The dental implant surface: a review of the past, present and future

Abstract

Purpose

The review provides an in-depth analysis of various factors that affect the long-term success of implants and scrutinizes all available techniques for dental implant modifications, along with their advantages and limitations. Along with established and proposed strategies, newer trends such as responsive coatings, ‘omics’ and AI-based possibilities for translating research into clinical settings are discussed.

Methods

The available scientific literature on dental implants, causes for their failures, and possible surface modification techniques was collected and analyzed. Strategies to prevent implant failures are presented as a comprehensive, structured review.

Results

A literature review of scientific research papers published over the last decade clearly indicates that surface modification of dental implants is critical for ensuring long-term success. Strategies aimed at surface changes consider the intrinsic antibacterial activity, surface texture, and geometry of the implant material. In both healthy and compromised patients, bio-functionalized surfaces can improve osseointegration and reduce peri-implantitis, boosting the success of dental implants.

Conclusions

Dental implants, while promising, face hurdles that hinder their long-term success. Modifying implants through physical, chemical, or mechanical methods could potentially address these challenges. These techniques would require clinical validation before being fully integrated into clinical practice. Moreover, crucial factors such as immune response and in vivo testing are often overlooked.

Introduction

Human teeth play a crucial role in mastication, phonation, and esthetics, and hence their loss poses serious social, psychological, and emotional impacts on the lives of edentulous patients [1]. Although teeth are resistant to abrasion and display lifelong architectural durability, teeth and the oral tissue are highly prone to many diseases, including chemical and microbial attacks and mechanical trauma. Restoring damage caused by such factors is challenging and needs advanced approaches [2]. Dental implants have proven to be a promising and successful substitute for missing or broken teeth. A study conducted in the United States observed a large increase in the demand for dental implants from 0.7% (1999–2000) to 5.7% (2015–2016) [3]. Annually, millions of implants are placed worldwide, with 3 million placed in the USA alone, at an increasing trend of 500,000/year [4]. Based on a report by International Market Analysis Research and Consulting Group (IMARC), in 2022 alone, about 8 lakh dental implant and abutment procedures were carried out in India, which will keep increasing with time [5].

One of the striking breakthroughs occurred in dental implant history when a Swedish orthopedic surgeon, Dr Per-Ingvar Brånemark, during his research on bone healing and regeneration, found that titanium (Ti) effectively integrated into bone without being rejected [6, 7]. The titanium alloy, Ti6Al4V, is currently more popular because of its strength and fatigue resistance compared to pure titanium [8]. Recent research, however, reports toxicity caused by the disintegration of aluminum and vanadium ions from corrosion of the implant and new alloys based on single-phase β titanium, such as Ti12Mo6Zr2Fe, Ti15Mo5Zr3Al, Ti15Sn4Nb2Ta0.2Pd, Ti13Nb13Zr, and Ti29Nb13Ta4.6Zr have been developed. The latter exhibit lower elastic modulus and better biocompatibility than Ti6Al4V [9, 10].

For the long-term success of dental implants, several factors play a key role, including the physicochemical properties of the implant as well as host physiology. Biomechanical factors affecting implant success include the geometry and structure of the material, the quality and quantity of the surrounding bone, and the surgical procedures involved [11]. Titanium and its alloys cannot directly bond with bone, and hence, efforts are made to modify implant surfaces to make them more osteoconductive for successful anchorage. Osseointegration involves the direct connection of the bone and the implant. The stability and long-term success of the implant depend on osseointegration [12]. The early stage is characterized by an initial healing phase, inflammation, the release of growth factors and cytokines, which contribute to healing. Subsequently, cell proliferation occurs, in which various biomolecules aid the differentiation of mesenchymal cells into osteoblasts. The osteoblasts produce a collagen-rich matrix, which is eventually mineralized to form bone anchoring the implant. This process may take 2 weeks to 4 months, depending on the level of osteoconductivity of the implant surface [13]. The long-term success and functional stability of titanium dental implants are critically dependent on the biomechanical outcomes (i.e., the biting force distribution, optimal micromovement at the bone-implant interface, bone remodeling, and growth of tissue on the implant) achieved through osseointegration.

Along with osseointegration, another important factor determining implant success is antibacterial activity. In natural teeth, Sharpey’s fibers are found to be perpendicular to the cementum and periosteum of the alveolar bone, which is responsible for firmly holding the teeth. Unfortunately, in the case of an implant, a deep peri-implant crevice is formed due to the parallel arrangement of fibers, which is more favorable for bacterial penetration [14]. This leads to increased bacterial contact with the implant surface, consequently allowing the formation of complex biofilms. Gram-positive aerobic bacteria, including species of Streptococcus and Actinomyces, are the first ones to initiate biofilm formation. Suitable growth conditions created by these bacteria enable the growth of anaerobic late colonizers, such as red complex species (Tannerella forsythia, Porphyromonas gingivalis, Treponema denticola) and orange complex species (Prevotella intermedia, Fusobacterium nucleatum), to participate in the biofilm [15]. In response to this bacterial infection, immune system cells migrate to the site of infection and initiate an immune response, leading to inflammation and loss of the bone surrounding the implant (peri-implantitis). A cross-sectional study carried out by Mombelli et al. showed that the prevalence of peri-implantitis seems to be in the order of 10% implants and 20% patients during 5–10 years after implant placement [16]. Peri-implant mucositis is another condition which is characterized by soft tissue inflammation around the implant without bone loss. Peri-implant mucositis can lead to peri-implantitis if it is not managed in its early stages by maintaining oral hygiene. At the transmucosal level, the interface between the implant and peri-implant soft tissue plays a crucial role ensuring soft tissue integration of the implant. This peri-implant soft tissue creates a seal between bone and the oral environment, which inhibits the entry of oral pathogens near the implant surface [17]. Current research focuses on imparting antimicrobial activity and enhancing the osseointegrative properties of the implant to prevent microbial contamination and promote early healing. Desirable characteristics of dental implant material are represented in Fig. 1.

Fig. 1.

Desirable characteristics of dental implant materials

Implant design: materials and geometry

Designing an implant with all the desired characteristics, as mentioned in Fig. 1, is a challenge. The physicochemical properties of implant surface materials influence various reactions, including angiogenesis, osteogenesis, and cell adhesion, migration, proliferation, and differentiation in the surrounding tissues.

Therefore, the implant material and its characteristics become a crucial factor in determining the success of implants. Important parameters that should be considered for long-term implant success include its chemical composition, hydrophilicity, corrosion resistance, elasticity, tensile strength, roughness, biocompatibility, cytotoxicity, and implant geometry, which will help in successful osseointegration and soft tissue integration.

Based on Dr Per-Ingvar Brånemark’s research, titanium has been considered the ‘Gold standard’ for dental implants and is also used in hip, joint and spine fixation implants. Before considering Ti as the gold standard, various other metals were tested. Stainless steel is ductile and strong, but it can stimulate bacterial infections, leading to failure. Similarly, Cobalt-based alloys are also a good candidate but are less ductile and preferred for hip and joint implants, which require high wear resistance [18]. The biocompatibility of titanium is attributed to the formation of a self-limiting and stable oxide layer, which prevents further oxidation and corrosion which leads to mechanical stability of the implant surface and surrounding tissue. Additionally, Ti is strong, lightweight, has a low modulus of elasticity (110 GPa), and is less rigid, making it an ideal contender for dental implants. Despite the favorable physicoochemical characteristics, Ti implants face several obstacles in implantology, which include inconsistency in osseointegration, lack of inherent antibacterial activity to combat biofilm formation, triggering an immune response due to leached particles, and the appearance of bluish-gray color, which compromises the esthetics [19]. Additionally, Titanium requires a stress-free healing period of at least 3–4 months, which increases the risk of infection and can lead to implant failure. Zirconia (Zirconium dioxide, ZrO₂), a crystalline ceramic material, has gained popularity because of its high esthetic value (it is opaque white), subpar toughness, biocompatibility, and chemical stability. It exhibits reduced bacterial adhesion and provides corrosion resistance. It is the best option available to people with metal allergies or who simply want an esthetic appeal. Unfortunately, Zirconia is prone to phase transformation and low temperature degradation (LTD), which leads to microcracking [20] however, advances in surface modifications, cementing techniques, use of monolithic implants and use of newer dual-cure or self-adhesive resin can be useful to address this issue [21–23] Along with Zirconia, PEEK (Polyether ether ketone) is currently considered an option against Ti-based dental implants. PEEK is a synthetic polymer with natural tooth color and excellent chemical, mechanical, and biocompatibility properties. Like most polymers, PEEK is also hydrophobic and hence has very low osteoconductive properties, which is a major issue. Another major issue observed in PEEK-based implants is the formation of bacterial biofilms, which might induce peri-implantitis. Considering these aspects, PEEK is very far from being an alternative to Ti-based implants. Bioceramics such as Calcium Phosphate (CaPO₄), Tricalcium Phosphate [Ca₃(PO₄)2], or hydroxyapatite-coated materials are relatively weak and structurally unstable, impeding their utility in load-bearing applications.

