
==== Front
Bioact Mater
Bioact Mater
Bioactive Materials
2452-199X
KeAi Publishing

S2452-199X(24)00380-3
10.1016/j.bioactmat.2024.08.042
Review Article
Surface modification strategies to reinforce the soft tissue seal at transmucosal region of dental implants
Jin Siqi 2311210517@bjmu.edu.cn
a1
Yu Yameng 2216395033@bjmu.edu.cn
a1
Zhang Ting 2201112423@stu.pku.edu.cn
b
Xie Daping yb97519@umac.mo
c
Zheng Yufeng yfzheng@pku.edu.cn
be⁎⁎
Wang Chunming cmwang@umac.mo
c⁎⁎⁎
Liu Yunsong liuyunsong@hsc.pku.edu.cn
d⁎
Xia Dandan dandanxia@pku.edu.cn
a⁎⁎⁎⁎
a Department of Dental Materials, Peking University School and Hospital of Stomatology & National Center for Stomatology & National Clinical Research Center for Oral Diseases & National Engineering Research Center of Oral Biomaterials and Digital Medical Devices & Beijing Key Laboratory of Digital Stomatology & NHC Key Laboratory of Digital Stomatology & NMPA Key Laboratory for Dental Materials, Beijing, 100081, China
b School of Materials Science and Engineering, Peking University, Beijing, 100871, China
c State Key Laboratory in Quality Research of Chinese Medicine, Institute of Chinese Medical Sciences, University of Macau, Macau SAR, 999078, China
d Department of Prosthodontics, Peking University School and Hospital of Stomatology, Beijing, 100081, China
e International Research Organization for Advanced Science and Technology (IROAST), Kumamoto University, 2-39-2 Kumamoto, 860-8555, Japan
⁎ Corresponding author. liuyunsong@hsc.pku.edu.cn
⁎⁎ Corresponding author. School of Materials Science and Engineering, Peking University, Beijing, 100871, China. yfzheng@pku.edu.cn
⁎⁎⁎ Corresponding author. cmwang@umac.mo
⁎⁎⁎⁎ Corresponding author. Department of Dental Materials, Peking University School and Hospital of Stomatology, Beijing, 100081, China. dandanxia@pku.edu.cn
1 Equal contribution, Joint first authors.

10 9 2024
12 2024
10 9 2024
42 404432
20 7 2024
29 8 2024
29 8 2024
© 2024 The Authors
2024
https://creativecommons.org/licenses/by-nc-nd/4.0/ This is an open access article under the CC BY-NC-ND license (http://creativecommons.org/licenses/by-nc-nd/4.0/).
Soft tissue seal around the transmucosal region of dental implants is crucial for shielding oral bacterial invasion and guaranteeing the long-term functioning of implants. Compared with the robust periodontal tissue barrier around a natural tooth, the peri-implant mucosa presents a lower bonding efficiency to the transmucosal region of dental implants, due to physiological structural differences. As such, the weaker soft tissue seal around the transmucosal region can be easily broken by oral pathogens, which may stimulate serious inflammatory responses and lead to the development of peri-implant mucositis. Without timely treatment, the curable peri-implant mucositis would evolve into irreversible peri-implantitis, finally causing the failure of implantation. Herein, this review has summarized current surface modification strategies for the transmucosal region of dental implants with improved soft tissue bonding capacities (e.g., improving surface wettability, fabricating micro/nano topographies, altering the surface chemical composition and constructing bioactive coatings). Furthermore, the surfaces with advanced soft tissue bonding abilities can be incorporated with antibacterial properties to prevent infections, and/or with immunomodulatory designs to facilitate the establishment of soft tissue seal. Finally, we proposed future research orientations for developing multifunctional surfaces, thus establishing a firm soft tissue seal at the transmucosal region and achieving the long-term predictability of dental implants.

Graphical abstract

Image 1

Highlights

• The inferior sealing capacity of peri-implant soft tissue increases the risk of implant failure.

• Adhesion-promoting, antibacterial, and immunomodulatory design strategies were proposed to enhance peri-implant soft tissue seal.

• The existing obstacles and potential future research orientations were discussed to construct advanced surfaces.

Keywords

Surface modifications
Dental implant transmucosal region
Soft tissue seal
Antibacterial
Immunomodulation
==== Body
pmc1 Introduction

Soft tissue seal, also referred to as soft tissue integration, is the attaching of epithelium and underlying connective tissue of peri-implant mucosa to the transmucosal region of dental implants [1,2]. Unlike that osseointegration describes a solid and compact two-part interface between the implant and bone [3], soft tissue seal exists as a relatively vulnerable soft tissue barrier that helps in sealing the implant-gingiva interface and isolating the underlying osseointegrated implant part from the highly contaminated oral cavity. Once the soft tissue seal is breached, bacteria in the oral cavity might invade the interface and cause peri-implant mucositis and peri-implantitis, leading to irreversible marginal bone loss [4], thus endangering the survival of dental implants. It was reported in the 11th European Workshop of Periodontology that the prevalence of peri-implant mucositis and peri-implantitis was 19 %∼65 % and 1 %∼47 %, respectively, based on various case definitions [5]. Under the most conservative estimates of D.P. Tarnow, about 10 % of all dental implants would develop a certain degree of peri-implantitis in approximately 10 years [6]. Though treatments for peri-implantitis produce positive results, it is more recommended to pay particular attention to the prevention of peri-implantitis. As such, establishing a firm soft tissue seal and preventing initial bacterial infections are of great significance, which aid in achieving the long-term survival and functioning of dental implants. However, in some failure cases, biofilm and bacterial metabolism product might have an indirect or negligible role in the loosening process, where the excessive immune responses may induce aseptic loosening [7]. To avoid the negative influence of uncontrolled immune responses on the establishment of soft tissue seal, as well as creating a repair-supportive immune microenvironment for tissue regeneration, immunomodulatory therapies should also be considered when constructing advanced transmucosal surfaces.

Nowadays, titanium (Ti) and yttria-stabilized tetragonal zirconia polycrystals (Y-TZP) are dominating materials applied in the dental implant transmucosal region, while polymeric poly-ether-ether-ketone (PEEK) has a predicable application prospect due to its outstanding material properties [8]. The transmucosal region made of the above materials was confirmed to be capable of supporting the attachment of peri-implant mucosa and establishing a biocompatible soft tissue seal rather than fibrous encapsulation [[9], [10], [11], [12], [13]]. Nevertheless, the pristine form of such materials cannot actively promote transmucosal soft tissue attachment, so that the soft tissue seal at transmucosal region of dental implants exhibits inferior sealing capacity compared with the integrated periodontal tissue around a natural tooth [14]. Moreover, the risks of bacterial infection and inflammatory responses around the materials should also be considered. From the viewpoint of bacterial infection, the unmodified surface lacks intrinsic abilities to defend against bacterial invasion, where bacterial colonization would adversely affect tissue cell adhesion and impair the establishment of soft tissue seal [2]. From the viewpoint of immune responses, the pristine materials might lack the properties to actively facilitate the formation of a pro-regenerative immune microenvironment for soft tissue. Specifically, the wear debris and ions released from the implanted devices have been shown to elicit inflammatory response [15]. This suggests that pristine Ti/Y-TZP/PEEK materials hardly contribute to the attenuation of inflammation and soft tissue regeneration.

Therefore, for accomplishing a robust peri-implant soft tissue seal, the ideal transmucosal surface would be necessarily equipped with the fundamental property of effectively promoting soft tissue attachment. On this basis, the transmucosal surface is also expected to be capable of eliminating bacterial colonization and regulating the immune microenvironment towards a repair-supportive state, decreasing the negative influence of infection and inflammation on the establishment of peri-implant soft tissue seal. Over the decades, numerous studies have taken efforts to improve the sealing capacity of transmucosal surface, with several articles summarizing the modifying methods in terms of certain materials, mainly focusing on the promotion of soft tissue attachment, yet offering limited description on the surface antibacterial properties and immunomodulating abilities [[16], [17], [18]]. Thus, surface modifications strategies include promoting soft tissue attachment, balancing antibacterial effect and soft tissue attachment, as well as immunomodulation are imporant aspects to reinforce the soft tissue seal.

Herein, an updated overview of the transmucosal region surface modifications in terms of adhesion-promoting, antibacterial, and immunomodulatory design strategies was proposed. In this review, we firstly depicted the structural features of the interface between the transmucosal region and peri-implant mucosa, and elucidated the biological healing process of peri-implant mucosa. Accordingly, current surface modification strategies at the transmucosal region made of Ti/Y-TZP/PEEK aiming to facilitate soft tissue attachment to the implant surface were presented. Moreover, rationales for preventing bacterial infections and regulating the immune microenvironment were summarized. Finally, we briefly concluded the existing obstacles in constructing versatile surfaces and proposed potential future research orientations.

