with decreasing pH, and elements with a higher ionization tendency than H become cations, a process that does not occur with precious metallic dental materials; however, some elements can be eluted from surgical biomaterials. Metallic ions released into the oral cavity contact adjacent tissues and cause adverse reactions, including toxicity, allergy, and mutagenicity. 225
Most precious metals used as dental materials have a face-centered cubic (fcc) crystal structure, which exhibits high workability and relatively low strength. The binary equilibrium phase diagrams of precious metals are generally divided into three categories: complete solid solution type, reaction type (eutectic/monotectic), and low-temperature ordered phase formation. The atomic diameters of Au, Ag, Pt, Pd, and Cu are 0.408, 0.407, 0.392, 0.389, and 0.362 nm, respectively. As a result, it is expected that solid-solution hardening will occur in copper-added alloys. In contrast, the Au-Cu, 226 Cu-Pt, 227 and Cu-Pd 228 form ordered phases that enhance their strength. In a representative cast of a restorative Au-Ag-Cu-based alloy, 229 the order-disorder transformation does not occur when the Au content exceeds 88.4 wt.%. The corrosion resistance of the alloy decreases when the Au content falls below 60 wt. %. Additionally, the Ag content should be limited to less than 10 wt. % to avoid uneven distribution of Ag within the alloy. Other representative casting alloys for restorative dentistry that are based on Au-Ag-Pd do not undergo age-hardening, as they form a complete solid solution across all composition ratios. 230 Dental components, such as inlays and crowns, are typically manufactured using the melt-and-cast method; therefore, both flowability and mold releasability are essential for the alloy. Increasing the Pd content in the alloy decreases fluidity due to its high viscosity and diminishes castability because of gas absorption. As a result, ceramic crucibles with lower gas emissions are utilized in place of graphite crucibles; however, the formation of oxides on the alloy surface hinders the melting process.
The expected functions of orthopedic materials differ based on their components. For instance, stents need to have high circumferential rigidity and significant longitudinal bending stiffness. 231 In the case of cerebral aneurysm clipping, it is essential to have a strong closing force along with the flexibility of the clipping blades. 232 Similarly, hip prosthesis stems and spinal fixation devices require materials with low Young's modulus while also maintaining high strength. 233 Achieving a balance between strength and deformability is crucial for orthopedic devices. Recently, Ti and its alloys have increasingly replaced stainless steels and Co-Cr alloys. Strengthening Ti is achieved by increasing the fraction of the hexagonal close-packed (hcp) α phase in the alloy. This can be done by adding α stabilizer (Al, Sn, O, N) and controlling the conditions of thermomechanical treatment. Figure 13 shows the relationship between tensile strength and Young's modulus in metallic materials, highlighting a trade-off between the two properties. When it comes to an artificial hip stem, the proximal portion must have high strength to support body weight, while the distal portion should have a low Young's modulus similar that closely matches that of bone. This is important to minimize stress shielding. The key strengthening element in Ti alloys is Al, and the Ti-6Al-4V alloy with α+b dual phases is widely used as an orthopedic biomaterial. Thompson and Puleo found that ions eluted from the Ti-6Al-4V alloy inhibited the normal differentiation of bone marrow stromal cells into mature osteoblasts in vitro. 234 V ions released from the Ti-6Al-4V alloy can enter blood vessels 235 and cause DNA and nuclear damage. 236 The adverse effects of inorganic vanadium compounds in mammalian cell cultures are considered carcinogenic, immunotoxic, and neurotoxic, 237 leading to significant endothelial cell damage. 238 In a systematic study of the toxicity of metallic ions, 39 the relative growth rate of cells was found to decrease with increasing metal ion concentrations, except for several metals (Figure 14). These studies indicate that ensuring the success and safety of metallic implants requires careful consideration of both mechanical properties and cytotoxicity in material selection.