In clinical and experimental settings, Ti and its alloys are still considered the gold standard of dental implants amongst options like Zirconia and PEEK. Well-defined international mechanical standards regulate the use of titanium as a gold standard for dental implants. The ISO/TC 106/SC 8 committee developed the ISO 5832 series, which defines the minimum requirements for implant-grade titanium and its alloys. These include a minimum yield strength of 795 MPa and a tensile strength of 860 MPa for Ti-6Al-4V, as well as limits for elongation and chemical composition. These values ensure that titanium implants consistently achieve high strength, ductility, and reliability necessary for long-term clinical function [24]. Standardized fatigue testing under worst-case loading conditions for long-term performance validation is specified in ISO 14801[25]. Together, these standards provide a strict regulatory framework that ensures predictable mechanical behavior. Although PEEK for surgical applications is standardized under ASTM F2026 (material specification) and related test methods (e.g., ISO 15309), standardization addressing full implant mechanical performance and fatigue testing is not as fully developed as the ISO standards for Ti and its alloys [26].

A comparison between dental implant materials is mentioned in Table 1.

Table 1.

Comparison between Ti, Zr and PEEK as dental implant materials [114–116]

Characteristics	Titanium	Zirconia	PEEK
Material	Titanium alloy, usually Ti6Al4V	Zirconium dioxide (Bioceramic)	Biocompatible polymer
Color	Metallic gray	Similar to natural teeth	Beige/cream/off-white
Esthetics	Less esthetically pleasing	Mimics natural teeth, no gray shadow casting	Better than Ti but not as good as Zr
Strength	High, Strong and durable	High, but brittle than Ti	Lowest of the three
Weight	Heaviest of the three	50% lighter than Ti	Lightest of the three
Availability	Widely available and practised	Widely available but not practised by all dentists	Less widely available
Surgical technique	Well-established	It may require different protocols for drilling	May require specific techniques
Biocompatibility	Excellent, well-established research	Excellent	Good
Osseointegration	Excellent, extensively documented results	Good but slower as compared to Ti, surface modifications may enhance the process	It has potential, but still limited data
Plaque formation	Susceptible	May resist plaque buildup better than titanium due to a smoother surface	Susceptible
Gum tissue response/inflammation	Observed in some patients	Less likely	Limited data for long-term
Metal allergies	Should be avoided in patients with Ti allergies	Suitable for patients with metal allergies	Suitable for patients with metal allergies
Cost	Generally, less expensive and affordable	Generally, more expensive than titanium	Comparable, maybe cheaper than Ti
Long-term success rates	Considered the Gold Standard because of its high success rate with proper care	Similar to titanium implants, with proper care (studies suggest comparable survival rates)	Limited data for long-term
Fracture risk	Very low	Less but higher risk than Ti	Higher risk
Ductility	High	Low, brittle material	High
Young’s Modulus	100–115 GPa	200–230 GPa	3–4 GPa
Comparison with Cortical Bone	6.7–7.7 times stiffer than cortical bone (high-stress shielding)	2–2.3 times stiffer than cortical bone (Low-stress shielding)	0.2–0.27 times stiffer than cortical bone (No or very little stress shielding)

Macro-level modifications, which can alter the implant geometry on a millimeter scale, can help alter Ti implant geometry, consequently improving primary implant fixation and stability. Such macro-level design, considering irregularities like thread pattern, shape (parallel/tapered), length, and diameter, is important in determining surface area, mechanical bone implant interlocking, bone ingrowth, and 1° stability [27]. Parallel implants in which the diameter is consistent throughout the length were traditionally used until technique-based sensitivity was observed. In the 1990’s Jack Hahn developed tapered implants which decreased in diameter along the length like real tooth roots. Tapered implants distributed occlusal forces towards neighboring bones and required less bone volume to be removed, making them superior for immediate placement [28]. In tapered implants, the overall diameter increased, which improved overall force distribution, elevated capacity for load bearing, and reduced stress throughout the implant length [29]. Dental implant threading improves 1° contact, stability, and stress distribution in the jawbone. Face angle, thread pitch, shape, and depth are the most important features of threading. The face angle is the angle between the face of a thread and the plane perpendicular to the implant; thus, lowering the face angle and reducing the shear force are vital for bone response [30]. The distance between two neighboring threads, known as pitch, must be low to contribute less stress and more bone-to-implant contact. Additionally, increasing the thread depth and width enlarges the functional surface and improves primary stability; however, it becomes more challenging to insert the implant. Overall, the physical geometry of the implant plays a crucial role in osseointegration and combating biofilm formation, which ultimately leads to long-term implant success.

Reasons for implant failure

Owing to their excellent physicochemical properties, such as superior mechanical strength, lower elastic modulus, high corrosion resistance, excellent biocompatibility, and immense research on implant geometry and shape, the success rate of implant surgery has reached 90–95% [31, 32]. The long-lasting survival rates of titanium and its alloys have made titanium the “gold standard” material for the manufacture of endosseous dental implants. The 5% failure rate can be attributed to multiple factors, including host-related factors, material characteristics of the implant, iatrogenic factors, as well as bacterial infections [33, 34]. It was found that bacterial infection is an important factor responsible for peri-implantitis and failure of dental implants [35, 36]. Despite excellent physicochemical properties, conventional titanium implants require at least three months of stress-free healing period, which may increase the risk of infection and may lead to implant failure [37]. Additionally, patients who require early restoration of their esthetic appearance will not accept such a lengthy healing period.

Host-related factors such as systemic conditions, cigarette smoking, and genetic factors have been reported to influence the incidence of peri-implantitis. Wagner et al. [38] report that patients with poorly controlled diabetes are more prone to peri-implantitis, especially during the post-implantation healing period. Due to impaired cellular and vascular responses in hyperglycemic patients, the healing process is delayed, and the tissue destruction is enhanced. Osteoporosis is a systemic bone disease that results in fragile bones, ultimately making them vulnerable to fractures, which retards the bone healing process in implant patients [39]. However, based on the observations by Giro et al., definitive conclusions regarding the impact of osteoporosis on dental implant therapy cannot be made in the absence of a randomized clinical trial evaluating the effect of osteoporosis in dental implant therapy [40]. Along with systemic diseases, cigarette smoking is also considered a risk factor for peri-implantitis. Beneficial commensals in the subgingival tissue are reduced in number due to smoking, allowing pathogenic bacteria to thrive [41]. A cross-sectional study found a significant rise in peri-implantitis (30.5%) among smokers compared to non-smokers (18.2%) [42]. Additionally, dysregulation in the secretion of pro-inflammatory factors due to a fault in genetic makeup results in excessive bone loss, thus increasing the risk of implant failure. Implant failures may also be observed in individuals receiving antiresorptive therapeutics (viz., bisphosphonates (BP) and specific monoclonal antibodies such as Denosumab), which reduce bone turnover. High doses of these compounds may lead to early onset of medication-induced osteonecrosis of the jaw (MRONJ), compromising osseointegration [43]. However, the impact of antiresorptive therapy on implant success remains debatable [43, 44]. Popular implant materials such as titanium can trigger immune responses in certain cases, causing local inflammation and hypersensitivity. These local reactions can be addressed by replacing Ti with zirconia, but the latter has also been reported to cause allergenic responses [45]. Besides implant materials, specific foods, drugs, etc., may induce allergic responses resulting in implant failure. However, such a dataset is extremely limited [46]. In addition to host factors, several implant-based mechanical factors significantly influence the long-term success of dental implants.