2 Structural differences of the interface between tooth-gingiva or implant-gingiva

One of the main differences between natural teeth and dental implants lies in their connective structures with surrounding tissues. As indicated in Fig. 1, a natural tooth depends on junctional epithelium, connective tissue and periodontal ligament to connect with gingiva and alveolar bone, where the junctional epithelium and connective tissue constitute the dimension of biological width [19]. Correspondingly, a dental implant depends on peri-implant junctional epithelium and connective tissue to form soft tissue seal [20], and directly contacts with bone to form osseointegration [3]. Due to their structural differences, the sealing capacity of peri-implant soft tissue is inferior to periodontal soft tissue. In specific, the transmucosal surface of a dental implant is in absence of natural substrates such as enamel and cementum, which fails to induce an extensively distributed hemidesmosomes or a perpendicularly protrusion of collagen fibers, compared with that of a natural tooth [22,23]. Such structural differences make the peri-implant soft tissue seal more susceptible to bacterial invasion and inflammatory destruction, compared with a natural tooth.Fig. 1 Illustration showing the biological width around a natural toot and the soft tissue seal around a dental implant. (a) The attaching structures of periodontal biological width, which exhibits good defensive ability against bacterial invasion. (b) The attaching structures of peri-implant soft tissue seal, which exhibits inferior defensive ability against bacterial invasion.

Fig. 1

In this review, we adopt “transmucosal region” to refer to either the abutment in a submerged system, or the implant collar in a non-submerged system. In particular, the transmucosal region of a dental implant, which faces the peri-implant mucosa and constituting the implant-gingiva interface, describes the abutment part or the collar of implant [24]. In a submerged dental implant system, the top of implant is at the level of bone surface, where an abutment is further required to contact with the surrounding soft tissue. In a non-submerged system, the top of implant is at the level of gingival margin, thereby covering the transmucosal region. Clinically, the submerged system receives wider applications due to the separated structures, which optimizes the transfer of occlusal force and attenuates marginal bone resorption [25].

2.1 Histological features of the tooth/implant-gingiva interface

Around the natural tooth, a firm soft tissue barrier sealing the enamel-epithelium interface is composed of junctional epithelium and internal basal lamina, which is the specific basement membrane-like extracellular matrix (ECM) mainly containing laminin 332 (also known as laminin 5) [26]. Laminin 332 plays an essential role in initiating the formation of hemidesmosome assemblies on epithelial cell membranes, thus anchoring junctional epithelium to the enamel surface [27]. Below the enamel-epithelium interface is the cementum-connective tissue interface. The ECM of connective tissue is composed of bulk collagen fibers, ubiquitous proteoglycans and adhesive glycoproteins (mainly fibronectin) [28]. Collagen I is the dominating ECM component in this region, with the main cell type being fibroblasts which tend to exist in isolation within the pool of collagen fibers [28]. Moreover, the attaching of connective tissue to the tooth primarily depends on the direct protrusion of collagen fibers (Sharpey's fibers) into the cementum [29]. On the other hand, fibronectin works as an adhesive in the binding of fibroblasts to ECM components, mainly collagen fibers [28].

Around the transmucosal region of dental implant, similar features were identified. It has been demonstrated that internal basal lamina and hemidesmosomes also existed, yet only in the lower apical region, at the interface between Ti transmucosal region and peri-implant junctional epithelium [22]. Meanwhile, a histological examination revealed that the peri-implant connective tissue was also in direct contact with the Ti transmucosal region [30]. Moreover interestingly, through a detailed analysis of the 200 μm wide region next to the transmucosal region, a 40 μm wide region featuring abundant fibroblasts interposed between thin collagen fibers which located immediately adjacent to the implant was observed, while the outer 160 μm wide region appeared to be less cellular and contained more collagen fibers and blood vessels, which could be seen as a transitional area towards typical connective tissue. The fibroblasts-rich region adjoining the implant transmucosal region might be responsible of producing collagen and adhesive glycoproteins, which would facilitate the tissue attachment to the implant surface.

Though having similar structures, peri-implant soft tissue seal only forms an inferior barrier compared with the biological width, due to epithelial downgrowth and the scar-like connective tissue around the transmucosal region. In specific, epithelial downgrowth is coronal-apical migration and proliferation of epithelial cells, resulting in a lower attachment of peri-implant junctional epithelium and a longer length of peri-implant sulcus/pocket [31]. On the other hand, peri-implant connective tissue is in analogous to scar tissue rather than periodontal connective tissue, since the tissue is less vascularized and the collagen fibers are arranged in a parallel orientation, undermining the tissue to be more prone to fibrosis and inflammation [23].

2.2 Protein-protein interactions involved in the tooth/implant-gingiva interface

At the epithelium level, internal basal lamina is characterized in abundant laminin 332 and the absence of collagen IV or collagen VII, differing from typical basement membrane [26,32]. Laminin 332 is a heterotrimer composed of α3, β3 and γ2 subunits [33], expressed and secreted by epithelial cells. There is a 14-amino-acid motif (FFPLMLLKGSTRFC) in large globule domain 3 of the laminin 332 α3 chain (LamLG3) interacting with its receptor integrin α6β4 [34]. Upon binding to the integrin α6β4 on the membrane of epithelial cell, laminin 332 activates downstream phosphorylation of phosphatidylinositol-3-kinase (PI3K) and Serine/Threonine kinase (Akt, also called protein kinase B, PKB) to initiate the formation of the hemidesmosomes [35], which enables the epithelial cell to robustly anchor to the surface of the tooth or material. The hemidesmosome assembly basically consists of two parts, with the transmembrane part including integrin α6β4 and bullous phemphigoid antigen 180, and the cytoplasmic part containing plectin and bullous phemphigoid antigen 230 with the intermediate filaments attaching on it, together resulting in an electron-dense plaque anchoring on the inner side of the membrane of the epithelial cell [Fig. 2a∼b] [35].Fig. 2 Illustration showing the structures of hemidesmosome and focal adhesion. (a) A transmission electron microscope image of a hemidesmosome. (b) A schematic drawing of the hemidesmosome. Adapted with permission from Ref. [35]. Copyright 2006 Elsevier Ltd. (c) The organization of green fluorescent protein (GFP)-paxillin-labeled focal adhesions (green) and phalloidin-labeled filamentous actin (red), on rigid (upper panel) or soft (lower panel) surfaces. Adapted with permission from Ref. [39]. Copyright 1969 Springer Nature Limited. (d) A schematic representation of the formation of focal adhesion. Adapted with permission from Ref. [36]. Copyright 2021 Elsevier B.V.

Fig. 2

In the underlying connective tissue, the ECM can be depicted as a pool of densely arranged collagen (mainly collagen I bundles) and dispersedly distributed adhesive glycoproteins (mainly fibronectin), on which isolated fibroblasts attach and synthesize more ECM proteins [28]. Similar to the anchoring property of hemidesmosomes, focal adhesion assembly is critical to fibroblasts in the attaching process, functioning as an ECM-integrin-cytoskeleton linkage [36]. Characterized by vinculin, paxillin and focal adhesion kinase (FAK) [37], the focal adhesion mediates the fibroblast anchorage to fibronectin and collagen I via integrins [Fig. 2c∼d] [36]. Specifically, the fibronectin interacts with fibroblasts through activating integrin α5β1 and triggering downstream FAK/proto-oncogene tyrosine-protein kinase (Src) signaling pathway while integrin α2β1 transduce the bidirectional regulation between collagen I and fibroblasts [[36], [37], [38]]. In this manner, fibronectin mediates the interactive binding of collagen, fibroblasts and the cementum via integrins.

3 Temporal sequence of soft tissue healing at the implanted zone

Upon implantation, a sequence of coordinated overlapping biological processes occur at the implanted zone, involving clot formation, inflammation, connective tissue proliferation, epithelial cell migration, proliferation and sealing, collagen deposition and orientation, finally reaching tissue maturation [40].

Specifically, the healing process of peri-implant mucosa was concretely deciphered through an implantation model of Beagle dog with an observation duration of 12 weeks, where a Ti non-submerged implant system was employed [41]. The biopsies were harvested at 2nd hour, 4th day, 1st week, 2nd week, 4th week, 6th week, 8th week and 12th week after implantation. The temporal sequence was demonstrated in Fig. 3. Within 2 h after implantation, the blood clot formed in the implant-gingiva interface. At 4th day, a dense fibrin network was seen at the interface with leukocytes infiltration, while the underlying connective tissue was under a degradation stage. About 1∼2 weeks after implantation, epithelial cells derived from surrounding oral epithelium commenced to migrate to the interface and proliferate, where the connective tissue was rich in proliferating fibroblasts and blood vessels. After 2∼4 weeks, the number of fibroblasts decreased, with functioning spindle-shaped fibroblasts secreting collagen. After 4∼6 weeks, collagen fibers became orientated. After 6∼8 weeks, a mature epithelial barrier was established, while the collagen organization and tissue maturation were completed. Furthermore, the as-inferred temporal sequence was verified in humans by via a biopsy collection protocol [42].Fig. 3 The temporal sequence involved in the wound healing of peri-implant mucosa. The biological events occur after the implantation include bleeding, clotting, inflammation, cell proliferation, collagen deposition, and final tissue maturation. ECM extracellular matrix.

