Plasma spraying technology is a method in which a plasma electric arc driven by a direct current is used as a heat source to heat materials like ceramics, alloys, and metals to a molten or semi-fused state before spraying the surface of a pretreated workpiece at a high speed to form a firmly adhered surface layer. Plasma spraying is an effective method to prepare bioactive ceramic coatings. Hydroxyapatite [Ca₁₀(PO₄)₆(OH)₂, HA] is a calcium hydroxide and tricalcium phosphate compound salt with a chemical composition and crystalline structure similar to the main minerals in human bones and teeth. It is also the main inorganic component of human bone tissue, and a typical bioactive material with good biocompatibility and chemical stability. It has been reported that spraying a hydroxyapatite ceramic coating on the surface of titanium-based implants leads to good cellular compatibility, promotes adhesion, proliferation, and osteogenic differentiation of BMSCs, and improves the implant’s bond to surrounding bone tissue. In one study, Dimitrievska et al[26] fabricated a new type of titanium alloy that possesses a layer of hydroxyapatite on titanium dioxide by plasma spraying. They studied the behavior of BMSCs on this titanium-based material. The results show that cells have stronger initial adhesion (improved by 20% after 2 h) and higher metabolic activity (improved by 20% after 2 h) on TiO₂-HA compared to the titanium dioxide group. Furthermore, the differentiation of BMSCs is evidenced by alkaline phosphatase (ALP) and osteocalcin (OCN), early indicators of osteogenic differentiation, which are significantly increased on TiO₂-HA. However, the pure HA coating also has some serious defects: High brittleness, poor fatigue resistance, and weak bonding strength with metal substrates. Porous tantalum has attracted much attention for its good biocompatibility and microstructure similar to cancellous bone[27]. In a recent study, Ta-incorporated HA coatings were developed by Lu et al[28] using the plasma spray technique on a titanium substrate. The result demonstrated that Ta-incorporated HA coating could promote initial adhesion and faster cell proliferation after incubation for 3 and 5 days, but it also promotes osteogenic differentiation of BMSCs compared to HA coatings. Akermanite ceramics can induce apatite mineralization. They also have moderate stability in simulated body fluid (SBF) and generally good mechanical properties, and support BMSC attachment[29,30]. The researchers found that the bonding strength between the plasma-sprayed akermanite bioactive coatings and Ti substrates is higher than hydroxyapatite (HA) coatings, and BMSC attachment and proliferation were more significant on akermanite coatings than on HA coatings[31].

Sol–gel process first described 150 years ago is still receiving great attention as one of the easiest ways to develop modified materials which possess required properties and are characterized by durability and stability[32]. Hence, sol-gel process is another method for preparing bioactive ceramic coatings. The sol-gel technology has some advantages compared to plasma spraying methods, including chemical uniformity, fine grain structure, and lower processing temperature[33]. In a study, a micro/nano-layered structure was prepared on a micro-structured titanium (Micro-Ti) substrate using a sol-gel method with a spin coating technique. The results confirmed that the micro/nano-level structure of large particles (80 nm) significantly promoted MSC proliferation and differentiation compared to other small particles (20 nm and 40 nm)[23]. Inzunza et al[34] prepared nanoporous silica coatings on Ti using the sol-gel method and evaporation-induced self-assembly method. The silica coatings with highly ordered sub-10 nm porosity accelerate the adhesive response of early BMSCs and promote BMSC osteogenic differentiation.

Chemical methods can be used to increase the thickness of the oxide film to improve the biocompatibility and bioactivity of titanium and its alloys. The surface chemical treatment of titanium and titanium alloys mainly includes alkali treatment, acid treatment, and acid-base two-step treatment. Alkali solution is used to modify the titanium surface to obtain sodium titanate gel with rich Ti-OH groups on the surface, endowing it with biological activity[36,37]. For this purpose, Cai and his team employed potassium hydroxide to modify the surfaces of titanium substrates; the formed potassium titanate layer enhances titanium’s corrosion resistance. The proliferation and differentiation levels of alkaline phosphatase and osteocalcin were significantly increased in MSCs cultured on alkaline-treated titanium after 7 and 14 d of culture, respectively[38].

