Titania (TiO₂) nanotubes (TNTs) can be fabricated on Ti or its alloys via electrochemical anodization (EA) ^[18][35]. Briefly, EA involves the immersion of a Ti implant as an anode and a bare Ti/Pt electrode (cathode) inside an electrolyte (containing fluoride and water), with the supply of adequate current/voltage ^[19][31]. Under controlled and optimized conditions and the attainment of an equilibrium (characterized by metal oxide formation and dissolution), the self-ordering of TiO₂ nanotubes (like test-tubes, open at the top and closed at the bottom) or nanopores (nanotubes fused together, with no distance between them) on the entire surface of the implant occurs ^[20][36]. It is noteworthy that EA represents a cost-effective and scalable Ti implant surface modification strategy. Recent attempts to optimize EA to enable clinical translation include fabrication of controlled nanostructures on clinical dental implants ^[20][36], superior mechanical stability (nanopores > nanotubes) ^[21][37], and fabrication of dual micro-nano structures ^[22][38] by preserving the underlying ‘gold standard’ micro-roughness of dental implants ^[23][39]. It is worth noting that EA is a versatile technique that can be used to nano-engineer controlled topographies on various biomedical implants, spanning various metals and alloys, including Ti ^[24][40], Ti alloys ^[25][41], Zr ^[26][42] and Al ^[27][43]. A schematic representation of TNTs and their various characteristics and research challenges is shown in Figure 1.

Attributed to improved bioactivity and the ability to load and release proteins/growth factors, TNTs are a promising surface modification strategy for orchestrating osteogenesis, as established by various in vivo investigations ^[29][30][29,45]. The incorporation of fluoride ions into TNTs during anodization and the mechanical stimulation of osteoblasts also contribute towards the enhancement of osseointegration ^[31][46]. Further, to ensure the successful establishment and maintenance of osseointegration, TNTs on Ti implants have loaded with various orthobiologics, including bone morphogenetic protein-2 (BMP-2) ^[32][47], platelet-derived growth factor-BB ^[33][48], alendronate ^[34][49], ibandronate ^[35][50], N-acetyl cysteine (NAC) ^[36][51], and parathyroid hormone (PTH) ^[37][52]. Lee et al. loaded TNT-modified dental mini-screws with N-acetyl cysteine [NAC, a reactive oxygen species (ROS) scavenger with anti-inflammatory and osteogenic properties], implanted them in rat mandibles in vivo and, at 4 weeks, observed significantly enhanced osteointegration at the NAC–TNT sites ^[36][51]. In another study, machined dental implant screws were modified with HF etching and EA to fabricate dual micro- and nanotubular structures, which, upon implantation in ovariectomized sheep in vivo for 12 weeks, showed significantly increased pull-out force and bone-implant contact ^[38][53]. Further, various nanoparticles, ions or coatings of Sr ^[39][54], Ta ^[40][55], La ^[41][56], and Zn ^[42][57] onto/inside TNTs have also shown upregulated osteogenic outcomes.

It is worth noting that various ions or NPs have exhibited favorable osseointegration through their use in in vitro and in vivo investigations; however, these may illicit immunotoxic reactions in a dose-dependent manner and remain the subject of active research. Further, with respect to bone-forming proteins, future investigations into the estimation of the local need for bioactive agents and the evaluation of their release inside the bone micro-environment are needed.

Studies relating to the use of TNTs for enhancing STI for dental implants are very limited, as reviewed elsewhere ^[43][58]. Recently, Gulati et al. reported the enhanced proliferation and adhesion of human gingival fibroblasts (HGFs) on dual-micro-nano anisotropic TiO₂ nanopores ^[22][38]. Further, beginning at 1 day of culture, the HGFs started to align parallel to the nanopores; and the gene expression analysis (type I collagen, type III collagen and integrin β1) indicated a wound-healing profile that promoted substrate–cell and cell–cell interactions ^[44][59]. Further, anodization combined with heat treatment has also been used to upregulate fibroblast activity. Briefly, the proliferation and adhesion of gingival epithelial cells were enhanced on heat-treated anodized Ti surfaces, which was attributed to hydrothermal treatment precipitation of hydroxyapatite crystals ^[45][60]. Alternatively, hydrothermally treated TNTs have been reported to upregulate the integrin α5 and β4 expressions of gingival epithelial cells ^[46][61], the adhesion of murine fibroblast-like NIH/3T3 cells and the expression of adhesion kinase ^[47][62], as compared to unmodified TNTs.