The long-term success of the implants is significantly influenced by the design of the prosthetic restoration. The control of plaque accumulation is considerably hampered by poorly designed implants, which raises the possibility of developing peri-implant disease. Compared to single crowns, full-mouth implant-supported prostheses have been linked to a 16-fold increase in peri-implantitis due to insufficient accessibility for plaque control [47]. Cemented implant restorations are 3.6 times more likely to develop peri-implantitis than screw-retained prostheses [48]. It has been reported that rough implants, having an average surface of more than 0.8 μm favor 25× more bacterial attachment than implants with 0.3 μm surface roughness [49]. Bacterial infection is a leading cause of implant failure. Bacterial adhesion on the implant surface triggers biofilm formation, which progresses through several stages, starting with the attachment of planktonic bacteria to a solid support, coaggregation with other bacteria, growth of bacteria within the biofilm until the maturation phase, and finally, detachment and migration of a part of the biofilm to another site [50]. The reasons for implant failures are summarized in Fig. 2. It is imperative to maintain proper oral hygiene to avoid peri-implantitis and ensure implant success. Obtaining an implant with all desirable characteristics and antibacterial activity is challenging, and therefore, various surface modification techniques can be explored. The surface modification allows us to modify an implant in such a way that it will possess all the desirable biomechanical characteristics, as well as promote osseointegration and display antibacterial activity.

Fig. 2.

Reasons for dental implant failure

In this review, we have summarized the recent surface modification techniques used for improving the antibacterial activity and osseointegration of titanium dental implants.

Surface modifications

Ti materials show high corrosion resistance and optimal reactivity with the host tissues and are therefore considered bioinert [51]. However, bacterial contamination at the site can delay the osseointegration, ultimately causing implant failure, posing a heavy economic, physical, and social burden on the patient [4, 52]. This issue can be resolved by surface modifications using several strategies to achieve reduced infections, enhanced bone formation, healing acceleration, and improving osseointegration in areas with defective bone quantity/quality.

Surface characteristics are one of the important parameters that decide the success of an implant. Modification of implant surfaces leads to an increase in biocompatibility, acceleration of osteointegration, and shortening of the treatment period. Surface modification techniques allow us to alter the morphology and surface chemistry of the implant without its bulk mechanical properties being compromised [53]. Surface modifications of the implant body are aimed at increasing implant surface roughness to maximize osseointegration and bone-to-implant contact (BIC), which together determine the long-term success of the implant. The abutment surface remains in contact with the transmucosal soft tissue and is partially exposed to the oral cavity, which includes specific bacterial communities. Thus, coatings that achieve soft tissue integration while preventing bacterial colonization are designed for abutments [54]. With advancements in research, the titanium implant surface can be modified while maintaining its chemical composition, or it can be coated with appropriate materials, thereby reducing the reactivity of native titanium with host tissue. Surface modification methods include mechanical, chemical, and physical techniques, which allow an increase in reactive surface area by either additive or subtractive methods. There are 3 important interfaces between dental implants and host systems, which are potential targets for modification as they play a vital role in implant success. These interfaces are represented in Fig. 3.

Fig. 3.

Interfaces between the implant and oral tissues: sites for modifications

Two types of surface modifications are described below in detail viz., surface modifications for enhanced osseointegration and surface modifications providing antibacterial activity.

Surface modifications for enhanced osseointegration

To enhance osseointegration activity, the various surface modifications can primarily be divided into native and biofunctional. In native surface modifications, the surface of the implant itself is modified using either mechanical, chemical, or physical techniques. Whereas, in biofunctional modifications, the implant surface is coated with either organic or inorganic molecules, which allow and enhance osseointegration. Various techniques are summarized in Fig. 4.

Fig. 4.

Surface modifications for enhanced osseointegration

Native surface modifications

The bone response to the implant is largely governed by the surface topography of the implant, and healing at the bone-implant interface, which occurs through two mechanisms: distance and contact osteogenesis. In distance osteogenesis, new bone growth starts on existing bone and is directed towards the implant surface. In contrast, in contact osteogenesis de novo bone formation occurs on the implant surface, and the direction of growth is towards the bone. Conventional implants support only distance osteogenesis, which takes more time to heal and implant loading, thereby increasing the patient’s edentulous period [55]. Persistent efforts must be made to make implant surfaces more osteoconductive. Although implant coatings with various organic and inorganic substances have shown very good osseointegration properties, peeling of the coating may lead to implant loosening and operational failure [56].

Surface micro-nano structuring is an alternative surface modification strategy that has been shown to improve the osteogenic activity of the implant. The rough and porous structure increases the effective contact area between the osteoblasts and the material surface, which encourages their spreading and proliferation [57]. Nanoscale surfaces possessing high surface energy allow faster protein adsorption, which is crucial for the initial cellular interaction with the implant surface [58]. Titanium surface topography modification can be done by various techniques, including physical, chemical, and mechanical [13]. These methods can be used alone or in combination, e.g., sandblasting and large grit acid etching (SLA).

All these methods, along with their limitations and results, are summarized in Table 2.

Table 2.