(a) Plot of Young's modulus vs. tensile strength for metallic biomaterials. Young's modulus and tensile strength are in a trade-off relationship, and artificial hip joint stem materials are required to have a low Young's modulus to suppress stress-shielding (b). TiNbSn has succeeded in achieving both a low Young's modulus and high strength through component design, orientation control, and structural control by gradient heat treatment. 264
Effect of various metallic concentrations on the relative growth ratio of (left) L929 and (right) MC3T3-E1. Increasing concentrations of Ni, Co, Cr, and Fe ions significantly decreased cell viability, whereas increasing concentrations of Ti, Nb, Ta, Zr, and Sn ions had little effect on cell viability. Reprinted with permission from reference. 39
The essential properties of dental and orthopedic implant materials include wear resistance to minimize wear debris generation and corrosion resistance to prevent metal ion elution. The tribocorrosion test enables simultaneous in situ measurements of properties, highlighting the combined degradation due to wear (mechanical tribology) and corrosion (chemical reaction).239,240 Implant failure can lead to exacerbation from biomechanical actions, and tribocorrosion may synergize to facilitate crack propagation. 241 In hip joint prostheses, sliding wear occurs between the acetabular cup and the femoral head, while fretting wear occurs between the femoral stem and the femur, or between the femoral head and the femoral stem.. 242 Sliding wear occurs in areas with large displacement amplitudes (several tens of millimeters) between friction surfaces while fretting wear takes place when there is a small displacement amplitude (several hundred microns) between them. 243 Corrosion accompanies both types of wear in the presence of body fluids 244 ; therefore, the implant surface must be resistant to corrosion and wear caused by bodily fluids. The relationship between hardness and relative wear resistance in metals and alloys indicates that the wear resistance of pure metals and steels increases linearly with hardness.245,246 This implies that enhancing the surface hardness is essential for improving wear resistance. To enhance corrosion resistance, precious metals with low ionization tendencies can be utilized to form a passive layer on the surface. Since precious metals used in dental materials typically have low hardness, they can strengthen the matrix through precipitation or solid-solution hardening. In contrast, the surfaces of orthopedic implant materials can be strengthened through surface modification. This difference arises from the distinct approaches used to fabricate implant devices; precious metal dental materials are produced through the melt-and-cast method, while transition metal orthopedic materials are fabricated using thermomechanical processing. In summary, when implant materials lack adequate corrosion and wear resistance, implementing surface modifications becomes a crucial strategy to significantly improve their performance. This approach not only enhances the durability of the implants but also guarantees optimal performance within the body. Given the recent rise in the use of Ti and its alloys as implant materials, the following paragraph will address surface modifications that aim to improve their mechanical properties.
A drawback of Ti and its alloys is their poor wear resistance,247,248 which leads to the generation of wear debris and corrosion products near the implant, causing inflammation249,250 and loosening of the implant.251,252 Many studies have investigated the surface modification of Ti and its alloys to enhance their wear and corrosion resistance. Techniques used for this purpose include nitrogen ion doping, 253 TiN coating, 254 TiN electrodeposition, 255 friction hardening, 256 acid etching, 257 and anodization. 258 Anodic oxidation has gained significant attention due to its ability to form a physical barrier against tribocorrosion259,260 and to enhance osteoblast adhesion 261 and osseointegration.156,262,263 Kubota et al. reported 96 that an anodized TNS alloy, which possesses a low Young's modulus and high strength, 264 exhibited superior fretting tribocorrosion properties in simulated body fluids compared to titanium. The enhanced performance can be attributed to the development of a thick, durable, and well-adhered layer of rutile-structured TiO2. In comparison to the anodized Ti-6Al-4V alloy, the anodized TNS alloy demonstrated superior mechanical properties, photocatalytic effects, osteoconductivity, and antibacterial properties. 87 This superiority can be attributed to several key factors: the hard, thick, and well-adhered oxide layer enhances wear resistance; the chemically stable oxide layer offers corrosion resistance; the porous oxide layer promotes osteoconductivity; and the highly crystallized oxide layer contributes to antibacterial properties through photoinduced functionality. These properties depend on the quality of anodic oxide, including its crystallinity, porous structure, and chemical composition. This enhancement is achieved by carefully controlling the electrochemical conditions during the anodizing process.