Fig. 3

4 Surface modification strategies for materials applied in transmucosal region

Nowadays, Ti and its alloys as well as Y-TZP are dominating materials for fabricating clinically used dental implant transmucosal region. Generally, commercially pure Ti and Ti alloys (mainly Ti6Al4V) constitute the major part of commercial dental implant transmucosal region due to their outstanding biocompatibility, superior mechanical properties and high corrosion resistance [43]. Commercially available Ti abutments have managed to adopt a novel Ti surface (Xeal™, Nobel Biocare, Sweden) featured with an arithmetic mean height (Sa) of 0.13 ± 0.02 μm and an oxide layer thickness of 153 ± 5 nm [44]. Such surface is smooth, non-porous, nanostructured and anodized, thereby capable of supporting soft tissue attachment, minimizing bacterial colonization and reaching a satisfying esthetic transmucosal color. On the other hand, Y-TZP has emerged as the second-generation clinical abutment/implant materials, which can avoid the radiological artifact, hypersensitivity of specific metal species and achieve satisfying gingival esthetic [45]. Though limited modifications on Y-TZP products were investigated to improve the soft tissue seal at transmucosal region, the Straumann® PURE Ceramic Implant System (Straumann, Switzerland) based on Y-TZP, which had a roughened transmucosal region (Sa = 1.27 ± 0.24 μm) [46], was proved to provide an ivory transmucosal color, improved soft tissue blood circulation, enhanced soft tissue seal, and less plaque attachment, compared with traditional Ti transmucosal region [[47], [48], [49]]. In addition, PEEK has been increasingly recognized as potential dental implant materials, since its elastic modulus value (3∼4 GPa) is much closer to the alveolar bone (6∼30 GPa) than that of Ti (100∼110 GPa) or Y-TZP (210 GPa) [45]. However, PEEK surface appears to be hydrophobic due to its bioinert nature, which infers an increased difficulty in integrating with the soft tissue, and requires further modifications, including improving wettability, fabricating porous topography and applying bioactive coatings [18,50].

Commercially available surface designs of abutments or implant transmucosal regions are listed in Table 1. Except for fabricating standard smooth surfaces, commercially available abutment/implant products still lack a guidance for producing a transmucosal surface that has high affinity to soft tissue. Apart from the ability to promote soft tissue attachment, it is also important for the transmucosal region to resist bacterial infection, suggested by the ‘race to invade’ theory. Specifically, once oral pathogens colonized and accumulated on the surface with biofilm formed, soft tissue cells were unable to adhere to the transmucosal region, and thus undermining the establishment of a robust soft tissue seal [2]. On the other hand, the process of placing the implant/abutment would trigger a series of immune responses and cause inflammation, which indicates the need of introducing immunomodulatory properties to the surface. Therefore, the following three chapters were developed based on the above strategies to discuss how to reinforce the transmucosal soft tissue seal [Fig. 4].Table 1 Commercially available surface designs of abutments or implant transmucosal regions.

Table 1Materials	Products or patented surface treatments	Modification methods	Surface properties	Ref.	
Ti	Xeal™ (Nobel Biocare, Sweden)	Anodization	A smooth non-porous, nanostructured and anodized surface possessing surface chemistry and topography that are specially designed to promote soft tissue attachment.	[51,52]	
Ti	Synthegra (Geass, Italy)	Laser technology	A smooth nanostructured surface that promotes cell adhesion, collagen organization and induces lower grade of inflammation compared with machined surface.	[53,54]	
Y-TZP	ZLA® (Straumann, Switzerland)	Large-grit sandblasting, acid-etching	A roughened hydrophilic surface.	[46]	
Y-TZP	CERALOG® (Camlog, Germany)	Ceramic injection molding	The neck area of the implant has a roughness with an Ra value of 0.5 μm, which additionally optimizes the deposition of soft tissue.	[46]	
Ti	Ankylos® (Dentsply Sirona, USA)	Macro-scale shape design	A concave abutment profile creates a three-dimensional space ideal for growing and maintaining soft tissue.	[55]	
Ti titanium, Y-TZP yttria-stabilized tetragonal zirconia polycrystals, Ra roughness average.

Fig. 4 Illustration showing the surface modification strategies to reinforce the soft tissue seal at transmucosal region of dental implants based on the three rationales: promoting soft tissue attachment, eliminating bacterial colonization, and regulating immune microenvironment. ECM extracellular matrix, PAR4-AP protease activated receptor 4-activating peptide, CLA conjugated linoleic acid, modSLA modified sandblasted, large-grit, acid-etched technique, IL-23 interleukin 23, CeO2 ceria, PDGF-B platelet-derived growth factor subunit B.

Fig. 4

5 Modification strategies to promote soft tissue attachment

Improve the bonding capacity of transmucosal region to peri-implant mucosa is of critical importance. In specific, the augmented bonding capacity can be realized through either promoting the cell adhesion (mainly the epithelial cells and the fibroblasts) with increased hemidesmosome and focal adhesion formation, or facilitating the perpendicular collagen attachment towards the transmucosal region surface. In terms of modifying the transmucosal region surface, the variables include surface chemistry, surface topography and surface energy [Fig. 5], with the last property measured by solid-liquid contact angle and reflected by surface wettability [56].Fig. 5 Illustration showing the surface modification strategies to promote soft tissue attachment, which mainly include improving surface wettability, manipulating surface topography, altering surface chemical composition, and constructing biomimetic coatings. UV ultraviolet, Li lithium, Mg magnesium, Zn zinc, Ta tantalum, Ca calcium, ECM extracellular matrix.

Fig. 5

5.1 Improving surface wettability

Surface wettability may influence the biological processes in terms of the adhesion of proteins and macromolecules on the surface, cell adhesion, as well as bacterial adhesion and subsequent biofilm formation [56]. Previously, a possible mechanism explaining how surface wettability affects cell adhesion was proposed, in which hydrophobic surface was considered undesired since there might be air bubbles entrapped in the cell-surface interface hampering the cell adhesion [56]. Therefore, plenty of efforts were made to make the surface more hydrophilic. To date, a modified sandblasted, large-grit, acid-etched technique (Straumann® SLActive®) can produce a rough and hydrophilic surface, compared with the rough and hydrophobic surface produced by traditional sandblasted, large-grit, acid-etched technique (SLA), and has been applied in clinical practice. Generally, the modified SLA surface is obtained through a standard SLA treatment and subsequent rinsing under nitrogen protection and storing in isotonic saline, again with nitrogen filling instead of exposing to ambient atmosphere [57]. In a recent study, machined surface (roughness average Ra = 0.07 μm, contact angle = 93.6°), SLA surface (Ra = 1.56 μm, contact angle = 126.7°) and SLActive® surface (Ra = 1.43 μm, contact angle = 0°) of Ti were tested in terms of blood coagulation, human gingival epithelial cell adhesion and human gingival fibroblast (HGF) adhesion, demonstrating that SLActive® surface accelerated the fibrinogen network formation and blood coagulation, human gingival fibroblast adhesion as well as epithelial attachment [58]. Nevertheless, the SLActive® technique builds a hydrophilic surface at the cost of increasing surface roughness [59], which could facilitate bacterial colonization and induce infection risk. Efforts should be made to improve wettability while maintaining the original surface topography.

5.1.1 Ultraviolet functionalization

The ultraviolet (UV) ray covers a wavelength under 400 nm in the optical spectrum, including UV-A (wavelength λ = 320∼400 nm), UV-B (λ = 280∼320 nm), UV-C (λ = 200∼280 nm) and vacuum UV (VUV, λ < 200 nm) which cannot spread in the atmosphere. Nowadays, UV photofunctionalization has been introduced as an effective pretreatment to dramatically improve the surface hydrophilicity of the Ti and Y-TZP substrates without altering the surface roughness or other morphological features [60,61]. Accordingly, the excellent wettability performance after UV treatment resulted in an enhanced expression of hemidesmosome-related proteins in epithelial cells, and focal adhesion-related proteins in fibroblasts [[62], [63], [64]]. In a rodent animal model, the increased amount of perpendicularly oriented collagen fibers attached to the UV-treated Y-TZP implant was also depicted, which indicated an enhanced soft tissue seal [61].

In essence, the increased wettability of Ti and Y-TZP substrates after UV photofunctionalization can be attributed to the semiconductor photocatalyst nature of the superficial Ti dioxide (TiO2) or zirconia (ZrO2). Under UV irradiation, the electron on the valence band of TiO2 or ZrO2 would be excited to the conduction band (producing ecb−) when the absorbed energy was big enough to help the electron leap across the bandgap (3.2 eV in TiO2 and 5.0 eV in ZrO2 [65]), leaving a positively charged hole on the valence band (producing hvb+), thus forming an electron-hole (ecb−-hvb+) pair. If the ecb−-hvb+ pair is not recombined, the ecb− would be able to reduce the electron acceptors absorbed on the material surface, for example, reducing the absorbed oxygen and water into superoxide anion (O2−), hydrogen peroxide (H2O2) and hydroxyl radical (·OH), thus generating reactive oxygen species (ROS) [66]. The generated ROS would be able to obliterate the hydrocarbon debris and kill bacteria on the surface. Also, in the reduction process, hydroxy groups (Ti-OH or Zr-OH) would be increased, which are quite essential in improving surface hydrophilicity [63]. Meanwhile, the hvb+ would endow the surface with positive charges, thereby attracting the electronegative proteins and cells to adhere [67].

However, PEEK does not possess the photocatalytic capacity like Ti and Y-TZP, yet interestingly it was demonstrated that UV treatment could still promote the protein expressions of laminin 332 and integrin β4 in human gingival epithelial cells on PEEK surface even without an obvious improvement in surface wettability [68]. Due to the inertness of PEEK, plenty of efforts have been made on surface activation of the material. For instance, a 200 nm thick nano strontium titanate (SrTiO3) coating was formed on PEEK disk via magnetron sputtering, combining the osteogenic activity of strontium and the photocatalytic property of titanate, successfully resulting in the improved adherent behavior of osteoblasts, fibroblasts as well as epithelial cells after UV functionalization [69].