Techniques for modification of native implant surface

	Technique	Methodology	Limitations	Findings and references
Mechanical	Sandblasting	Hard ceramic particles bombarded on surface using a pressurized air stream	Release blasted particles in tissues, leading to toxicity	1) 96% osseointegration [117] 2) Accelerated bone tissue growth, reduced pathogen attachment [118] 3) 97% implant survival after 10 years [119]
	Shot peening and Laser peening	Surface repeatedly shocked or hit with a laser	Cannot modify at subsurface levels	1) Improved wettability and cell adhesion [120] 2) Enhanced fatigue strength, hardness, wear and tear resistance [121]
Chemical	Acid etching	Submerge implant in corrosive acids (Nitric acid, sulfuric acid, hydrofluoric acid)	Undesired chemical reactions at implant surface	1) Form micro pits, enhance osseointegration, bone formation and osteoconductivity [122] 2) Ti NP formation on the surface leading to protein adhesion [123]
	Anodization	Electrochemical deposition of oxide layers	Varied color formations, high initial costs, overall inconsistencies	1) Enhanced adhesion, and differentiation of HGF and HBM mesenchymal cells [60] 2) 40% increase in mineralized bone formation [124] 3) Antibacterial activity against B. cereus [125]
	Hydrothermal treatment	Use hot-compressed water as a reaction system between the substrate and test material	Very high temperatures can cause phase changes in Ti implants	1) TiN coating caused increased fibroblast cell adhesion, cell proliferation, wettability, abrasion resistance [126] 2) Distilled H2O- improved cell integration, soft tissue development [127] 3) NaOH- Less thrombosis, lower fibrinogen adsorption, good blood compatibility [128]
	Chemical vapor deposition (CVD)	Deposition of film on solid substrate by gas phase chemical reactions	Limited adhesion, limited applicable chemical reactions	1) Increase osseointegration, bone mineralization in cortical, cancellous bone [129] 2) Increased osteogenic differentiation and calcium deposition [130]
	Micro Arc Oxidation (MAO)	Electrochemical deposition of thick oxide layers using high applied potential	Residual stress on deposited film, variable parameters change coating characteristics	1) Ag incorporated MAO reduced S. aureus and E. coli attachment [131] 2) Increased proliferation of human bone marrow-derived mesenchymal cells [132]
	Fluoride treatment	Deposit Fluoride ions on implant surface	Concentration higher than 1 mM is cytotoxic	1) Increased expression of BMP2, sialoprotein, osteoblast differentiation [133] 2) 100% survival rate after 10 years in mandibular implants [134]
	Sandblasted, Large grit, Acid etched (SLA)	Large grit particles (∼250–500 μm) are sandblasted on the surface followed by acid etching	Heterogenous nano surface, Limited oxide layer thickness	1) 92.9% success rate after 17 years, increased osseointegration [9] 2) Ti-NW-Zn upregulated osteogenic differentiation, adhesion, proliferation [135] 3) Antibacterial activity against S. aureus and improved osteogenic activity of MC3T3-E1 cells[136]
Physical	Plasma spraying	Powder sprayed on implant surface using Ar, He, N2 etc as a plasma source	Bonding issues, formation of cracks in coatings, copper contamination from sprayer nozzles	1) Increased MG-63 proliferation, adhesion after decontamination, decreased mitochondrial damage after using Ar plasma[137] 2) Atmospheric plasma spraying to fabricate Zn-Sr-Mg-HAp doped Ti which inhibited P. gingivalis, P. nigrescens[138]
	Laser ablation	Use different types of lasers to create exact, complicated, high-resolution surface patterns	High power, high duration exposure might cause deleterious effects on the implant	1) Increased BIC after formation of micro pits[139] 2) 30% reduction in E. coli adherence after using a nanosecond pulsed laser[140]
	Physical vapor deposition (PVD)	Coat substances using the vapor condensation phenomenon	High Process cost, low output, complex process	1) Enhanced abrasion, corrosion resistance, osteodifferentiation of Human bone marrow cells[141] 2) Maintain aseptic conditions around implant, promote osseointegration and antibacterial activity against peri-implantitis pathogens[142]
	UV-PF	Treatment with UV-A and UV-C light	Exact parameters like irradiation time, irradiation distance still poorly researched	1) Enhanced osteoblast cell adherence and activity[143] 2) Possible Ti ageing reversal agent[144] 3) 80% strength increase in implant integration, cell attachment, proliferation[145] 4) Better soft tissue seal formation after 12 min exposure to 250 nm and 360 nm UV light mixture[146]

Biofunctional modifications

As stated earlier, native Ti possesses excellent properties such as biocompatibility, high corrosion resistance, and a high strength-to-weight ratio, but the success of an implant depends on its integration with the surrounding tissue, the quality of surrounding bone tissue, and the presence of infection. Therefore, the surface of the implant is the most crucial factor that dictates the long-term implant success. Hence, to enhance the bioactivity and augment the osseointegration potential, surface engineering of the implant surface through functional coatings (with bioactive molecules, including organic and inorganic molecules) or micro/nanostructuring has been proposed.

Organic coatings

Growth factor-based coatings: Bone morphogenetic proteins (BMPs) belong to the transforming growth factor-β (TGF-β) superfamily, triggering a cascade of events that leads to bone development. BMP-2 is involved in the osteogenic differentiation of mesenchymal stem cells and is also involved in several other regeneration processes. BMP-2, BMP-7 and alendronic acid immobilized separately on plain Ti surfaces showed significantly improved adhesion, proliferation and differentiation of bone marrow-derived mesenchymal stem cells compared to plain and SLA surfaces [59]. Recently, recombinant human BMP-2 immobilized on poly glycidyl methacrylate (pGMA) coated Ti efficiently differentiated human adipose-derived stem cells (hASC) into osteoblasts. Cells grown on rhBMP-2 (Recombinant human bone morphogenetic protein) coated surface expressed a higher level of OCN, RunX2, and COL1 mRNA, which is characteristic of osteoblast cells; in addition, ALP activity and calcium deposition were much higher on the coated surface [60]. TiO₂ nanotubes can act as excellent reservoirs of drugs, and controlled release of hBMP-2 is essential to avoid adverse effects such as bone overgrowth, ectopic bone formation, and immune response. Considering these facts, Li et al. immobilized hBMP-2 on Ti nanotubes (TNT) and found 1.5-, 1.7- and 1.9-fold increase in adherent cell number of MC3T3-E1 cells on the TNT, hBMP-2/Ti, and hBMP-2/TNT surfaces, respectively, compared to the control (polished Ti surface). Also, cells on hBMP-2/TNT displayed more cell proliferation, ALP activity, osteogenic marker gene expression, spreading, and improved vinculin expression, indicating superior cell adhesion. In vivo studies in rats revealed better bone formation and enhanced bonding strength at hBMP-2/TNT and tissue interface [61].

The BMP-2/GDF-5 dual coating supported remarkable bone formation at the implant and host bone interface with significantly increased bone volume (BV), Bone volume to total volume ratio (BV/TV), BIC, and removal torque values [62] The same research group in another study showed that Ti surface modified by hydroxyapatite (HA) and Heparin-BMP-2 complex supports higher proliferation, ALP activity, calcium deposition, and OCN gene expression by the MG-63 human osteosarcoma cells. Results of an in vivo study after 4 weeks of implantation in New Zealand white rabbits revealed that the Ti/HAp/Hep/BMP-2 sample resulted in higher bone formation around the implant. In histology evaluation, large-scale osteons with the highest number of osteocytes were formed at the Ti/HAp/Hep/BMP-2 and tissue interface [63].

Extracellular matrix (ECM) based coatings: Implant surface tethering with short synthetic peptides derived from the bone extracellular matrix proteins has been used to promote the attachment and differentiation of osteoprogenitor cells. Ti surface modification using components of extracellular matrix (ECM) such as collagen, fibronectin, vitronectin, and bone sialoprotein (BSP) has also been studied as an effective way of promoting implant stability. Coating implant surfaces with ECM proteins modifies the initial immune response, and they provide cell adhesion motifs, which act as scaffolds for bone cells and enhance osseous anchorage and reduction in implant micro-movements. Type 1 collagen is the main component of bone ECM and regulates various aspects of cell behavior, such as migration and attachment, through interaction with surface integrin receptors on the osteoblast cells [64].

Fibronectin and vitronectin are the glycoproteins present in the ECM and are known for supporting initial cell attachment; α5β1 and αVβ1 integrin receptors on the osteoblasts recognize the ligand present in fibronectin and vitronectin, respectively [65]. Fibronectin grafting to glow discharge plasma (GDP) pretreated Ti benefited the cell adhesion, migration, and proliferation of MG-63 cells compared to the control Ti surface [66]. Recently, vitronectin (ECM protein) derived peptide motif (VnP-16) physically adsorbed on the SLA Ti surface showed strong cell attachment and spreading activity by human osteoblast-like cells (HOS and MG-63). In vivo study showed no inflammatory or immune response toward the peptide-coated implant, and sufficient mineralization was observed in all the groups [67]. Human osteoblasts possess a net negative charge, which would favor electrostatic attraction to the positively charged material surfaces. Based on this assumption, Nebe et al. [68] developed a positively charged plasma polymerized allylamine (PPAAm) nanolayer on the Ti and Y-TZP ceramic surface and demonstrated significant spreading of MG-63 cells compared to an uncoated surface. In a recent study, the effects of type 1 collagen-coated SLA Ti implant on bone healing and bone formation in rabbits showed improved osseointegration. The collagen-coated surface was surrounded by greater bone formation inside the screw thread (bone area inner thread, BAIT) and in contact with the screw surface (bone area outer thread, BAOT) with no evidence of soft tissues in between the implant and bone [69]. Surface modification with P-15 peptide that mimics the cell-binding domain of human type 1 collagen resulted in improved osteoblast adhesion and calcium deposition [70]. An ECM mimetic, bio-composite osteogenic nano-fibers consisting of polycaprolactone, gelatin, hydroxyapatite, dexamethasone, β-glycerophosphate, and ascorbic acid electrospun on Ti implant placed in rabbit tibia showed significantly greater pull-out test compared to control implants. Additionally, extensive mature bone formation with higher BIC % around the coated implant without graft rejection or inflammatory reaction signs indicated the osseointegration power [71].