The biofunctionalization of TNTs has also been explored in order to enhance the functions of fibroblasts and epithelial cells towards augmenting STI. For instance, Xu et al. reported that the inhibition of human gingival epithelial cells on TNTs was reversed when the electrochemical deposition of CaP was performed on TNTs, which was attributed to the local elution of Ca and P ions ^[48][63]. Next, Liu et al. investigated the influence of bovine serum albumin (BSA) loading inside TNTs on HGF functions ^[49][64]. Unmodified TNTs promoted early HGF adhesion and COL-1 secretion; however, BSA-TNTs enhanced early HGF adhesion, while suppressing late proliferation and COL-1 secretion. It is interesting that contradictory behaviors among bioactive coatings on TNTs have been reported and further in-depth investigation into the influence of these modifications on the STI performance is needed. Furthermore, the local elution of fibroblast growth factor-2 (FGF-2, immobilized on Ag nanoparticles) from TNTs effectively enhanced the proliferation, adhesion and extra-cellular matrix formation in the cultured HGFs ^[50][65]. Augmented proliferation, adhesion, and expression of VEGF and LAMA1 genes in vitro was observed, which were pronounced after the loading of 500 ng/mL of FGF-2.

The local release of therapeutics from TNTs has been widely explored towards optimizing the loading and local elution of potent antibacterial agents ^[51][30]. It is noteworthy that within minutes of implantation, saliva proteins adhere to the dental implant, forming a pellicle, and early colonizers such as Streptococci adhere to these pellicles within 48h ^[52][66]. This can be followed by secondary colonizers, including Fusobacterium nucleatum, Aggregatibacter actinomycetemcomitans and Porphyromonas gingivalis ^[53][67]. These bacteria can further lead to peri-implantitis ^[54][68]. Once a biofilm is established, the routine administration of antibiotics is insufficient and, hence, local therapy using dental implants has been proposed. Further, TNTs can enhance bacterial adhesion due to their nano-scale roughness, increased number of dead bacteria and amorphous nature. Hence, the synergistic antibacterial functions of TNT-modified dental implants are needed to prevent bacterial colonization and implant failure. Further, the size and crystal structure of TNTs influences bacterial adhesion properties. Ercan et al. investigated the influence of the size and the heat treatment of TNTs on their antibacterial effect and reported that heat-treated and 80 nm diameter TNTs exhibit strong antibacterial effects ^[55][69]. Similarly, when comparing 15, 50 and 100 nm diameter TNTs, the lowest number of adherent bacteria were reported on the smallest-diameter TNTs ^[56][70]. Further, annealed TNTs show the best bactericidal response, as reported by Mazare et al. ^[57][71] and Podporska-Carroll et al. ^[58][72].

Various commonly prescribed antibiotics including Gentamicin ^[59][73], Vancomycin ^[60][74], Minocycline, Amoxicillin, Cephalothin ^[61][75], Cefuroxime ^[62][76] and Cecropin B ^[63][77] have been incorporated inside TNT-modified Ti implants to enable local antibacterial functions. Further, to target methicillin-resistant Staphylococcus aureus (MRSA), antimicrobial peptides (AMPs) such as HHC-36 have been loaded inside TNTs to achieve a bactericidal effect of almost 99.9% against MRSA ^[64][78]. Biopolymer coatings have also been applied to antibiotic-loaded TNTs to: (a) control drug release, (b) promote bioactivity, and (c) harness the inherent antibacterial property of biopolymers in order to provide long-term antibacterial functions. As a result, bare/drug-loaded TNTs have been modified with chitosan ^[65][79], polydopamine ^[66][80], silk fibroin ^[67][81] and PLGA (poly(lactic-co-glycolic acid)), which exhibited synergistic bioactivity and antibacterial enhancements. In addition, various antibacterial ions and nanoparticles (NPs), such as Ag ^[68][82], Au ^[69][83], Cu ^[70][71][84,85], B, P, Ca ^[72][86], Ga ^[73][87], Mg ^[74][88], ZnO ^[75][89], etc., have also been immobilized on or incorporated inside TNTs, with or without the use of hydroxyapatite or biopolymers, using techniques such as micro-arc oxidation, chemical reduction, photo-irradiation, spin-coating, and sputtering.