5.1.2 Non-thermal plasma treatment

Plasma, is described as partially ionized gas composed of electrons, free radicals, photons, ions, neutral molecules and atoms, yet being electrically neutral [70]. Generated by applying energy (usually in the form of electric current) to gas [71], plasma can be categorized into thermal plasma and non-thermal plasma (NTP, also referred to as cold plasma) depending on the electron density. The temperature of NTP is usually low and acceptable towards live tissues, leading to its extensive applications in modifying biomaterials, especially thermolabile ones.

NTP treatment resembles UV photofunctionalization in terms of the capacity to transform hydrophobic Ti/Y-TZP/PEEK surfaces into hydrophilic ones, thus improving the adhesion and proliferation of epithelial cells and fibroblasts [68,72,73]. Studies indicated that NTP treatment was capable of drastically improve the surface wettability without altering the topography, and boosting the attachment of connective tissue to the Ti implant, promoting the maturation of collagen, and encouraging the expression of fibronectin receptor integrin α5 [Fig. 6a∼f] [72]. Moreover, in another study where rats received implant surgeries 6 weeks after the tooth extraction, it was found that the NTP-treated Y-TZP abutments promoted the synthesis and secretion of laminin 332 in peri-implant junctional epithelium, though the length of peri-implant junctional epithelium seal was not examined [21]. In addition, on a PEEK substrate, the application of NTP to the surface was shown to enhance the expression of integrin β4 and laminin 332 in epithelial cells [68]. Nonetheless, the influence of NTP on PEEK surface distinguished from that on Ti or Y-TZP surface, which was reflected through the changed surface morphology after NTP treatment, where intact and barely changed Ti/Y-TZP surfaces compared with the roughened PEEK surface were seen [21,74].Fig. 6 Illustration showing that NTP treatment could drastically increase surface wettability, leading to improved epithelial sealing and connective tissue attachment. (a) Photographs of Ti disk (Control) and non-thermal plasma-treated Ti disk (NTP). (b) Scanning electron microscopy images of the surfaces. (c) Water contact angles of the surfaces. (d) H&E staining of peri-implant mucosa. Black arrowheads indicated the lowest point of the epithelium. (e) Picrosirius Red staining of collagen in peri-implant connective tissue. (f) Immunohistochemistry of integrin α5 (indicating fibronectin-fibroblast interaction) during wound healing of peri-implant connective tissue. Adapted with permission from Ref. [72]. Copyright 2024 Atsuro Harada et al.

Fig. 6

The improvement on surface hydrophilicity by NTP treatment can be explained by the increase of surface hydroxyl groups. Examined via X-ray photoelectron spectroscopy (XPS), a similar decrease of C/O ratio was observed on Ti/Y-TZP/PEEK surface, indicating the removal of hydrocarbon compounds and the production of oxygen-containing groups like hydroxyl groups, thus improving the surface energy and hydrophilicity [74,75]. Furthermore, potential molecular mechanisms of promoting tissue cell adhesion by NTP treatment were proposed, where NTP treatment was applied on Y-TZP and tested the biological response of HGFs, resulting in an enhanced spread morphology of cells after 2 hours of incubation [76]. In the meantime, upregulated calcium ion (Ca2+) signals inside the fibroblasts were recorded after receiving adenosine triphosphate (ATP) stimulation. This was further explained by increased expression of ATP receptors in fibroblasts on the plasma-treated surfaces, which led to augmented level of inositol triphosphate (IP3), and induced the Ca2+ transfer from the endoplasmic reticulum to cytoplasmic matrix, finally resulted the stretching of actin cytoskeleton and the strengthened fibroblast adhesion.

As mentioned above, UV photofunctionalization and NTP treatment are both effective methods for increasing surface wettability. The advantages of UV photofunctionalization include its topography-protective property and convenience, since the treatment can be conducted at chairside right before the insertion of the implant. Nevertheless, the UV treatment can be time-consuming. Conventional protocols adopt an UV exposure time ranging from 12 minutes to 24 hours, according to the individual wavelength and intensity applied [77,78]. As such, a high-energy VUV (vacuum UV, λ < 200 nm) functionalization was applied on Y-TZP to shorten the exposure time to 1 min, successfully resulting in an advanced fibroblast behavior, however the safety consideration of VUV application remaining to be further discussed [64]. On the other hand, NTP treatment also needs to be done soon before implantation, since the cleaned surface could be again occupied by hydrocarbon contaminants after hours of atmospheric exposure, causing a rebound in surface hydrophobicityc [68].

5.2 Manipulating surface topography

Surface topography is an important feature that impacts cell behaviors. Physiologically, the cells are embraced by highly dynamic ECM, whose topography features are constantly changing and providing spatiotemporal mechanotransduction cues to regulate cell adhesion, spreading, migration, proliferation and differentiation [79]. In specific, controlled topographies (from minimal irregular roughness to periodic micro/nano textures) can adjust tissue cell adhesion and collagen attachment. Among others, the irregular roughness is always characterized by Ra. However, mere Ra values seemed unable to predict tissue cell adhesion. In one study, the machined Ti surface with a Ra value of 0.069 μm exhibited inferior HGF attachment compared with the sandblasted surface (Ra = 1.514 μm) [80], whereas in another study, similar HGF attachment was observed on electrochemical polished (Ra ≈ 0.100 μm) and laser cut (Ra ≈ 1.000 μm) Ti surface [81]. In either study, there was an obvious increase in surface roughness, yet one study showed improved HGF adhesion while another did not. Therefore, similar Ra values did not necessarily indicate identical topographies, nor cell behaviors. As such, the following sections focused on the discussion of micro/nano textures.

5.2.1 Laser structuring: micro-textures

Laser technology can bring microscale textures to Ti/Y-TZP/PEEK surfaces, mainly including microgrooves and micropores. Microgrooves (36∼88 μm wide and 12∼137 μm deep) were fabricated on Y-TZP substrates via neodymiun-doped yttrium aluminium garnet (Nd:YAG) laser with an output power 6 W, a spot size of 3 μm and a pulse width of approximately 35 ns, which was able to increase the mechanical binding strength between the surface and the artificial gingiva (made of poly-vinyl siloxane) [82]. On the other hand, the HGFs seeded on Ti and PEEK surfaces with laser-produced microgrooves exhibited more elongated cells along with more cytoplasmic pseudopodia, signifying an enhanced fibroblast affinity on laser-treated surfaces [83]. As for epithelial cells, an in vitro study demonstrated an upregulated expression of hemidesmosome-related proteins including integrin α6β4 and laminin 332 of cells incubated on Y-TZP surfaces, which received Nd:YAG laser treatment and presented microgrooves texture, resulting in an enhanced stretching of cytoskeleton [84]. In addition, it was proved via a dog model that the Ti abutment with microgrooves texture could induce the development of perpendicularly oriented collagen fibers and decrease the epithelial downgrowth around dental implants [85].

Micropores texture is another popular surface topography produced by laser treatment. Periodic micropores with specific diameter and depth could be obtained through various laser processing systems, while micropores with specific diameters could improve the attachment of peri-implant mucosa. In a human study on how Ti alloys healing abutments with different topography influence soft tissue attachment, a micropores texture with the diameter of 5 μm was fabricated via laser treatment on the surface of the Ti alloys abutment, and a downregulation of pro-inflammation cytokines including tumor necrosis factor α (TNF-α) and the matrix metalloproteinase 9 (MMP-9) and an upregulation of anti-inflammation cytokines like interleukin 10 (IL-10) and tissue inhibitor of metalloproteinase 1 (TIMP-1) were detected in peri-implant soft tissue biopsies, compared with the tissue around smooth surfaces, which predicted a better soft tissue seal [86]. In another study, a micropores texture with the diameter of 7 μm or 15 μm that was created on Ti substrates by applying femtosecond laser technology, which could induce vertical migration of HGFs into the micropores (with the migration depth of more than 10 μm) [87]. In addition, micropores with diameter of 10 μm, 60 μm or 180 μm were also fabricated on Y-TZP disks via laser treatment with further NTP activation [73]. After incubation with the specimens for 24 hours, HGFs were observed to proliferate and spread better on smooth surfaces or textured surfaces with the micropore diameter of 10 μm. Meanwhile, micropores with the diameter of 60 μm or 180 μm were identified as unsuitable topographical cues for encouraging fibroblast spreading and might even contribute to cell apoptosis. Such relation between the increasing diameter of laser-produced micropores and the decreasing surface adhesion of fibroblasts may be explained in terms of cell size. In specific, fibroblasts are typically 15 μm∼25 μm in size, being able to modify their shapes to adapt to the surface topography [88]. However, such size adaption capacity is with limits, therefore fibroblasts would be unable to adapt to surfaces with excessively large micropores.

5.2.2 Electrochemical anodization: nano-textures

While the laser technology can introduce microscale topography onto the implant surface, electrochemical anodization has emerged as a facile and low-cost way to fabricate nanostructures, yet only applicable to Ti substrates. According to the articles reviewed, the conditions, produced nanostructures of electrochemical anodization, and the response of soft tissue cells are displayed in Table 2. Usually, organic-based electrolyte containing fluoride ion is selected to form TiO2 nanostructures rather than a compact TiO2 layer, and further facilitate the evolution from nanopores into nanotubes via dissolution of the interpore TiO2 [89].Table 2 Fabrication conditions and results of electrochemical anodization.