The binding of integrin receptors on the osteoblast membrane with the RGD sequence of extracellular matrix proteins (ECM) such as collagen Type 1, fibronectin, and vitronectin induces intracellular signaling pathways regulating osteoblast adhesion, survival, proliferation, and differentiation [37, 70]. Hoyos et al. [72] reported that RGD and KRSR (heparin-binding peptide) dual bio-molecular interface mimicking the natural ECM microenvironment improved the number, area, proliferation, and mineralization of SaOS2 osteoblast cells compared to control Ti. RGD and KRSR synergistically induced the production of calcified matrix by the osteoblast. Mussel adhesive protein (MAP) coated Ti surface showed superior rBMMSC cell adhesion and spreading, calcium deposition, and expression of osteogenic differentiation-related genes Runx2 and ALP than the uncoated Ti. In vivo studies showed more bone formation around the MAP-coated implant with higher Bone volume/Total volume (BV/TV), trabecular number (Tb. N), and trabecular thickness (Tb. Th) values. After 8 weeks of surgery, both BA and BIC were significantly higher on the MAP surface than on the uncoated one [73].

RNAi can also be used for enhanced osseointegration besides ECM and growth factor-based proteins. The RNAi (siRNA, miRNA mimics, or anti-miRNA) have been identified as key regulators to promote the osteogenic potential of MSCs [74]. It has been found that several miRNAs are involved in the differentiation of mesenchymal precursors into osteoblast cells through the regulation of transcription factors [75] Overexpression of miR-29c-3p in BMSCs showed enhanced osteoblastic differentiation and osseointegration by directly suppressing Dvl2 and Fzd4 in the Wnt/ βcatenin signaling pathway [76]. Knowing the importance of miRNAs in osteoblast differentiation and osteogenesis, various researchers have modified the Ti surface with miRNAs and have studied their effects on implant osseointegration. Microporous Ti surface functionalized with anti-miR-138 and miR-29b led to the enhanced osteogenic activity of MSCs due to upregulation of osteogenic gene expression, ALP activity, collagen production, and ECM mineralization [77]. PEG/PEI dual-modified graphene oxide condensed with siRNA targeting Ckip-1, (Casein kinase-2 interacting protein-1, which is a negative regulator of bone formation) deposited on titania nanotubes induced in vitro osteogenic differentiation and significantly higher bone formation around the coated implant in vivo [78]. In another study, Ti functionalized with Chitosan/siRNA complex targeting long noncoding RNA MIR31HG showed excellent knockdown efficacy and enhanced osteoblastic differentiation of BMSCs and ectopic bone formation without any in vivo cytotoxicity [79].

Inorganic coatings

An alternative strategy to alter surface chemistry to encourage bone cell attachment and differentiation is surface functionalization with inorganic elements involved in the osseointegration processes, such as calcium, phosphate, fluoride, magnesium, strontium, and zinc, which are the basic elements of bone tissue [80]. Ca alone or in combination with other elements has been extensively applied for implant surface modification [81, 82]. Some of the techniques are summarized below.

Hydroxyapatite (HA, Ca₁₀(PO₄)₆(OH)₂), owing to its outstanding properties such as chemical similarity with bone, osteoconductivity, and lack of immune response, has attracted attention as a surface-coating compound. Hydroxyapatite (HA), a bioceramic, is one of the most common molecules used because of its beneficial bioactivity and biocompatibility. HA is a highly osteotropic bioceramic that helps in the strong chemical bonding of bone with the implant by creating a calcium phosphate precipitate on the Ti surface. HA can be coated on the implant surface by plasma spraying, vacuum deposition, sol-gel dip coating method, or electrolytic process [83, 84]. HA can also be used to modify Zirconia implants [85] The HA coatings on MAO and Ti nanotube surfaces have shown promising osteogenic activity, such as enhanced attachment, proliferation, and differentiation of osteogenic cells without any toxic effects [86].

Graphene-based bio-mimetic coatings are a new approach for dental implant modification. Graphene significantly reduced wear (50%) and brittleness, which leads to enhanced tribological behavior [87]. The hexagonal, 2D structure of graphene resulted in characteristic mechanical, optical, and electronic properties, which lead to resistance to degradation and provide stability to Ti₆Al₄V implants [88]. Zeng et al. [89] fabricated a Graphene Oxide (GO)/Hydroxyapatite (HA) composite coating using an electrochemical deposition technique. This composite coating resulted in increased crystallinity of deposited apatite particles and increased the bonding strength of the coating. Dubey et al. [90] coated the Ti substrate with graphene using a dry transfer technique based on the hot-pressing method. This coated Ti surface was cytocompatible, non-toxic, induced human osteoblast maturation and also inhibited biofilms of Streptococcus mutans and Enterococcus faecalis. Carbon nanotubes are currently gaining momentum as a medical biomaterial owing to their biomechanical, electrical, and chemical properties. Such coatings were reported to promote surface roughness, cell proliferation, and adhesion [91]. Adding C nanotubes to polymethyl methacrylate resins prevented the adhesion of Staphylococcus aureus, Streptococcus mutans, and Candida albicans on the Ti surface. Such coatings were non-toxic to oral keratinocytes [92]. Plasma-sprayed CaO-MgO-SiO2-based glass-ceramic-coated Ti implant after 3 months of implantation in the rabbit femoral defect model showed thicker, direct, active bone formation than the HA-coated implant without any sign of inflammation and formation of fibrous connective tissues [86]. The in vitro study on the comparative osteogenic capacity of divalent cations Ca and Mg coatings on the Ti surface using MSCs showed that both coatings were able to enhance the early spreading and differentiation of MSCs, but the Mg-modified surface exhibited a more potent capacity to promote early osteogenic differentiation than the Ca coating [93].

Surface modifications against bacterial infections

Modifications of Ti with several synthetic and natural agents viz. antibiotics, metals, and peptides are illustrated in Fig. 5 and briefly described in the following sections.

Fig. 5.

Surface modifications against infections

Antibacterial coatings

Surface modification by coating the implant surface with antibacterial compounds is the best strategy to inhibit initial bacterial cell attachment and subsequent biofilm formation, responsible for peri-implantitis. Several compounds have been tested for this application, such as metals, polymers, antibiotics, antimicrobial peptides, composite nanomaterials, etc.

Antibiotics are the most important medication used for bacterial disease management and are widely used by dentists for the treatment and prevention of odontogenic infections. Penicillin, amoxicillin, metronidazole, clindamycin, azithromycin, ciprofloxacin, gentamicin, etc. [94, 95]. However, overuse or misuse of broad-spectrum systemic antibiotics is associated with severe complications to human health, such as gastrointestinal disturbance, fatal anaphylactic shock, and the emergence of devastating antibacterial resistance [96]. A prophylactic approach for treating peri-implantitis is antimicrobial implant coating, which ensures sustained local delivery of antibiotics and killing of microbes without any systemic toxic effects on the human body [35].

Various coating methods, such as silanization, sol-gel silica coating, TiO₂ nanotubes, PLGA, and HA have been used to couple the antibiotics to the biomaterial surfaces [97]. Although many antibacterial coatings are investigated, very few have been developed and adopted in commercial and practical settings [98].

Antimicrobial peptides (AMPs) are naturally occurring polypeptide sequences comprised of cationic and hydrophobic amino acids (~12–20 residues) with direct antibacterial activity [99] AMPs are produced by all eukaryotic and prokaryotic life forms as a part of their innate immunity [100]. The underlying mechanisms of the antibacterial action of AMP include membrane disruption, metabolic interference, and targeting cytoplasmic components (nucleic acids and proteins) [101, 102]. Recently, AMPs alone or in combination with other molecules have been used as a bio-functional coating on titanium implants to improve their antibacterial and osteogenic potential. Apart from implant coatings, they have also been used as an additive in the guided bone regeneration (GBR) membrane and hydrogels used in dental implant infection treatments [103, 104]. Even with all the great advancements and promising results offered by antimicrobial peptides, there is still limited information about how they work, thus limiting their scope of applications. A few references investigating antibiotics and antimicrobial peptides are summarized in Table 3.