Multiple synergistic therapies, including osseointegration, immunomodulation, soft-tissue integration and antibacterial functions can also be enabled using nano-engineered Ti with TNTs. For instance, TNTs modified by Ag via plasma immersion ion implantation (PIII) showed excellent antibacterial effects against P. gingivalis and A. actinomycetemcomitans, while enhancing the bioactivity of epithelial cells and fibroblasts in vitro and reducing inflammatory responses in vivo ^[76][90]. Similarly, hydrothermally doped Mg-TNTs exhibited upregulated osteoprogenitor cell adhesion and proliferation (without cytotoxicity) and suppressed osteoclastogenesis, while showing long-lasting antimicrobial effects against methicillin-susceptible S. aureus (MSSA), methicillin-resistant S. aureus (MRSA) and E. coli ^[74][88].

The modulation of the host immuno-inflammatory response is crucial to the timely establishment of osseointegration. Hence, attempts have been made to obtain immunomodulatory functions from modified TNTs ^[28][44]. These include the influence of physical/chemical characteristics and the local elution of anti-inflammatory drugs from TNTs. The influence of Ti nanotopography on immune cells, including macrophages, monocytes and neutrophils, has supported the attenuation of inflammation ^[77][78][91,92]. Clearly, the presence of nano-scale cues controls macrophage adhesion and inflammatory cytokine production. Similarly, in vitro cultures of such cells on TNTs have also established the influence TNTs nanotopography on immuno-inflammatory responses ^[22][38].

Smith et al. reported reduced functions (viability, adhesion, proliferation and spreading) of immune cells on TNTs, as compared with bare Ti ^[79][93]. Alternatively, other studies have shown enhanced nitric oxide and the absence of foreign-body giant cells on TNTs ^[79][80][93,94]. With respect to the nanotube diameters, inconsistent results have been obtained, with some studies indicating 60–70 nm diameters as the most immuno-compatible ^[79][81][93,95]. Further, Ma et al. compared the functions of monocytes/macrophages on nanotubes and polished Ti, and reported post-attachment stretching inhibition (repulsed adhesion), enhanced M2 phenotype (wound healing) and suppressed M1 phenotype (pro-inflammatory) polarization for TNTs anodized at 5V ^[82][96]. Furthermore, to understand the mechanism behind selective immunomodulation due to TNTs, Neacsu et al. reported that this effect is attributed to the suppression of the phosphorylation of MAPK (mitogen-activated protein kinase) signaling molecules (p38, ERK1/2, and JNK) on TNTs ^[83][97]. More recently, using 50 and 70 nm diameter anodized anisotropic TiO₂ nanopores, we showed that macrophage proliferation was significantly reduced on the 70 nm nanopores ^[22][38]. Further, the spread of macrophage on nanopores indicated an oval morphology, which was suggestive of an inactivated state.

The local elution of potent drugs, such as non-steroidal anti-inflammatory drugs (NSAIDs), bypasses the limitations associated with systemic administration (delayed bone healing and toxicity). These drugs have been loaded inside TNTs for the purpose of local release. Briefly, Ibuprofen ^[84][98], Indomethacin ^[85][99], Dexamethasone ^[86][100], Aspirin ^[87][101], Sodium naproxen ^[88][102], Quercetin ^[89][103], Enrofloxacin ^[90][104], Propolis ^[91][105] and immunomodulatory cytokines ^[92][106] have been successfully loaded and locally eluted from TNTs in vitro. Further, to achieve substantial loading and the delayed/controlled release of anti-inflammatory drugs, approaches including biopolymer coating on drug-loaded TNTs ^[93][94][107,108], polymeric micelle encapsulation of drugs prior to loading ^[95][109], the periodic tailoring of TNTs ^[96][110], the chemical intercalation of drugs inside TNTs ^[97][111] and trigger-based release ^[98][99][112,113] have been reported for TNT-based Ti implants. Additionally, metal ions and nanoparticles (NPs), including Au ^[69][83], Ag ^[100][114] and Zn ^[101][115] have also been incorporated on/inside TNTs to impart synergistic immunomodulatory functions with antibacterial or osteogenic activity. More recently, super-hydrophilic TNTs were fabricated via anodization and hydrogenation, and significantly reduced macrophage proliferation; upregulated M2 and downregulated M1 surface markers were exhibited on the modified TNTs, translating into effective immunomodulation and wound healing functionality ^[102][116]. It is also noteworthy that various in vivo tests of TNT modifications intended for use in various therapies, including antibacterial ^[51][30], osteogenic [29] or anti-cancer ^[103][117] applications, have established the immuno-compatibility of TNTs.