Table 2Electrolytes	Electrolytic conditions	Additional treatments	Textures	Diameters (nm)	Influence on tissue cells	Ref.	
Ethylene glycol (0.1 M NH4F, 8 M H2O)	35 V × 30 min, 35 V × 10 min	A tape peeling between two sections of anodization	Nanonets	77.7 × 47.4	HGFs: 30 min cell adhesion (−); 14 d collagen deposition (↑)	[90]	
Ethylene glycol (0.1 M NH4F, 1 M H2O)	60 V × 30 min, 60 V × 10 min	A tape peeling between two sections of anodization	Nanopores	52.9	HGFs: 48 h metabolic activity (↓)	[90]	
Ethylene glycol (0.11 M NH4F, 1.1 M H2O)	30 V × 3 h	Post sonication: 30 % H2O2× 5 min, 0.1 M CH3COOH × 60 min	Nanotubes	67.0	ECs: 3 h cell adhesion (−)
HGFs: 24 h cell adhesion (−); 30 min,1 h, 3 h TGF-β secretion (↑)	[91,92]	
Ethylene glycol (0.08 M NH4F, 0.55 M H2O)	60 V × 10 min	–	Nanopores	60	HGFs: 24h cell spread (↑); 7 d collagen deposition and ITGB1 mRNA expression (↑)	[93]	
Ethylene glycol (0.15 M NH4F, 5.5 M H2O)	10 V × 3 h/30 V × 3 h/60 V × 3 h	Prior acid cleaning: 2 mol/L HF-5 mol/L HNO3 mixture; post annealing: 450 °C × 2 h, air	Nanotubes	30/100/200	HGFs: 2 h adhesion (↑/ ↑/ ↓);	[94]	
Glycerol/water (50:50) with 0.27 M NH4F	30 V × 4 h	Post annealing: 450 °C × 1 h, air	Nanotubes	100∼170	HGFs: 4 h cell adhesion (−)	[95]	
Ethylene glycol (0.15 M NH4F, 5.5 M H2O)	50 V × 15 min	Post annealing: 500 °C × 2 h, air; post hydrogenation:
500 °C × 4 h, H2	Nanotubes	100	HGFs: 4 h cell spreading (↑); 24 h FAK, ITGA2, ITGB1, VCL, FN, COL1A1 mRNA expression (↑)	[96]	
Ethylene glycol (0.15 M NH4F, 5.5 M H2O)	50 V × 15 min	Post annealing: 550 °C × 2 h, air	Nanotubes	100	HGFs: 4 h cell adhesion (↑); 24 h FAK, ITGB1, VCL, MAPK3, ERK2, FN, COL1, ITGB1 mRNA expression (↑)	[97]	
Ethylene glycol/water (70:30) with 0.07 M NH4F	30 V × 2 h/40 V × 2 h/50 V × 2 h	Prior acid cleaning: 1 mol/L HF-1 mol/L HNO3 mixture	Nanotubes	61.4/103.4/45∼150	HGFs: 24 h cell spreading (↑/ ↑/ ↓); 12 h ITGB1, ITGB4, ICAM1, COL1 mRNA expression (↑)	[98]	
Electrolyte part of the concentrations of NH4F and H2O were converted from wt%, vol%, w/v% and v/v% units, please check the original article for accurate concentrations, HGFs human gingival fibroblasts, ECs epithelial cells, TGF-β transforming growth factor β, ITGB1 integrin β1, FAK focal adhesion kinase, ITGA2 integrin α2, VCL vinculin, FN fibronectin, COL1A1 collagen I α1 chain, MAPK3 mitogen-activated protein kinase 3, ERK2 extracellular regulated protein kinase 2, COL1 collagen I, ITGB4 integrin β4, ICAM1 intercellular adhesion molecule 1, − unchanged level, ↑ elevated level, ↓ decreased level.

Based on different electrolytes, a series of pre-experiments should be done to determine the optimal combination of voltage and time. Generally, lower voltages produce nanostructures with smaller diameters, while higher voltages resulted in larger diameters [91]. On the other hand, as time prolonged, the nanopores would transform into nanotubes with the dissolution of interpore TiO2, and the nanotubes keep extending in length until the peeling of the brittle tube layer occurs [98]. On this account, manual tape peeling or acid sonication were adopted to wipe off the redundantly long and mechanically unstable nanotubes, forming organized nanopores or short-length nanotubes [90].

It can be concluded from Table 2 that nanopores or nanotubes with the diameter of 50∼100 nm were proved to promote the early adhesion and focal adhesion-related genes expression of HGFs, probably because of the nanoscale ordered topographical cues that electrochemical anodization produced. Meanwhile, when the mean diameter of nanotubes is around 200 nm, the surfaces could inhibit the adhesion and proliferation of HGFs [94]. Nevertheless, the investigations of nanopores and nanotubes on epithelial cells were scarce, with in vivo test revealing no significant changes of peri-implant junctional epithelium around abutment surfaces with nanotubes [99].

TiO2 nanotubes and nanopores could not only act as mechano-transduction cues that might induce an enhanced focal adhesion formation in fibroblasts, but are also widely adopted as a pretreatment that would receive further modifications. The as-obtained nanotubes was further loaded with gold nanoparticles (AuNPs), endowing the surfaces with the ability to respond to visible light rather than UV light, and resulting in the inactivation of multispecies biofilm via ROS generation, while enhancing the early attachment of HGFs [95]. In addition, collagen I fibers were perpendicularly inserted into TiO2 nanotubes via electrophoretic fusion technique, obtaining a surface that encouraged both adhesion of platelets and HGFs, which induced a higher level of platelet-derived growth factors that would aid in wound healing [92]. Rather than loading TiO2 nanotubes/nanopores with other biomaterials, TiO2 nanopores were creatively converted into Ti nanopores through magnesiothermic reduction, improving the inferior electrical conductivity of TiO2, and confirmed that the electrical stimuli could facilitate the attachment of HGFs while exerting antibacterial effect against oral salivary biofilm [100].

5.3 Altering surface chemical composition

Introducing metallic ions onto the surfaces could change the surface chemical composition of the material and influence tissue cell adhesion, where non-metallic elements including fluoride (F) and phosphorus (P) were also demonstrated to impose positive effect on the transmucosal soft tissue seal [101,102], though with limited investigations. As presented in Table 3, lithium ion (Li+), magnesium ion (Mg2+), Ca2+, zinc ion (Zn2+) and tantalum ion (Ta5+) could be modified on Ti/Y-TZP surfaces (PEEK surface modifications based on this rationale are limited) through either traditional hydrothermal treatment and sol-gel synthesis, or novel techniques including plasma immersion ion implantation and magnetron sputtering. With the above-mentioned metals doped, the surfaces were proved to play a promotional effect on soft tissue cell adhesion and capable of achieving in an enhanced soft tissue seal. Mechanisms under such promotional phenomena of Ca2+, Mg2+ and Zn2+ doping were discussed further in the following paragraphs, while Li + or Ta5+ coatings were not explained in detail yet due to the limited investigations on their influence on soft tissue regeneration.Table 3 Metallic ions-releasing coatings that also enhance soft tissue cell adhesion.

Table 3Released metallic ions	Substrates (pretreatments)	Fabrications	In vitro results	In vivo results	Ref.	
Li+	Ti (mixture acid washing; alkaili-heat treatment)	Hydrothermal treatment	HGFs: 24 h adhesion and spreading (↑)	–	[103]	
Mg2+	Ti	Plasma immersion ion implantation	HGFs: 4 h spreading and adhesion (↑); 12 h vinculin and fibronectin expression (↑); p-Akt/Akt level (↑)	–	[104]	
Mg2+	Ti (mixture acid washing)	Plasma immersion ion implantation	HGFs: 3 d collagen I and fibronectin secretion (↑)	–	[105]	
Mg2+ and Ca2+	Y-TZP powder (pre-sintering)	Sol-gel synthesis	HGFs: 5 h adhesion (↑); 24 h mRNA expression of ITGA5 and ITGB1 (↑); 3 d mRNA expression of MAPK3, SMAD2, ERK2, SMAD3, COL1 (↑)	–	[106]	
Ca2+	Ti	Hydrothermal treatment	Rat OECs: 4 d laminin 332 secretion (↑)	Male Wistar rats: right maxillary first molar implantation, 4 w coronal laminin 332 secretion of PIE (↑); 4 w apical HRP penetration (↓)	[107,108]	
Ca2+	Ti	Hydrothermal treatment	HGFs 1 h, 8 h, 12 h adhesion (↑); 3 d integrin β1 expression (↑)	Sprague-Dawley rats: right maxillary first molar implantation, apical HRP penetration (↓)	[109]	
Zn2+	Ti	Plasma immersion ion implantation	HGFs: 6 h, 12 h adhesion (↑); 3 d collagen I and fibronectin secretion (↑)	–	[110]	
Zn2+	Y-TZP	Hydrothermal treatment	Human OECs: 3 d integrin β4 expression (↑); 3 d PI3K, Akt, ITGB4 mRNA expression (↑)	Male Wistar rats: right maxillary first molar implantation, 2 w apical HRP penetration (↓)	[111]	
Ta5+	Ti (sputter-cleaning)	Magnetron sputtering	HGFs: 1 h adhesion (↑); 24 h VCL, ITGB1, ITGA5, COL1A1, FN mRNA expression (↑)	–	[112]	
Ti titanium, Y-TZP yttria-stabilized zirconia tetragonal polycrystals, HGFs human gingival fibroblasts, p-Akt phosphorylated Serine/Threonine kinase, Akt Serine/Threonine kinase, ITGA5 integrin α5, ITGB1 integrin β1, MAPK3 mitogen-activated protein kinase 3, SMAD2 SMAD family member 2, ERK2 extracellular regulated protein kinase 2, SMAD3 SMAD family member 3, COL1 collagen I, OECs oral epithelial cells, PI3K phosphatidylinositol-3-kinase, ITGB4 integrin β4, VCL vinculin, COL1A1 collagen I α1 chain, FN fibronectin, PIE peri-implant junctional epithelium, HRP horseradish peroxidase, ↑ elevated level, ↓ decreased level.