Acid treatment is often used to remove the oxide layer and contaminants on the surface of the medical titanium material to obtain a clean and uniform surface. The acid treatment results in a 10-nm thick oxide layer, while the titanium oxide in the air is only 3-6 nm thick[39]. Maekawa et al[40] treated titanium with polyphosphoric acid solution for 24 h at 37 °C. Surface texture measurement results show that the maximum surface roughness of the treated titanium surface significantly increased. Significantly higher cell attachment and proliferation were also found on titanium treated with polyphosphoric acid in contrast to untreated titanium (control). By comparing the effects of acid-treated titanium and pure titanium on osteogenic differentiation of bone MSCs, Perrotti et al[41] concluded that 1 wk of treatment was more than enough for osteoblast differentiation on acid-treated titanium. Silva and his group suggested that rough surfaces submitted to acid-etching favor undifferentiated mesenchymal cell differentiation into osteogenic lineage cells compared to smooth titanium surfaces without acid treatment[42]. Although many studies have shown that surface acidification can increase the degree of roughening and improve the biological activity of titanium implants, acid treatment may cause hydrogen to penetrate below the oxide layer, thereby triggering hydrogen embrittlement[43].

The acid-alkali two-step method is also used for titanium surface modification. Strong acid erosion could cause micropores on the surface of titanium and titanium alloys to increase surface area. Meanwhile, alkaline solution can form a thicker microporous titanium oxide layer on the titanium surface, improving the titanium implant’s biological activity[44-46]. Li et al [47] first placed titanium in oxalic acid solution (5 wt%) at 100 °C for 2 h to remove the oxide layer and acquire a homogeneous micropit surface. Each pretreated titanium plate was treated in 5 mmol/L NaOH solution at 80 °C for 24 h. An in vitro cell experiment demonstrated that BMSC adhesion and osteogenesis can be better promoted on a micro/nanoporous surface than on an acid etched titanium surface. However, BMSC proliferation was significantly inhibited on treated surfaces after culturing for 4 and 7 days, which may due to the high pH around the implant. The high pH at the cell/material interface may cause alkalosis and inhibit BMSC proliferation and viability[48].

Hydrogen peroxide can also be used for activation treatment of titanium. Hydrogen peroxide treatment of titanium is a chemical dissolution and oxidation process, which could alter surface roughness, thickness, and hydrophilicity, with improvements in titanium osteoconductivity[49-51]. In one study, titanium was treated with 30% volume (v/v) of H₂O₂ (5 mL H₂O₂/g disc) for different times in an unsealed covered container under darkness at room temperature. The modifications induced by 6-24 h H₂O₂-treated surfaces are most beneficial for maintaining or promoting the attachment, proliferation, and osteogenic differentiation of BMSCs[52].

Anodization refers to the use of an electric field and various dilute acids as electrolyte solutions. A series of REDOX chemical reactions take place on the anode surface to form an oxidation layer. Due to anodization’s simplicity, versatility, and low cost, it has gained widespread attention in the surface treatment of titanium implants. In a study, Xu et al[24] found that tube diameter had a significant effect on adhesion, proliferation, and differentiation of MSCs. Titanium was used as the working electrode, platinum sheet was used as the cathode, and 0.50 wt% NH₄F + 10 vol% H₂O mixture was used as the electrolyte. The anodic oxidation was carried out at 10, 30, and 60 V, which were designated as NT10, NT30, and NT60, respectively. Finally, NT10, NT30, and NT60 were obtained with pore diameters of 30, 100, and 200 nm, respectively. By comparison, although NT60 can promote osteogenic differentiation to the greatest extent, it significantly inhibits cell adhesion and proliferation. NT10 can promote cell proliferation and adhesion, but it is useless for osteogenic differentiation of cells. NT30 supported adhesion and proliferation of BMSCs, and the cells on NT30 became increasingly elongated with increased diameter and showed a large number of prominent filamentous pseudopods. Moreover, it showed better osteogenesis-inducing ability. In another study, Grimalt et al[53] produced a nanonets structure on titanium discs. BMSCs cultured on nanonet structured titanium surfaces present a high frequency of alignment and promote osteogenic differentiation of the cells, while cells on untreated titanium surfaces exhibited a random orientation.