Ag NPs are one of the most widely used dentistry restoration and dental implant doping choices due to their outstanding antimicrobial properties ^[107][121]. Ag adheres to the bacterial cell wall and the cytoplasmic membrane electrostatically, which causes structural disruption ^[106][120]. This results in extensive damage to bacterial DNA, proteins and lipids, resulting in the inhibition of bacterial growth/viability and effective bactericidal action. Besides, Ag NPs can also stimulate osteogenesis and soft-tissue integration, making them an ideal choice for dental implant surface modification ^[108][122]. For instance, dental abutments modified with Ag NP suspension prevented C. albicans contamination, in comparison with the controls of unmodified abutments ^[109][123]. Further, citrate-capped Ag NPs offered bactericidal effects against S. aureus and P. aeruginosa ^[110][124]. Ti implants deposited with Ag NPs using anodic spark deposition have also been co-doped with Si, Ca, P and Na ions, to offer synergistic antibacterial (S. epidermidis, S. mutans and E. coli) and osteogenic (human osteoblast-like cells, SAOS-2) functions ^[111][125]. Similarly, to confirm that the used dosage of Ag NPs is safe, a culture of HGFs on Ag NPs/Ti was performed in vitro and the results confirmed no adverse effects ^[112][126]. Further, Ag NPs have also been immobilized on Ti implants pre-modified with hydroxyapatite ^[113][127], hydrogen titanate ^[114][128], chitosan/hyaluronic acid multilayer ^[115][129], nanoporous silica coatings ^[116][130], Pt and Au ^[117][131], and sandblasting and acid-etching ^[118][132] in order to achieve superior antibacterial and bioactivity effects. However, while Ag NPs offer effective antimicrobial action, they may cause cytotoxicity via the release of free Ag+ ions, ROS production, transport across blood-brain-barrier, and inflammation ^[106][120]. In a manner that is also applicable to other NPs discussed below, the toxicity of NPs depends on their chemical composition, surface charge, size and shape ^[119][133]

Like Ag NPs, Zn/ZnO NPs are not only antimicrobial but also osteogenic, hence their use in the modification of dental implants ^[104][118]. Zn is an essential element in all biological tissues and offers antibacterial effects against a wide range of microbes; however, its aggregation can cause cytotoxicity in mammalian cells ^[120][134]. To demonstrate its effectiveness against oral biofilms, Kulshrestha et al. reported that graphene/zinc oxide nanocomposite showed a significant reduction in biofilm formation ^[121][135]. Further, Hu et al. incorporated Zn into TiO₂ coatings on Ti implants through plasma electrolytic oxidation and observed superior bactericidal and bone-forming effects ^[122][136]. In 2017, Li et al. synthesized N-halamine labeled Silica/ZnO hybrid nanoparticles to functionalize Ti implants to enable antibacterial functions ^[123][137]. The hybrid NP-modified Ti exhibited excellent antibacterial activity against P. aeruginosa, E. coli and S. aureus, without any cytotoxicity against MC3T3-E1 preosteoblast in vitro. Recently, selective laser-melted porous Ti was biofunctionalized using Ag and Zn NPs via plasma electrolytic oxidation and tested against methicillin-resistant Staphylococcus aureus (MRSA) ^[124][138]. The results confirmed that 75% Ag and 25% Zn fully eradicated both adherent and planktonic bacteria in vitro and ex vivo. Further, Zn-modified Ti (0% Ag) enhanced the metabolic activity of preosteoblasts, indicating its suitability for dual osteogenic and antibacterial implant modification. Further, it is worth noting that ZnO NPs may cause cell apoptosis or necrosis and DNA damage ^[125][139]