5.3.1 Ca2+ incorporation

As the critical component of the most investigated hydroxyapatite, Ca2+ plays a positive role in promote peri-implant soft tissue seal. One of the assumptions that explained how Ca2+ promoted cell adhesion was proposed in terms of ECM molecules that had Ca2+-binding ability. For example, heparan sulfate proteoglycan (HSPG) is an anion matrix proteoglycan [113]. It was proposed that HSPG exhibited high affinity with both Ca2+ and laminin 332, which enabled it to bridge the gap between the Ca2+-releasing Ti surface and the epithelial cell with integrin α6β4 on its membrane [114]. Briefly, the Ca2+-HSPG-laminin 332 bonding would activate the integrin α6β4 of the epithelial cell and drive the formation of hemidesmosomes. Similarly, the fibrin, an adhesive glycoprotein that has Ca2+-binding property in the ECM, facilitating the interaction between the Ca2+-releasing surface and fibronectin, thus initiating the formation of focal adhesion in the fibroblast [Fig. 7a] [114].Fig. 7 Illustration showing the mechanisms of how Ca2+, Mg2+ and Zn2+ aid in the adhesion of soft tissue cells. (a) Schematic diagram showing the hypothesized mechanism of cell adhesion to the surface of Ca2+-modified Ti substrate. Adapted with permission from Ref. [114]. Copyright 2012 Hideyuki Okawachi et al. (b) Illustration for the possible action of Mg2+-modified surfaces to HGFs. Adapted with permission from Ref. [104]. Copyright 2019 Wiley Periodicals, Inc. (c) Illustration for the possible action of Mg2+-modified surfaces to HGFs. (d) Illustration for the possible signal pathways of the effect of Mg2+/Zn2+ on the behaviors of HGFs. Adapted with permission from Ref. [105]. Copyright 2020 Lanyu Wang et al.

Fig. 7

5.3.2 Mg2+ incorporation

Two of the articles in Table 3 indicated that the relative mRNA expression level of integrin β1, FAK, collagen I, fibronectin and vinculin were elevated in fibroblasts seeded on Mg2+-modified Ti surfaces [104,105]. In addition, increased phosphorylation of Akt, which could be blocked by PI3K inhibitor LY294002, was observed in fibroblasts on the coated surfaces, indicating the activation of PI3K/Akt pathway on Mg2+-doped surfaces [104]. The whole surface-fibroblast interaction can be illustrated through a mechanism as below: clustering of integrins, especially integrin β1, phosphorylation of FAK, the recruitment of PI3K, the activation of PI3K/Akt signaling transduction, the augmented synthesis and secretion of collagen I and fibronectin, and stimulated expression of cytoskeleton-related proteins, finally increasing the formation of the adhesive apparatus focal adhesion [Fig. 7b∼c] [104]. However, how Mg2+ interferes with this complicated pathway required more exploration.

5.3.3 Zn2+ incorporation

It is demonstrated that the incorporation of Zn2+ could be accomplished via both hydrothermal treatment and plasma immersion ion implantation, thereby endowing the Ti surface with superior affinity with epithelial cells and fibroblasts [105,108,110]. In addition, zinc oxide (ZnO) nanorod crystals (releasing Zn2+ into the surrounding environment) were also successfully fabricated on the Y-TZP abutment through hydrothermal treatment, and were verified that the coating could tighten the epithelial attachment by promoting the binding of laminin 332 and integrin β4 and upregulated the PI3K/Akt pathway [111]. On the other hand, zinc as an essential trace element with considerable abundance in the connective tissue, was proposed to participate in regulating fibroblast adhesion via SLC39A (also referred to as ZIP) and transforming growth factor β (TGF-β)/SMAD signaling pathway [Fig. 7d] [105,115].

5.4 Constructing biomimetic coatings

The soft tissue seal can be interpreted as the result of cell-ECM interaction. As mentioned above, peri-implant ECM is composed of proteoglycans (consisting of glycosaminoglycans and the core protein) scaffolds, bulk collagen fibers and adhesive glycoproteins (mainly laminin 332 and fibronectin). Therefore, various surface modifications based on these molecules were fabricated to mimic ECM and prosper tissue cell adhesion, thus achieving peri-implant soft tissue seal.

5.4.1 ECM protein coatings

ECM proteins include dispersive adhesive glycoproteins and bulk collagens. Among others, laminin 332 and fibronectin are important adhesive glycoproteins since their essential roles in mediating the formations of attaching apparatus. Laminin 332 is the initiator of hemidesmosome formation, which is critical for epithelial attachment [27]. Modifying the surface with laminin 332 could be realized through coatings of laminin 332 protein and the functional LamLG3 peptide [[116], [117], [118]], or coatings of laminin subunit α3 (LAMA3) gene, which editing the α3 chain of laminin 332 [119,120]. Among them, LAMA3 gene coatings advantage over protein/peptide coatings in achieving stability and lasting protein release, since the activation of fibrinolytic system might cause the degradation of laminin 332 protein [121], while the cell uptake of genetic vectors warranted the continuous synthesis of laminin 332. Specifically, LAMA3 gene coatings could be fabricated via genetic vectors like plasmid and recombinant adenovirus [Fig. 8a] [119,120]. After a successful laminin 332 modification, the surface could boost the proliferation and of epithelial cells, driving the epithelial differentiation of gingival mesenchymal cells, and promoting the formation of hemidesmosomes at the upper 1/3 of peri-implant junctional epithelium [120], which would effectively improve the sealing capacity of peri-implant junctional epithelium.Fig. 8 Illustration showing that ECM protein coatings could facilitate the adhesion of epithelial cells and fibroblasts. ECM extracellualr matrix (a) Surface modification via plasmid-mediated pLAMA3-CM gene transfection promoted the attachment of gingival epithelial cells to Ti and improved soft tissue seal at the transmucosal region. CS chitosan, COL collagen, pLAMA3-CM plasmid encoding a motif of the C-terminal globular domain of laminin α3 chain, S-Ti smooth titanium, SEM scanning electron microscope, HRP horseradish peroxidase, B new bone formation, E epithelial tissue, C connective tissue, BV blood vessel. Adapted with permission from Ref. [119]. Copyright 2019 The Royal Society of Chemistry. (b) Schematic illustration of the silanization-mediated fibronectin modification process, and confocal microscopy images of HGFs 24 h after cultivation on fibronectin modified (left panel) and pure (right panel) surfaces. Immunostaining indicated pFAK-Y397 (red) and phalloidin-staining (green) of HGFs seeded on specimens. White arrows labeled pFAK-Y397 expression at the end of actin filaments. Adapted with permission from Ref. [122]. Copyright 2021 Alena L. Palkowitz et al. (c) PDA-mediated RGD functionalization process and its effect on HGFs and bacterial adhesion for enhanced peri-implant soft tissue seal. (d) Confocal laser scanning microscopy images of HGFs on pristine zirconia (upper panel) and RGD-functionalized zirconia surfaces after 3 h (left column) and 24 h (right column) of culture. Adapted with permission from Ref. [125]. Copyright 2020 The Royal Society of Chemistry.

Fig. 8

Fibronectin participates in the ECM-fibroblast adhesion via focal adhesion, where integrin α5β1 on fibroblast membrane as the fibronectin receptor. Fibronectin also plays a critical role in connecting different proteoglycans in the ECM. For example, a fibronectin coating was fabricated on surfaces of Ti alloys and Y-TZP, where an elevated Tyr 397 autophosphorylation of FAK at the edge of fibroblasts was detected, indicating increased focal adhesion formation around fibroblasts that were incubated on the coated surfaces [Fig. 8b] [122]. Since fibronectin circulates in blood in a soluble inactive form, and functions in tissues as insoluble fibrils when activated by cell surfaces, imprudently introducing fibronectin-coated surface into the bleeding wound might cause the accident formation of insoluble emboli in circulation, which can be considerably dangerous [37]. This led to the identification of the functional peptides in fibronectin. Fibronectin has an Arg-Gly-Asp (RGD) motif, which is recognized as a cell-adhesive sequence through binding to integrin α5β1 and is widely used in promoting HGF attachment to the surfaces of dental implant materials [123,124] A linear KGG-RGD-SP peptide coating and cyclic c(RGD-fK) peptide coating were produced on Y-TZP surfaces through polydopamine (PDA)-assisted immobilization of peptides, with both coating being confirmed to enhance early adhesion and spreading of HGFs [Fig. 8c∼d] [125]. In addition, the heparin binding II (HBII) domain in fibronectin has an affinity to the TGF-β, which is a key growth factor in wound healing. Accordingly, a coating of RGD-mutated HBII fragment was immobilized on Ti surfaces, together resulting in promoted fibroblasts adhesion and TGF-β adsorption [126].