Micro-arc oxidation (MAO) is a new type of anodic oxidation technology that deposits a ceramic coating on the metal surface, and it has been widely applied in the surface modification of titanium and its alloys to enhance biological activity and osteogenic capacity. Based on ordinary anodization, arc discharge is used to enhance and activate the reaction occurring on the anode, thereby forming a ceramic film in situ on the surface of titanium[54-56]. Zhou et al[25] reported that porous coatings prepared by MAO promote BMSC adhesion and osteogenic differentiation. In addition, the larger the pore size, the more conducive to BMSC adhesion and osteogenic differentiation when the pore size is in the range of 3-10 μm. A similar phenomenon was observed in BMSCs in another study. Li et al[57] developed two kinds of coatings (MAO and MAO-Alkali coatings) with similar micro-morphologies, both of which significantly promote BMSC adhesion and osteogenic differentiation by mediating the integrin β1 signaling pathway.

Plasma ion implantation (PIII) is known to modify the surface and near surface regions of materials, and it has many advantages for surface modification of materials, including the following: (1) Changing the surface characteristics of the material alone without affecting the properties of the material; (2) The modified layer will not fall off or fail in combination; and (3) PIII is a low-temperature process (approximately 100 °C), and there is no change in the size of the workpiece due to thermal distortion.

PIII surface modification mainly uses plasma generated after Ar, N₂, O₂, and other gases or metal gasification to treat the material surface. Under the action of plasma, the surface of the material is bombarded with high-energy particles in the plasma. Chemical bond breakage occurs, and large molecular radicals are generated. At the same time, the material is etched to change the surface properties. PIII of metal materials can effectively improve the mechanical properties, wear resistance, and corrosion resistance of orthopedic implants, thus enhancing their biocompatibility[58]. Yang et al[59] explored the effect of titanium treated with oxygen plasma immersion ion implantation (O-PIII) on the behavior of BMSCs with different oxygen doses. The results showed that O-PIII treatment could enhance BMSC adhesion, and there was no significant difference in the titanium surface treated with O-PIII when the oxygen ion dose differed. In their later study, Yang et al[60] compared the effects of three doses of oxygen ion implantation into titanium on BMSC behavior. Among these treated titanium disks, the group treated with the highest concentration of oxygen ions has the best effect on cell adhesion, migration, proliferation, mineralization, and differentiation of BMSCs. It has been reported that calcium-ion-implanted titanium also remarkably improved BMSC adhesion and proliferation compared to the untreated sample[61]. Similarly, other studies have evaluated the response of BMSCs to titanium surfaces that had been implanted with Ca and Mg ions using the PIIID technique. The results showed that initial cell attachment on a titanium surface can be improved by Ca and Mg ion implantation. Cells on the Mg ion-implanted surface showed more extended filopodia after 4 and 24 h of cultivation. In addition, the expression of genes associated with osteogenic differentiation like RUNX₂ and type I collagen was higher in the Mg ion-implanted surface[62]. These results are consistent with previous studies showing that significant cytotoxicity was not observed after Mg ion implantation into a titanium implant, and initial BMSC adhesion was improved with resulting osteoblast differentiation enhancement[63].

Laser beam treatment is a controllable and flexible approach to modifying surfaces, which results in surfaces with increased surface area and enhanced wettability, and it displays negligible corrosion and high removal torques of established implants in preclinical bone models[64,65]. Laser-modified titanium surfaces could enhance upregulation of expression of the osteogenic markers and enhance alkaline phosphatase activity of BMSCs[66]. A recent investigation on the direct metal laser sintering (DMLS) titanium surface found that topographical cues of DMLS surfaces could enhance both protein adsorption ability and BMSC adhesion performance. Moreover, DMLS titanium surface could efficiently induce osteogenesis-associated gene expression in BMSCs via H3K27 demethylation and increases in H3K4me3 levels at gene promoters after osteogenic differentiation[67]. In another study, dynamic analyses of early cellular events showed that BMSCs exhibited a more elongated, spindle-like morphology and higher spreading speeds on femtosecond laser-modified surfaces compared to commercially pure titanium[68].

The extracellular matrix (ECM) is composed of several molecules secreted by cells. In addition to providing structural and mechanical support for tissues to interact with cells, these molecules can also bind to soluble molecules like growth factors that are present in extracellular fluid and regulate the occurrence of tissues and physiological activities of cells. The ECM provides a framework for tissue construction and plays an important role in regulating the survival, migration, proliferation, morphology, and other functions of cells in contact with it. Therefore, ECM components are the first choice for the biochemical surface modification of titanium-based bone implant materials.