CuO NPs offer advantages over Ag NPs, including cost-effectiveness, chemical stability and ease of combining with polymers, which makes them an attractive choice for biomaterial applications ^[126][140]. Further, Cu NPs have antibacterial, osteogenic and angiogenic properties ^[127][141], and have been applied towards the enhancement of both the bioactivity and the antimicrobial properties of Ti dental implants ^[128][142]. More recently, van Hengel et al. incorporated varying amounts of Ag and Cu NPs into TiO₂ coating on additively manufactured Ti–6Al–4V porous implants via plasma electrolytic oxidation ^[129][143]. Further, 75% Ag and 25% Cu caused the eradication of all bacteria in a murine femora model ex vivo, while only Cu NP-modified implants (0% Ag) augmented the metabolic activity of pre-osteoblastic MC3T3-E1 cells in vitro. Alternatively, Ti-6Al-7Nb alloy dental implants were coated with Cu NPs and cultured with P. gingivalis in vitro, and the findings suggested that Cu NPs can aid in local infection control around implants ^[130][144]. In 2020, Xia et al. reported the use of plasma immersion ion implantation and deposition (PIIID) technology to modify Ti implants with C/Cu NPs co-implantation ^[131][145]. The modified implants displayed superior mechanical and corrosion resistance properties and enhanced the antibacterial performance of Ti implants (against S. aureus and E. coli) without causing cytotoxicity (to mouse osteoblast cells) in vitro. In a more dental implant setting, Cu-deposited (micro-/nanoparticles) commercially pure (cp) grade 4 Ti discs (via spark-assisted anodization) were shown to exhibit dose-dependent antibacterial effects against peri-implantitis-associated strain P. gingivalis ^[132][146]. Similarly, micro-arc oxidation Cu NP-doped TiO₂ coatings showed excellent antibacterial activity, while augmenting the proliferation and adhesion of osteoblast and endothelial cells in vitro ^[71][85]. The interaction of Cu NPs with microbes and the bioactivity and toxicity evaluations of Cu NPs can be found elsewhere ^[133][147].

Zirconium (Zr) and zirconia (ZrO₂) are rising as dental implant material choices due to their biocompatibility, corrosion resistance and superior mechanical properties ^[26][42]. It is established that Zr⁴⁺ ions can interact with negatively charged bacterial membranes and cause cell damage and death ^[134][148]. Furthermore, Zr-based implants have been electrochemically anodized in order to fabricate controlled ZrO₂ nanostructures, including nanotubes and nanopores, which can augment implant bioactivity due to their nanoscale roughness ^{[26][135][136]}[42,149,150]. For instance, anodized Zr cylinders were placed in rat femur osteotomy models in vivo, and accelerated bone formation was obtained, in comparison with the controls of unmodified Zr ^[137][151]. Further, Indira et al. reported the dip coating of Zr ions into anodized TNTs to form ZrTiO₄ over the nanotubes, which exhibited enhanced bioactivity (HAp formation in Hank’s solution in vitro) and corrosion resistance ^[138][152]. Similarly, the application of a Zr film on a TiNi alloy via plasma immersion ion implantation and deposition (PIIID) augmented its corrosion resistance ^[139][153]. Nanotube formation has also been extended to TiZr alloys. For instance, Grigorescu et al. used two-step EA to fabricate nanotubes of varied diameters and observed an increase in hydrophilicity with reduction in diameter ^[140][154]. Further, the smallest nanotube diameters exhibited the highest antibacterial effects against E. coli. While ZrO₂/Zr is extensively used as a dental implant material, the leaching of Zr NPs may initiate cytotoxicity. For instance, the application of both Zr and TiO₂ NPs in a dose-dependent fashion could lead to osteoblast morphology changes and apoptosis, affecting both osteoblast differentiation and osteogenesis at high dosages ^[141][155].