Collagen, usually collagen I, is the main component in forming connective tissue attachment around a natural tooth, where the collagen fibers directly protrude in the cementum as Sharpey's fibers, and bind to fibroblasts via integrin α2β1 [38]. Inspired by natural teeth, collagen I was coated on surfaces through 1-(3-dimethylaminopropyl)-3-ethylcarbodiimide (EDC)/N-hydroxyl succinimide (NHS) chemistry and layer-by-layer assembly, through which an increased fibronectin adsorption and a boosted the expression of integrin β1 and vinculin in HGFs were achieved [127,128]. More importantly, perpendicularly protruding collagen I is the core of soft tissue seal. Perpendicularly arranged collagen I could be fabricated through electrophoretic fusion technique, where the negatively charged collagen C terminal was perpendicularly navigated to anodic TiO2 nanotube in a polyacrylamide gel system [91]. Afterwards, an improved affinity to fibroblasts of the as-obtained surface was confirmed, which supported the promising future endogenous collagen deposition secreted by fibroblasts [92]. As expected, obvious vertical oriented collagen fibers were detected around the abutment with above-mentioned surface modification in a rat model 2 weeks post implantation, compared with circumferential collagen array around bare Ti abutment [128]. Taken together, collagen I fibers were coated on biomaterial surfaces, in order to induce the focal adhesion formation and the perpendicular extension of newly secreted endogenous collagen fibers, thus promoting the healing of peri-implant connective tissue.

5.4.2 Polysaccharides-inspired coatings

Inspired by the structure of polysaccharides in the ECM, chitosan and cellulose, which are scaffold polysaccharides with hydroxyl and/or amino functional groups, were coated on Ti surfaces to promote the spreading of HGFs. Specifically, chitosan was introduced onto the Ti surface through a soaking method, with the surface receiving an alkaili-heat pretreatment and thus rich in hydroxyl groups, resulting in increased cell spreading area [129]. Cellulose nanocrystals (CNC), on the other hand, were also spin-coated onto Ti surfaces after polyethylenimine modification [130]. The surface anisotropy provided by CNC exhibited guidance on the growth and alignment of HGFs, where dense cytoskeleton was found in the peripheral region of the cell, suggesting the formation of focal adhesion which indicated the cell adhesion of HGFs and further soft tissue seal at the connective tissue level. Except for polysaccharides, proteins were chosen to simulate polysaccharides morphology as well. In specific, Amyloid is a chemically stable protein with abundant β-sheet structure, which could be obtained by mixing cystine and lactoferrin and successfully immobilized onto the Ti surface via a simple solution incubation. The coating presented an impressive performance in promoting the formation of hemidesmosomes in epithelial cells and upregulating the expression of focal adhesion-related proteins in HGFs [131]. Moreover, the as-obtained amyloid coating exhibited little degradation in artificial saliva at 28th day, warranting the coated surface to exert a lasting positive effect on the soft tissue until the wound healing was completed.

While no related surface modifications on Y-TZP have been identified, one study proposed fabricating an ECM-mimic structure on PEEK surface. Since PEEK itself is a polymer with aldehyde groups that can offer potential carbon radicals, acrylic acid was used to link to the carbon radicals before self-polymerizing under UV activation to form poly-acrylic acid (PAA) chains, following by the ethylenediamine (EDA) covalently linking to the carboxyl groups in PAA through EDC/NHS chemistry, and resulting in a brush-like surface on PEEK with superhydrophilicity and upregulated FAK signaling pathway in myofibroblasts [132].

6 Modification strategies to balance the antibacterial effect and soft tissue attachment

According to the “race for the surface” theory, a competitive occupation of the biomaterial surface by tissue cells and bacteria occurs upon implantation [133]. If the race was won by bacteria and the surface was rapidly covered by biofilm, there would be less tissue cells adhering to the surface and the tissue integration would be hampered [134,135]. The formation of oral biofilm would increase the risk of developing peri-implant mucositis [136]. Without timely treatment, peri-implant mucositis would evolve into peri-implantitis and lead to final implant failure [137]. Therefore, antibacterial activity should be equipped with the surface to aid soft tissue seal.

Peri-implant diseases have polymicrobial etiology. The microbiological profiles of peri-implant conditions are predominated by Gram-positive cocci (Streptococcus species) and facultative anaerobic rods [138]. Specifically, tissues of peri-implant mucositis are associated with increased pathogens in red complex (Poryphyromonas gingivalis, Tannerella forsythia, and Treponema denticola) and orange complex (Fusobacterium nucleatum, Prevotella intermedia, Campylobacter rectus, and Parvimonas micra). Tissues of peri-implantitis are identified with a predominant colonization of red complex bacteria and unique species like Filifactor alocis and Fretibacterium fastidiosum. Meanwhile, opportunistic pathogens like Staphylococcus aureus, Escherichia coli and Candida albicans are also detected in tissues of peri-implantitis, which are perceived as contributing to disease progression.

Treatment modalities for peri-implantitis are always remedial. According to S3-level clinical practice guideline developed by European Federation of Periodontology, there are non-surgical and surgical approaches for peri-implantitis management [139]. Specifically, non-surgical treatments mainly depend on sub-marginal instrumentation (mechanical scaling, laser, air-polishing, etc.) and adjunctive use of local/systemic antimicrobials to control peri-implant biofilms and inflammation. Surgical treatments universally incorporate flap elevation, removal of inflamed tissues and implant surface debridement, with or without additional reconstructive approaches (bone grafts, bone replacement grafts, barrier membranes, etc.) to manage peri-implant osseous defects. As indicated, debridement and anti-infection therapy are the essence in the management of peri-implantitis.

Though the afore-mentioned treatments provide remedies for peri-implantitis, it is a laborious work and the therapeutic outcomes are hard to guarantee. Therefore, numerous studies have investigated the application of surface modifications to exert antibacterial effects while maintaining good biocompatibility with soft tissue [Table 4]. As shown in Fig. 9, the included strategies can be concluded as below: applying antimicrobial peptides, modifying with metallic elements, employing intrinsic antibacterial agents, and developing stimuli-responsive antibacterial surfaces, which were discussed in the following sections, respectively.Table 4 Antimicrobial strategies that also maintain soft tissue attachment.