Table 3.

Surface modification strategies against bacterial infections

Agent and process used	Findings	Reference
Antibiotics
PLGA loaded with Norfloxacin using spray drying technique	1) Biphasic release with initial burst for 48 H and slow release for 15 days 2) 99% decrease in E. coli viability 3) No cytotoxicity against mouse L929 cells	[147]
PLGA loaded with Amoxicillin using a simple dipping method	1) Increased drug release with time, inhibiting S. aureus and S. epidermidis. 2) Negative effect on viability, cell adhesion, spreading ofMG-63 cells	[148]
Silk fibroin nanoparticles loaded with Gentamicin	1) Reduced adhesion of S. aureus 2) Enhanced adhesion of MG-63 cells without cytotoxicity. 3) Increased Calcium deposition and ALP secretion	[149]
Doxycycline loaded in gelatin solution using sonication. Locally delivered using TiO2 nanotubes	1) Suppressed growth of P. gingivalis 2) 1.29 µg/mL (mean) drug released for 28 days.	[150]
Polymeric nanoparticles doped with doxycycline/doxymethasone	1) No significant increase in osteoblast proliferation (MG-63) 2) Upregulation of osteogenic proliferation genes (TGF-β1, TGF-βR1 and TGF-βR2)	[151]
Gentamicin incorporated in polyethylene glycol and ethylene glycol	1) Complete inhibition of E. coli, S. aureus 2) No bacterial adhesion indicating antifouling activity	[152]
Tetracycline incorporated in polymer nanofibers consisting of PLA/PCA and gelatin using an electrospinning technique	1) Inhibited growth of A. actinomycetemcomitans, F. nucleatum, P. gingivalis and P. intermedia 2) Increased osteogenesis indicated by enhanced proliferation and ALP activity of MC3T3-E1	[153]
Minocycline hydrochloride loaded in Graphene oxide	1) 85.3, 91.5, and 100% antibacterial rate against S. aureus, E. coli, and S. mutans respectively 2) improvement in cell adhesion, spreading and proliferation of HGF cells	[154]
Cefotaxime sodium (CS) immobilized on dopamine	1) Inhibited adhesion, proliferation and biofilm formation of E. coli and S. mutans 2) Enhanced MG-63 cell attachment and growth	[155]
Antimicrobial peptides
Cationic polypeptide immobilized on TiO2 nano spikes	1) 99.9% killing ratio against S. aureus and E. coli 2) log 6.65 reduction in bacterial attachment in subcutaneous rat model	[156]
Hexapeptide consisting of two fluorinated phenylalanine (4F-Phe) residues, an RGD sequence and 3,4-dihydroxyphenylalanine (DOPA) synthesized by solid phase peptide synthesis	1) 25% and 50% reduction in the attachment of early (V. parvula and S. sorbents) and late (P. gingivalis) colonizers respectively	[157]
Bi-functional peptide with Ti binding domain and AMP domain linked by a spacer (TiBP-spacer5-AMP)	1) Can be used for initial placement or retreatment 2) MIC against S. mutans was 64 µM 2) Antibacterial and antifouling activity over 4 fouling cycles 3) Enhanced antibacterial and antifouling activity against S. mutans	[4]
Biofunctionalized Ti with Laminin 332 and ameloblastin-derived peptides- Lam and Ambn.	1) Individual coatings increased keratinocyte proliferation 2) Combined coating increased Hemidesmosome formation. Both these formations are important for creating a seal between subgingival soft tissue and the implant.	[158]
Chimeric peptide in which AMP JH8184 (derived from histatin) is linked with Ti-binding peptide (min-TBP1) using flexible/rigid linkers	1) AMP peptide connected through rigid linker killed a higher number of attached cells of S. gordonii and S. sanguis. 2) Such peptide-treated surfaces inhibit bacterial colonization for 72 h with a gradual decrease.	[159]
GLK13 cationic peptide (derived from human parotid secretory protein) mobilized on Ti micro-grooves through Silane linkers	1) No effect on adhesion of P. gingivalis but CFU was decreased. 2) Significant increase in adhesion and proliferation of HGFs	[160]
Tet213 linked to Collagen IV deposited on chitosan and hyaluronic acid through layer-by-layer deposition method	1) Sustained antibacterial activity up to 29 days against S. aureus and P. gingivalis 2) 99% inhibition rate against both the pathogens 3) Strong inhibition of early biofilm development of S. aureus	[126]

Metal coatings

For thousands of years, metals have been used as antibacterial compounds in medicine and agriculture [105]. D-block transition metals (V, Ti, Cr, Co, Ni, Cu, Zn, Tb, W, Ag, Cd, Au, Hg) and a few other metals and metalloids belonging to groups 13-16 of the periodic table (Al, Ga, Ge, As, Se, Sn, Sb, Te, Pb and Bi) have been considered for the antimicrobial applications [106]. The discovery of organic antibiotic compounds in the 1920s reduced the use of metals for antibacterial applications, but with the growing threat of antibiotic-resistant bacteria, the use of metals is getting popular again [107] Also, as metals target multiple cellular processes for killing the bacteria, a very low level of resistance has been observed against them compared to antibiotics [106]. The beneficial properties of inorganic antibacterial compounds such as chemical stability at physiological pH, thermal resistance, protracted action, broad-spectrum antibacterial activity and low propensity for bacterial resistance make them an ideal alternative for organic antibacterial compounds [35]. The mechanisms by which metals poison bacterial cells include, ROS-mediated cellular damage, site-specific damage to proteins and enzymes, membrane dysfunction due to lipid peroxidation, interference with nutrient assimilation, and genotoxicity [108, 109].

Table 4 reviews various metals deposited on the Ti surface along with their biocompatibility, antibiotic activity, and ion release studies.

Table 4.