Antibacterial properties
Infections in orthopedic implants can lead to severe consequences, including the need for costly hospitalization and revision surgery.265,266 The global infection risk in orthopedics has been reported as 2–5%. 48 Staphylococcus aureus is a cause of surgical wound infections and, together with Staphylococcus epidermidis, can cause infections associated with implanted medical devices. 267 Despite the biocompatibility of Ti and its alloys commonly used for implants, these surfaces are highly susceptible to bacterial colonization. 268 Once bacteria adhere to implant surfaces, they rapidly multiply and form a biofilm that protects them from antibiotics and immune responses.269–271 Thus, early intervention is critical to prevent biofilm development. Various approaches have been employed to control bacteria upon contact. Bactericidal coatings eliminate pathogens directly, unlike anti-fouling coatings, which only delay bacterial adhesion and may allow recolonization.272,273 However, it's essential to recognize that completely killing all pathogenic microbes to eliminate the risk of future infections remains challenging with current strategies. Bactericidal coatings can effectively reduce the immediate risk of infection and, when combined with a healthy immune system, may eliminate the risk. Silver, the most commonly used bactericidal agent in clinical settings and research, offers broad-spectrum antimicrobial properties that are highly effective against device-related infections. 274
Table 3 shows a summary of the antimicrobial processing technology for Ti-based alloys. Silver's broad-spectrum antimicrobial properties have made it a valuable tool in controlling infections associated with medical devices. 274 In orthopedics, one standard method to imbue orthopedic materials with antimicrobial qualities without significantly altering or coating the implant's surface is through a process known as silver implantation, mainly silver plasma immersion ion implantation [Ag-PIII]. 275 Ag-PIII has demonstrated its effectiveness in controlling relevant bacterial species and promoting bone growth in laboratory settings (in vitro) and living organisms (in vivo). During the Ag-PIII process, silver nanoparticles are created and embedded in substrates to ensure their long-term presence.275,276 These Ag-PIII surfaces have been proven to be bactericidal rather than simply anti-fouling. This distinction was demonstrated by collecting adherent bacteria and growing them in fresh culture, resulting in a remarkable 99% reduction in viability for S. aureus. 277 The Ag-PIII surface finish also showed good compatibility with osteoblasts and good bone formation at the interface with the metal, as observed in the rat model. 278 Ag-PIII has also been combined with other elemental implantations, such as magnesium, zinc, and calcium, to enhance antibacterial and bone-forming properties. 279
Table 3 summary of antibacterial treatments for titanium-based alloys. This table summarizes various antimicrobial strategies, including silver ion implantation, metal oxide coatings, and TiO2 nanotubes, detailing their associated materials, processing technologies, antibacterial mechanisms, and bone regeneration-related outcomes.
This table summarizes antimicrobial methods applied to titanium alloys, each with unique mechanisms and clinical benefits.
Metal oxides including TiO2, copper oxide (CuO), and zinc oxide (ZnO) have also been widely introduced to produce antimicrobial coatings.280–282 A previous study reported that sonochemically prepared CuO and ZnO nanoparticles exhibit a highly crystallized structure after heat treatment, resulting in higher ROS generation activity and stronger antimicrobial performance. 283 In general, metal ion/oxide antimicrobial coatings are an applicable technology for the antimicrobial coating of orthopedic implants due to their excellent heat resistance, chemical stability, long antimicrobial efficacy period, and broad antimicrobial spectrum. In particular, TiO2 nanotube coatings have emerged as a fascinating avenue to combat bacterial adhesion. These coatings also exhibit promising interactions with osteogenic cells, potentially serving a dual purpose in enhancing osseointegration. TiO2 nanotubes were synthesized using a Ti surface through an anodization process involving sodium fluoride and sulfuric acid. To introduce antimicrobial properties, the nanotubes were further anodized with silver nitrate. Notably, human osteoblast precursor cells adhered equally well to nanotube coatings with or without the silver nitrate coating. However, the viability of Pseudomonas aeruginosa (P. aeruginosa) was significantly reduced, demonstrating the antimicrobial effectiveness of the silver-coated nanotubes. 