It has been found that some short peptides in ECM proteins play important roles in cell behavior regulation[80,81]. Among different ECM proteins, fibronectin (FN), a multifunctional cell adhesive glycoprotein, is one of the most well-known and commonly used to functionalize biological materials. It contains several domains that mediate many cellular processes like cell adhesion, migration, growth, and differentiation. The use of FN-functionalized titanium implants has been shown to improve bone conduction capacity for its ability to attach cells to ECM components via integrin receptor interactions[82]. Chen et al[83] fixed FN on the surface of titanium, and BMSCs exhibited substantial actin polymerization, in the form of lamellipodia, pseudopodia, and actin stress fiber. However, the cells retained a rounded morphology on untreated surface. Besides, FN-functionalized titanium had a significant positive effect on BMSC proliferation compared to the control. However, its use for clinical applications is hampered due to poor stability, high production costs, and poor ECM protein immunogenicity, which have reduced their biomedical potential[84]. The use of ECM-derived synthetic peptides containing the functional domains of ECM proteins is an effective method to overcome these problems. Therefore, the synthesis of short peptide fragments representing ECM proteins and the modification of titanium-based implants have been gradually developed[85,86]. The most commonly used peptide sequence for surface modification is the arginine-glycine-aspartic acid (RGD) motif. RGD-functionalized titanium can improve early bone growth and matrix mineralization, and it can enhance the combination of materials and new bone[87]. There have been several reports on the effects of RGD on BMSCs. In a study, Karaman et al[88] covalently attached RGD peptide to titanium discs. The results indicated that RGD peptide treatment significantly enhanced BMSC adhesion and proliferation. Furthermore, this effect was enhanced by combining cold temperature plasma treatment and RGD peptide coating. Consistent with this, Herranz et al[89] concluded that the RGD motif was more favorable for BMSC adhesion, proliferation, and osteogenic differentiation in contrast to fibronectin. In another study, Jordi and his group covalently attached a novel molecule on the titanium surface. The novel molecule possesses adhesion capacity by an RGD gain-of function DNA mutation installed to the heparin binding II (HBII) fragment. The presence of RGD in the HBII domain stimulated focal adhesion formation at BMSC edges where filopodia were spikier compared to bare titanium samples with completely round cells. In addition, HBII-RGD-functionalized titanium surfaces could also stimulate BMSC differentiation and mineralization[90].

Growth factors are a class of proteins secreted by cells that act as signaling mediators for the relevant target cells to perform specific behaviors. Growth factors can promote cell proliferation, differentiation, protein synthesis, and migration of specific cells. Growth factors released from the implant surface can increase osteoblast activity and facilitate bone tissue regeneration[91]. Many researchers have been depositing growth factors on biomaterials to affect cell behavior. In one study, Bauer et al[92] showed the covalent immobilization of two growth factors, epidermal growth factor (EGF) and bone morphogenetic protein-2 (BMP-2), on the surface of TiO₂ nanotubes and their effects on BMSC behavior. Cell adhesion and proliferation were dramatically increased by covalently grafting EGF on a surface of a 100 nm nanotube, but covalently grafted BMP-2 did not. The result was consistent with the finding of previous studies that BMP-2 promotes BMSC differentiation into osteoblast lineages but does not contribute to the cell attachment, adhesion, or proliferation like EGF[93]. Studies on BMP-2’s effect on BMSC differentiation have shown that BMP-2 has a significant effect on osteoblast differentiation potential[94].

Platelet derived growth factor (PDGF) has been shown to play critical roles in bone regeneration after injury, and it significantly contributes to all stages of bone regeneration after trauma[95,96]. Among three types of dimerism, PDGF-AA, -BB, and –AB, PDGF-BB exerts the most potent chemotactic effects on BMSCs[97]. Ma et al[98] fabricated a nano-micro hierarchical TiO₂ clustered nanotubular structure using anodization, and PDGF-BB was functionalized with PhoA (11-hydroxyphosphonic acid)/CDI (carbonyldiimidazole). The resulting new material had almost no cytotoxicity to host cells, and it significantly enhanced BMSC attachment and osteogenesis-related functions (early proliferation, extracellular matrix synthesis, and mineralization).