Metal coatings on Ti surface to combat infections

Metal	Deposition method	Surface roughness	Wettability	Ion release	Antimicrobial property against	Biocompatibility	Reference
Silver (Ag)
Ag NP-coated TiO2 nanotube	Electrochemical method (TiO2 nanotubes) + chemical reduction (AgNPs) using δ-gluconolactone	ND	ND	3.35–14.6 ppm Ag release upon 24 h incubation in SBF according to different silver ammonia solutions used	S. aureus	ND	[125]
Nanoscale Ag film	DC sputtering	Increase in roughness value after deposition	ND	Extended Ag ion release up to 28 days	S. mutans, P. aeruginosa and C. albicans	The coating showed acceptable biocompatibility with human gingival fibroblasts (HGF)	[54]
Combined Ti-GO-Ag (Graphene oxide)	Electroplating + UV reduction	Increase in Ra (average roughness) with increasing GO concentration (5.36–10.6 nm)	No significant increase	Maximum release observed - 0.34 ppm	S. mutans	Microtubules disorganized, disorderly centered and reduced cell area, length, and width in MC3T3-E1	[161]
Ag nanostructure coated on TiO2 nanotube	Plasma Electrolyte Oxidation (PEO) + Magnetron sputtering	Increase in roughness if less sputtering time used	Can be varied according to sputtering time	Burst release for first three days which decreased until 7 days	S. aureus	Increase in cytotoxicity against MC3T3-E1 cells with an increase in sputtering time	[162]
Sr, Ag loaded TiO2 nanotube	Anodization + hydrothermal treatment + dip coating	ND	ND	Initial rapid release of Ag for first 3 days followed by constant 0.002 ppm/day after. A similar release trend followed by Sr ion release	Methicillin resistant S. aureus (MRSA), Methicillin sensitive S. aureus and E. coli	No cell cytotoxicity against MC3T3-E1 cells, enhanced cell spreading and expression of osteogenesis-related genes, accelerated longitudinal bone healing observed	[163]
Nanosilver-loaded poly L- lysin/sodium alginate coating	Electrostatic self-assembly + in-situ reduction of AgNPs by dopamine	ND	Increase in hydrophilicity after coating (23 ± 2°)	Sharp release of Ag for 6 days and then stabilized at ~2/3 µg per day for 21 days	S. aureus and S. mutans	Mild toxicity against MC3T3-E1 cells but mineralized coating showed excellent biocompatibility	[52]
PLGA/Ag/ZnO nanorods coating	Hydrothermal treatment (ZnO nanorods) + sol-gel method	ND	Increase in hydrophilicity after coating (CA ≤ 78°)	Slow and sustained release of Zn and Ag for 100 days	S. aureus and E. coli	Initial definite cytotoxicity was observed but increased incubation time (7 days) resulted in a lowering of cytotoxicity against MC3T3-E1, also low Ag samples showed higher ALP activity compared to control Ti	[164]
AgNPs immobilized SLA titanium	Sandblasting and acid etching (Ti-SLA) + Plasma immersion ion implantation (PIII)	No significant difference in surface roughness (Ra = 1.45 µm)	ND	Ag release in PBS after 30 days of immersion for 30-, 60- and 90-min deposited samples was 3.6, 4.3, and 4.6 ppb, respectively	S. aureus and F. nucleatum	No adverse effect on cell adhesion, spreading, proliferation and ALP activity of rBMSCs	[165]
Ag nanocoat on Ti surface	DC sputtering	Increased	Improved	Sustained release over 21 days	Streptococcus oralis, Streptococcus sanguinis, Aggregatibacter actinomycetemcomitans, Porphyromonas gingivalis, P. gingivalis 93 (Indian strain)	Increase osteoblast adhesion and related bioactivity	[142]
Copper (Cu)
Cu doped Titanium	Spark-assisted anodization method	ND	ND	ND	P. gingivalis	ND	[107]
Cu-TiO2 bio ceramic coating	Micro-arc oxidation (MAO) + Physical vapor deposition-Thermal evaporation (PVD-TE)	No significant difference in average surface roughness (~1.10 µm)	Hydrophilic surface (CA <90 which decreased as time increased	ND	S. aureus and E. coli	Greater in vitro crystalline apatite formation on Cu-MAO surface	[166]
Cu/Zn co-implanted TiN coating	Plasma immersion ion implantation (PIII)	ND	Increase in hydrophilicity, CA reduced from 98 to 78°	A stable release of ions Zn = ~0.09, Cu = ~0.17 µg/mL observed for 28 days	E. coli	Good cell adhesion, spreading and proliferation of L929 cells on Cu/Zn coated surface	[167]
Nanostructured Cu-TiO2 coating	Anodization (TiO2) + Electrodeposition (Cu-TiO2)	ND	ND	ND	S. aureus and E. coli	ND	[168]
Cu-Ti alloy	Vacuum melting and forging	ND	ND	Trace level of Cu (0.003 µg/mL per day release	S. aureus and E. coli	In vivo study in beagle dogs showed, reduced bone resorption, close contact with peri-implant bone tissue, and Higher BIC rate (39.1 ± 2.2%) than the control (7.3 ± 3.6%)	[169]
Cu-Ti alloy	Vacuum melting and forging	ND	ND	Cu ion release followed a linear trend with an average of 1.584 ×10−4 µg/mL/day for 35 days	S. mutans and P. gingivalis	No cytotoxic effects on rBMSCs, no significant difference in adhesion and spreading compared to control Ti	[170]
Zinc (Zn)
Biomimetic nano-ZnO	Hydrothermal treatment + Dip coating	ND	ND	Biphasic Zn release, initial sharp release followed by slow release up to 14 days	S. aureus and E. coli	No initial cytotoxicity but later (7 days) started showing cytotoxicity on human fibroblast cells	[171]
Ti surface with Zn-containing Nanowires	Sandblasting and acid etching (Ti-SLA) + acid (0.4% HF) and alkali (10% NaOH aq.) etching (Ti-NW) + Dip coating in 10 mM ZnSO4 (Ti-NW-Zn)	Reduction in surface roughness after modification, Ti-SLA = 2 µm Ti-NW and Ti-NW-Zn = 1 µm	Increase in hydrophilicity after modification, Ti-SLA (CA) > 90° Ti-NW and Ti-NW-Zn (CA) < 10°	Initial rapid release of Zn+2 followed by a steady high release for 4 days	S. aureus, P. gingivalis, and A. actinomycetemcomitans	MC3T3-E1 cells showed improved cell adhesion, spreading and proliferation along with higher expression of osteogenesis-related genes (Runx2, ALP, OSX, and OCN) on Ti-NW-Zn compared to Ti-SLA surface	[135]
TiO2 nanotubes decorated with nano ZnO and HA	Anodization (TiO2) + hydrothermal treatment + dip coating	Increase in surface roughness after ZnO coating but a reduction is observed after HA coating	ND	Slow release of Zn for 24 h, total Zn conc. 4.24 ± 0.36 µg/L	S. aureus	ND	[172]
Zn incorporated Chitosan/gelatin nanocomposite	Electrophoretic deposition	ND	ND	Sustained release of Zn for 4 weeks in PBS, only 10–20% of total loaded Zn was released	S. aureus and E. coli	Enhanced cell proliferation and ALP activity of rBMSCs on low Zn-containing samples, higher Zn samples showed cytotoxicity	[173]
ZnO doped with HA coating	Pulse electrochemical deposition	ND	ND	Overall release of ions decreased to a non-toxic concentration	S. aureus and E. coli	Enhanced adhesion, proliferation and differentiation of bone marrow-derived mesenchymal cells	[174]
Ti6Al4V coated with Zn-doped ZrO2/TiO2	Magnetron sputtering + Micro-arc oxidation	Increased after sputtering treatment	ND	High release concentration of Zn on the first day which gradually diminishes. No release of Zr ions	S. aureus	Higher cell proliferation of MC3T3-E1 osteoblastic cells	[175]
Gallium (Ga)
TiO2 nanopores doped with Ga	Electrochemical anodization	Increase due to deposition of Ga nano-rice	Increase in hydrophilicity	Initial burst release of 2 μg within first 10 min, peaked at day 1, declined until day 7 and sustained up to day 10	Unstimulated oral saliva sample	No cytotoxicity was observed against gingival fibroblasts even during the highest ion release.	[176]
Bilayer coating of Ga-modified chitosan/polyacrylic acid	Electrochemical deposition	Increased	ND	Initial burst within 24 h, after 48 h 36 μg was released. Until 7 days, only 14% gallium was released	E. coli, P. aeruginosa	No negative effect on osteoblastic marker (BMP2)	[177]
Ga containing calcium titanate	Chemical and heat treatments	ND	ND	Rapid release of 1.58 ppm Ga within 1 h, slow release up to 7 days, no release after that	Acinetobacter baumanni	Expected to inhibit bone resorption.	[178]
Ga and Ag ions grafted on the Ti surface	Anodic Spark deposition	Formation of micropores	ND	ND	S. mutans	No adverse effects on human gingival fibroblast viability	[179]
Other metals
Bismuth oxide coating on Ti	Magnetron Sputtering	No significant difference	No significant difference	2.67 ± 0.15 ppm	ND	No toxicity against rBMSCs and favored osteogenic differentiation	[180]
Bi-Nitrogen co-doping on Ti surface	Plasma Electrolyte Oxidation	Increased after PEO treatment	No significant difference	ND	S. sanguinis, A. naeslundii	No interference in Fibroblast cell viability, can cause oxidative stress after overproduction of ROS after photocatalytic activity.	[181]
Se incorporated on Ti nanotubes	Anodization followed by chemical dipping	Increase in roughness	ND	Increase steadily for 30 days	E. coli, S. aureus	Reduced macrophage activity under controlled Se concentration.	[170]
Ta implanted on Ti Surface	SLA + Magnetron sputtering	Decreased to 1.46 ± 0.09 mm as compared to SLA	No significant difference	ND	P. gingivalis, F. nucleatum	Increased early adhesion of rBMSCs, enhanced secretion of WNT/β catenin pathway activators, stimulated osteogenesis	[111, 182]
Co on Ti surface	DC sputtering	Increased	Improved	Cumulative release 12.6% over 28 days	Streptococcus oralis, Streptococcus sanguinis, Aggregatibacter actinomycetemcomitans, Porphyromonas gingivalis, P. gingivalis 93 (Indian strain)	Increase osteoblast adhesion and related bioactivity	[183]
TaN-Ag coating on Ti plates	Magnetron sputtering	Increased to 1.03–1.07 µm as compared to uncoated Ti	Contact angle reduced from 92° to 71°	ND	S. aureus	Improved HGF cell viability and proliferation	[184]