284 Subsequent research on TiO2 nanotubes, such as that conducted by Peng et al., has shown promising outcomes. They applied a straightforward nanotube array to Ti substrates and evaluated the coating's osteogenic and antibacterial properties. The presence of nanotubes improved the adhesion of osteogenic cells, with better adhesion observed on smaller-diameter nanotubes. Interestingly, the adhesion of Staphylococcus epidermidis followed the opposite trend, with smaller nanotubes inhibiting its adhesion. These findings suggest that optimizing nanotube properties can yield coatings conducive to bone formation and infection resistance. 285 Another system involved the creation of TiO2 nanocolumns on Ti-6Al-4V surfaces oriented approximately perpendicular to the substrate. These densely packed columns had a minimal effect on the attachment of osteoblasts but significantly reduced surface coverage and biofilm formation by various clinical strains of Staphylococcus aureus. 286 Recent advances in TiO₂ nanotube coatings have demonstrated strong photocatalytic antiviral activity, particularly against SARS-CoV-2. Under UV light, TiO₂ nanoparticles and nanotubes can inactivate up to 99.99% of the virus. 287 Doping TiO₂ allows activation under visible light, expanding indoor applicability. 288 Among various nanostructures, nanotubes show higher virucidal effects. 289 Composite coatings like silver-anatase TiO₂ enhance antiviral efficacy. 290 These coatings offer potential as self-disinfecting surfaces to reduce fomite-based transmission of SARS-CoV-2 and its variant.291,292
Nanotubes offer a versatile modification platform that can provide alternative antimicrobial and osteogenic mechanisms. One promising approach involves the incorporation of zinc, an essential element in multiple bone formation processes known for its bactericidal properties. In this method, Zinc ions are loaded onto a titania nanotube surface coating, aiming to enhance osseointegration while preventing infections. 293 The introduction of zinc-loaded nanotubes resulted in several improvements, both in vitro and in vivo. In laboratory settings, these modified nanotubes showed enhanced osteogenic markers, including increased alkaline phosphatase activation, mineral deposition, and elevated mRNA expression levels of essential bone-related proteins like collagen type I, osteocalcin, and osteopontin. 294 In vivo experiments conducted in a rodent tibial insert model demonstrated a significant increase in osseointegration facilitated by the Zinc-loaded nanotubes. 295 Additionally, the Zinc acted as a potent inhibitor of Staphylococcus aureus growth. The slow release of zinc ions over a month indicated longer-term bactericidal activity. Furthermore, researchers have explored other modification strategies, such as incorporating Ag and Cu into nanotubes to serve as bactericidal agents.296,297 These approaches leverage nanotubes’ high surface area and loading capacity to control bacterial infections effectively.
The development of antibacterial metals has included the use of antibacterial metal ions like Ag and Cu, as well as antibacterial agents like iodine and vancomycin.298–301 However, there is a concern that these added toxic substances could leach into the bloodstream, leading to possible side effects. 49 TiO2 has previously been recognized as a photocatalytic material and is commonly used for air and water purification because it can break down redox species when exposed to light with energy corresponding to its bandgap. 302 In the medical field, there have been reports of TiO2 exhibiting antibacterial properties and even demonstrating an anti-tumor effect.303,304 When water is irradiated on TiO2 with ultraviolet (UV) light, it undergoes decomposition through the photocatalytic effect, generating ROS such as hydroxyl radicals (•OH), superoxide anions, and hydrogen peroxide (H2O2). 301 ROS have been shown to possess antibacterial and anti-tumor effects by disrupting the structure of bacteria and tumor cells. 305 In contrast, TiO2 is generally considered a stable and safe substance used as an additive in food and drugs. 306 Coating biocompatible alloys with TiO2 has the potential to be an ideal technology, offering antibacterial performance through UV irradiation while maintaining inherent safety and stability.
There are limited reports on the antimicrobial performance of anodic oxidation in Ti alloys such as Ti-6Al-4V alloys. 286 Achieving antimicrobial activity by anodic oxidation in Ti-6Al-4V alloys, widely used as orthopedic implants, is still challenging. The study by Kurishima et al. showcased the photocatalytic activity of anodic oxides produced from TNS alloys using electrolytes containing sodium tartrate. These resulting anodic oxides possess a porous microstructure, including TiO2, with well-crystallized anatase and rutile structures.184,307 The research confirms that these anodic oxides are highly effective in generating •OH and demonstrate photocatalytic activity when exposed to UV irradiation. The superiority of antimicrobial performance by photocatalytic activity is that it is possible to achieve antimicrobial performance regardless of drug resistance. In this study, in addition to Staphylococcus aureus, which is a typical organism that causes inflammation in orthopedic surgery, methicillin-resistant Staphylococcus aureus and Escherichia coli also exhibited excellent antimicrobial performance (Figure 15). Anodic oxidation of TNS alloys in a sulfuric acid electrolyte to impart antimicrobial activity has also been investigated. 308 Similar to anodic oxidation with sodium tartrate, anodic oxidation with sulfuric acid electrolyte was also confirmed to have antimicrobial activity.
Schematic model shows an antimicrobial effect due to reactive oxygen radicals generated from H2O by UV irradiation on the anodized TiO2 surface. Photographs showing the results of antimicrobial assays on anodized TNS alloys and glass plates under low doses of ultraviolet irradiation. MSSA: methicillin-sensitive Staphylococcus aureus, MRSA: methicillin-resistant Staphylococcus aureus, E. coli: Escherichia coli. Reproduced with permission from reference. 308
On the other hand, in the anodic oxidation of Ti-6Al-4V alloy with sodium tartrate, the properties of the anodic oxide film generated were lower than those of TNS alloy, and photocatalytic activity and antibacterial effect by UV irradiation could not be confirmed. 264 Surface modification techniques that leverage the photocatalytic properties of Ti alloy anodic oxides have shown promising results in terms of antimicrobial performance, especially under UV irradiation conditions. This suggests their potential utility in various applications where antibacterial properties are desired. Antimicrobial performance is also being investigated for other orthopedic implant material metals. The incorporation of metal ions such as zinc, aluminum, manganese, copper, and zirconium into the Mg matrix at low doses has been investigated to overcome the rapid degradation rate of Mg-based implants.309,310 In particular, researchers have recently shown that combining Mg with Al and Cu [Mg-1Al-Cu] may reduce the in vitro degradation of Mg-based implants and simultaneously introduce antimicrobial properties. 311 On the other hand, no satisfactory research results have yet been reported on the loading of antimicrobial performance by anodic oxidation of magnesium, and this is a subject for future research. Ta is one of the most promising next-generation materials for orthopedic implants, but opinions are divided on the antimicrobial performance of Ta itself. 312 There is also a report on the antimicrobial performance of an anodized Ti-Ta alloy, which has shown excellent antimicrobial performance against Staphylococcus aureus. 313 While research on surface treatment for imparting antimicrobial performance, including anodic oxidation, has been conducted for other metallic materials, research has been accumulated on Ti and Ti alloys. The TNS alloys have the potential to simultaneously improve bone affinity and support antimicrobial performance. Clinical application is expected to be realized by accelerating research in the future.
Conclusions
Advances in medical technology have made it possible to treat diseases previously considered difficult to cure, and, at the same time, engineering technology is increasingly contributing to medical care. Implant metallic materials for treatment have also attracted attention, with their functions playing a crucial role in improving patients’ QOL. In this review, we outlined the role of Ti-based implant materials and introduced materials with excellent biofunctions. We expect this review will deepen our understanding of implant materials.
Footnotes
Acknowledgements
The authors would like to express sincere gratitude to Prof. Hideyuki Kanematsu from National Institute of Technology, Suzuka College, Japan for his invaluable support in completing this review article.
Declaration of conflicting interests
The authors declared no potential conflicts of interest with respect to the research, authorship, and/or publication of this article.
Funding
This study was conducted using the research resources of the Japan Society for the Promotion of Science (20H02458).
Japan Society for the Promotion of Science, (grant number 20H02458).