ND Not defined

From the foregoing literature, surface modification of implant materials is indeed a field of active research. Several studies have documented the use of planktonic cells to evaluate the surface modifications that lead to antibacterial effects. A constant depth film fermentor (CDFF) was demonstrated as a novel in vitro microcosm model that provides a high-throughput alternative for establishing biofilms resembling the progression from peri-implant health, peri-implant mucositis and peri-implantitis [110] reported changes in the microbial community associated with these conditions. Similar studies need to be carried out to evaluate the modified implant surfaces.

Recent trends in dental implant modifications

Abutment modifications for soft tissue integration

Currently, it is unclear to what extent various titanium abutment modifications influence soft tissue response and bacterial invasion. Soft tissue integration with implant material is of utmost importance for the long-term success of dental implants. If adhesion is not proper, pathogens may have the opportunity to evade the junctional epithelium interface, and bacterial evasion can lead to the dislodging of the implant. Canullo et al. [17] reviewed available literature to understand the effect of abutment modifications on peri-implant tissue behavior. According to their analysis, surface modifications of Ti abutment do not negatively affect the peri-implant soft tissue. But, at the same time, there are contrasting results when assessed for long duration and modification technique. Llopis-Grimalt et al. successfully synthesized a nanostructured titanium surface (Nanonets) with oriented cell alignment to promote soft and hard tissue integration and proved the same by monitoring the expression of differentiation markers [60].

3D printed implant surface with in-built modifications

With advancements in technology, 3D-printed implants are gaining popularity in biomedical implants. 3D printing is a powerful tool that allows personalized designing and manufacturing with high efficiency. Zhang et al. were successful in 3D printing Ti6Al4V surfaces with hierarchical hybrid structures ranging from nanoscale to macroscale, just like natural bone. SLA-modified Ti implants are the only modified implants that display hierarchical structures that include micron-scale concaves created because of sandblasting and sub-micron level pits caused by acid etching. Such SLA implants are well-documented to promote osseointegration. In a contemporary manner, 3D-printed Ti6Al4V inherently possesses micron-level concaves on the surface, like sandblasting. They subsequently employed acid etching to these 3D printed implants (3DA) to create sub-micron level pits, creating a hierarchical textured surface. 3DA surface displayed enhanced adhesion, differentiation, and proliferation of bone marrow-derived mesenchymal stromal cells (BMSCs) as compared to traditional SLA. Despite such results, the overall performance of 3DA surface was sub-par as compared to traditional SLA especially in osteogenic gene expression levels [111].

Smart implant surfaces

The formation of bacterial biofilms on the implant surface, leading to infections and subsequent implant failure, is a major concern in dental implants. As a result, stimuli-responsive strategies to activate antibacterial activity are currently gaining momentum. These stimuli can be either endogenous (pathological characters) or exogenous (photo, ultrasound). Fang et al. were successful in designing an antibacterial phototherapeutic system by combining 3 entities, which included polydopamine (PDA), Black-Phosphorous (BP) nanosheets, and ZnO nanowires on Ti substrate. The surface was irradiated with 808 nm light (Near-InfraRed), immense heat was produced from PDA and BP to dissipate bacterial biofilms, thus inhibiting biofilm formation. The authors reported >98% antibiofilm activity both in vitro and in vivo [112].

Lee et al. developed a novel thermo-sensitive anti-fouling Ti surface, which was successful in detaching bacteria after buffer rinse. They coated the Ti surface with poly glycidyl methacrylate (pGMA) using the iCVD technique and then chemically bound PIPAAM (poly N-isopropyl acrylamide), which is a thermo-responsive polymer. PIPAAM molecules are fully hydrated at temperatures below 32 °C and dehydrated at temperatures above 37 °C (oral temperature). These molecules, when hydrated, do not allow bacterial adhesion, and hence oral temperature can be decreased below 32 °C following a buffer rinse, detaching the bacteria [113]. This antifouling strategy can be useful to prevent peri-implantitis during the healing period.

Chair-side diagnostics

Considering the emergence of implants and the broader oral healthcare industry, there is a need to develop reliable analytical techniques for detecting oral diseases and other syndromes. These techniques should be capable of sensitive, accurate, rapid, chair-side diagnostics that satisfy the needs of the dentist. According to Singh et al., human oral saliva consists of biomarkers and components of the serum proteome, which can help in the identification of oral and systemic diseases [110]. These biomarkers can be evaluated and utilized as a valuable tool for assessing the activity of dental implant coatings. Albeshri et al. reviewed various biomarkers as predictors of bone regeneration around different biomaterials. Characterizing such biomarkers using non-invasive and inexpensive methods will allow us to evaluate the success of implant biomaterials [112].

Artificial intelligence (AI)-based tools

Utilizing AI and its tools is a relatively modern and novel method in the field of implantology. Artificial Intelligence (AI) and its various branches, such as Machine Learning (ML), Deep Learning (DL), and Neural Networks (NN), are utilized for various purposes such as assessing implant placement feasibility and forecasting post-placement success. AI is being utilized for clinical diagnosis and treatment of cysts and tumors, dental caries and periapical lesions, gingivitis, periodontitis, orthodontic treatment etc. However, the prognosis of implants, which involves determining the marginal bone level, peri-implantitis, osseointegration, and dental implant failure, has not yet made use of these technologies. A recent scoping analysis suggests that AI partially satisfies clinical demands when taking implants’ prognosis into account. Since altering the implant surfaces is suggested as a means of achieving success, algorithms for assessing in vitro bioactivity (anti-biofilm effect and osseointegration) based on pre- and post-images may be developed in the future to supplement radiological data and images [113].

Conclusion

Dental implants, being transmucosal in nature and exposed to oral cavity pathogens, present a considerable challenge to long-term success. Surface modifications of the implant, including physical, chemical, mechanical, or other techniques, help in the initiation of tissue regeneration by enhancing cell adhesion, migration, proliferation, and differentiation. Especially in compromised individuals with diabetes, osteoporosis, etc., the understanding of the mechanisms underlying enhanced osseointegration and wound healing is still limited, and research will help to ensure the success of dental implants. Interactions against human immune cells are unaccounted for and must be assessed to launch surface-modified implants into clinical settings. After reviewing numerous articles, the lack of in vivo research for surface-modified implants is evident. Only after considerable in vivo testing using relevant animal models will it be possible to establish the actual applications of surface-modified implants. Cost-effectiveness is always a daunting question in developing countries where esthetics-based dental treatments are not affordable to everyone. Considering all the techniques available to modify Ti with antibacterial activities and study specific modes of action, synthesis mechanisms must be researched to avoid long-term complications. Another concern is the lack of a universal standard for surface modifications accepted by both the scientific community and regulatory authorities. With the research and development in the ‘omics’, a tailored implant surface modification protocol will soon be available.

Compliance with ethical standards

Conflict of interest

The authors declare no competing interests.