Table 4Antimicrobial designs	Surface modifications	Tested bacteria	Influence on soft tissue	Ref.	
Antimicrobial proteins/peptides	GL13K-LamLG3 coimmobilization	S. gordonii	Promoting the proliferation and adhesion of human epithelial cells.	[117]	
Lactoferrin-derived amyloid coating	S. aureus
P. gingivalis	Promoting the proliferation and adhesion of both human epithelial cells and human gingival fibroblasts.	[131]	
Metallic elements	TaN-Ag coating	S. aureus	Promoting the proliferation of human gingival fibroblasts.	[140]	
AgNPs-loaded chitosan-heparin polyelectrolyte multilayers	P. gingivalis	Promoting the adhesion and proliferation of human gingival fibroblasts.	[129]	
Ag-doped TiO2 nanotubes	P. gingivalis
A. actinomycetemcomitnans	Promoting the gene expression of FN in epithelia-like cells and COL-I in fibroblast-like cells.	[141]	
Ag linear-beam ion implantation	S. mutans
P. gingivalis	Preserving the viability of human gingival fibroblasts.	[142]	
ZnO nanorods-nanospheres hierarchical structure	S. aureus
E. coli	Inhibiting the proliferation of human fibroblasts.	[143]	
ZnO@ZnS nanorod-array	S. aureus
E. coli	Reducing the cytotoxicity to human gingival fibroblasts.	[144]	
Ta2O5 coating	S. aureus
A. actinomycetemcomitnans	Promoting the proliferation of human skin fibroblasts.	[145]	
Ga-doped TiO2 nanopores	human saliva-derived biofilms	Favorably maintaining the viability of human gingival fibroblasts.	[146]	
Li doping	S. aureus human saliva-derived biofilms	Promoting the viability of human gingival fibroblasts.	[103]	
Intrinsic antibacterial agents	GO layers deposition	S. aureus
E. coli	Maintaining the viability of human gingival fibroblasts in a coculture model with bacteria.	[147]	
GO coating	P. gingivalis
F. nucleatum
S. mutans	Preserving the adhesion and proliferation of human gingival fibroblasts.	[148]	
PDA coating	S. gordonii
S. mutans	Promoting the adhesion, proliferation and collagen secretion of human gingival fibroblasts.	[149]	
PDA-RGD functionalization	S. mutans
P. gingivalis	Promoting the adhesion, proliferation of human gingival fibroblasts.	[125]	
Chitosan coating	S. aureus
E. coli	Preserving the viability of human fibroblasts.	[150]	
Chitosan-based hydrogel coating	S. aureus
E. coli	Preserving the viability of human fibroblasts.	[151]	
HEMA coating	S. aureus
E. coli	Preserving the viability of human foreskin fibroblasts.	[152]	
DEP coating	S. aureus
C. albicans	Preserving the viability of murine fibroblasts.	[153]	
PDA-HA-DMAHDM coating	S. aureus
S. sanguinis human saliva-derived biofilms	Preserving the viability of human gingival fibroblasts and periodontal ligament stem cells.	[154]	
Renewable N-halamine polymeric coating	S. aureus
P. gingivalis	Exerting no significant influence on soft tissue regeneration in a subcutaneous model.	[155]	
Stimuli-responsive platforms	PDT:
TiO2 nanotubes,
UV-C (100∼280 nm) for 24 h	P. gingivalis	Preserving the proliferation of human gingival fibroblasts.	[60]	
PDT:
AuNPs-loaded TiO2 nanotubes, blue light (420∼480 nm) for 10 min	S. gordonii
P. gingivalis
F. nucleatum	Promoting the adhesion and proliferation of human gingival fibroblasts.	[95]	
PTT:
PDA/HA-embedded gelatin coating,
NIR (808 nm) for 5 min	S. mutans	Preserving the viability of human gingival fibroblasts.	[156]	
PTT:
PDA-GO-PDA-mineralized collagen sandwich structure,
NIR (808 nm) for 10 min	S. sanguinis
F. nucleatum
P. gingivalis	Promoting the adhesion and proliferation of human gingival fibroblasts.	[127]	
PTT:
Ce-doped TiO2-x coating,
NIR (808 nm) for 10 min	S. aureus
E. coli	Preserving the proliferation of human gingival fibroblasts.	[157]	
EST:
Ti nanopores,
1.5 V for 5 min	human saliva-derived biofilms	Promoting the proliferation of human gingival fibroblast.	[100]	
EST: rechargeable HPI coating	S. aureus
E. coli	Promoting the adhesion of human gingival fibroblasts.	[158]	
SDT:
BaTiO3-x/L-arginine coating,
US irradiation (1 MHz) for 15 min	MRSA	Decreasing the soft tissue inflammation in an infection implantation model.	[159]	
SDT:
AuNPs-loaded TiO2 nanotubes,
US irradiation (1 MHz) for 5 min	P. gingivalis
F. nucleatum
S. sanguinis	Decreasing the soft tissue inflammation in a peri-implantitis model.	[160]	
LamLG3 the large globule domain 3 of the laminin 332 α3 chain, TaN tantalum nitride, Ag silver, TiO2 titania, AgNPs silver nanoparticles, ZnO zinc oxide, ZnS zinc sulfide, Zn zinc, CeO2 ceria, Ta2O5 tantalum oxide, Li lithium, Ga gallium, PDA polydopamine, GO graphene oxide, RGD Arg-Gly-Asp, HEMA 2-hydroxylethyl methacrylate, DEP diethyl phosphite, HA hydroxyapatite, DMAHDM dimethylaminohexadecyl methacrylate, PDT photodynamic therapy, UV-C ultraviolet C, AuNPs gold nanoparticles, PTT photothermal therapy, Ce cerium, TiO2-x defective titania, NIR near infrared, EST electrical stimulation therapy, Ti titanium, HPI hydrazine hydrate-derived polyimide, SDT sonodynamic therapy, BaTiO3-x oxygen-deficient barium titanate, US ultrasonic sound, FN fibronectin, COL-I collagen I, S. gordonii Streptococcus gordonii, S. aureus Staphylococcus aureus, P. gingivalis Porphyromonas gingivalis, A. actinomycetemcomitans Actinobacillus actinomycetemcomitnans, S. mutans Streptococcus mutans, E. coli Escherichia coli, F. nucleatum Fusobacterium nucleatum, C. albicans Candida albicans, S. sanguinis Streptococcus sanguinis.

Fig. 9 Illustration showing the surface modification strategies balancing antibacterial effect and soft tissue attachment, which include applying antimicrobial proteins/peptides, modifying with antibacterial metallic elements, employing intrinsic bacterial agents, and developing stimuli-responsive bactericidal surfaces. Ag silver, Zn zinc, Cu copper, Ta tantalum, Ga gallium, GO graphene oxide, PDT photodynamic therapy, PTT photothermal therapy, EST electrostimulation therapy, SDT sonodynamic therapy.

Fig. 9

6.1 Applying antimicrobial proteins/peptides to the surface

Antimicrobial proteins or peptides (AMPs) were introduced onto Ti surfaces to eliminate the colonized pathogens while encouraging soft tissue attachment. One of the action mechanism can be concluded as the electrostatic interaction between positively charged proteins/peptides and negatively charged bacterial cell membranes, which can cause disruption of membrane integrity and activation of autolytic enzymes, thereby preventing bacterial colonization and the subsequent infection [161]. Previously, an artificial GL13K peptide was designed and synthesized according to parotid secretory protein/short palate, lung and nasal epithelium clone 2 (PSP/SPLUNC2), and modified to carry increased amount of positive charge, thus performing bactericidal effect [162]. Based on its extraordinary antimicrobial activity, GL13K peptide was immobilized on Ti surfaces through alkaili-heat pretreatment and silanization-mediated coating method to decrease the viability of the adherent Streptococcus gordonii [117]. To balance the antibacterial and adhesion-promoting function, a laminin-derived peptide, LamLG3 was co-immobilized onto the surface with the antibacterial GL13K peptide, thus improving the adhesion of epithelial cells. As for cytocompatibility, the application of GL13K peptide showed no significant influence on epithelial cell proliferation, yet slightly compromised fibroblast metabolic activity.

Another action mechanism relies on the ability of the protein/peptide to interfere with the bacterial metabolism. For instance, lactoferrin is a natural protein with broad-spectrum antibacterial ability, due to its high affinity to iron, thus impeding the utilization of iron in bacteria [163]. The protein has further capacity to bind to the lipoteichoic acid in Gram-positive bacteria or the lipoprotein in Gram-negative bacteria, which might interfere with the metabolism of bacteria and lead to their declined viability. Recently a lactoferrin-derived amyloid was obtained which retained the antibacterial activities of lactoferrin, and coated onto Ti surfaces, leading to a bactericidal effect on Gram-positive S. aureus and Gram-negative P. gingivalis [Fig. 10 a∼d] [131]. Meanwhile, based on the abundant amino and carboxyl groups of the lactoferrin-derived amyloid, the proliferation and adhesion of epithelia cells and fibroblasts were promoted, since lactoferrin can bind to the integrins on cell membrane and promote cell proliferation through activating PI3K/Akt pathway, leading to enhanced soft tissue seal [131].Fig. 10 Illustration showing the antimicrobial activities of antimicrobial peptide, metallic element, and external stimuli-responsive platform. (a) Schematic of the synthesis of lactoferrin-derived amyloid coating on Ti (LAT). (b) Images of the bacterial colonies formed by S. aureus and P. gingivalis that adhered to the polished titanium control (PT) and LAT. (c) Antibacterial activity against S. aureus and P. gingivalis that were in contact with PT and LAT. Significant differences between PT and LAT were labeled with different letters (p < 0.01, Student's t-test, n = 6) (d) SEM images of S. aureus and P. gingivalis grown on the surfaces of PT and LAT. Adapted with permission from Ref. [131]. Copyright 2023 Wiley-VCH GmbH. (e) The preparation of AgNPs-loaded chitosan-hepatin polyelectrolyte multilayers (PEMs) on Ti substrate. (f) SEM images of P. gingivalis on Ti (left) and PEMs (right) samples. Adapted with permission from Ref. [129]. Copyright 2019 Springer Science Business Media, LLC. (g) Nanoporous Ti implants toward electrical stimulation therapy (EST). Schematic representation: (left to right) anodization of Ti implants to fabricate TiO2 nanopores (NPs) and magnesiothermic reduction to convert TiO2-NP to Ti-NP; and EST using nanopores toward soft-tissue integration and antibacterial efficacy. (h) Live/Dead staining of biofilms on Ti and EST-applied Ti-NP surfaces. EST parameters: 1.5 V for 5 min per day. Adapted with permission from Ref. [100]. Copyright 2024 Karan Gulati et al.

Fig. 10

6.2 Modifying the surface with antibacterial metallic elements

Various metallic elements and metal oxide could be utilized as alternative candidates to support soft tissue attachment and prevent bacterial infections. Silver (Ag) possesses superior bactericidal effect and broad-spectrum antibacterial property against various oral pathogenic microorganisms [[164], [165], [166]]. Ag nanoparticles (AgNPs) possess enhanced antibacterial efficiency due to their large total surface area and highly active surface for bacterial interaction [[167], [168], [169]]. In specific, AgNPs tend to accumulate on the bacterial membrane and cause damages to the bacterial membrane, while the oxidation of AgNPs to silver ion (Ag+) induce the generation of ROS, with Ag+ possessing a particularly high affinity to thiols and leading to irreversible aggregation of the thiol-containing molecules, together causing the inactivation of bacteria [170]. Therefore, AgNPs were coated onto the transmucosal region surface in various forms. For example, a composite coating of crystalline tantalum nitride (TaN) and AgNPs was deposited onto Ti surface through a twin-gun magnetron sputtering system, where TaN worked as a biocompatible structure and AgNPs exerted bactericidal effect [140]. Besides, AgNPs can be deposited onto Y-TZP surface via silver linear-beam ion implantation to inhibit the growth of Streptococcus mutans and P. gingivalis [142]. The antibacterial activity of AgNPs is a dose-dependent manner. To balance the adhesion-promoting and antibacterial function of AgNPs-based coating, plasma immersion ion implantation technique was employed to incorporate AgNPs onto TiO2 nanotubular surface, and the concentration and depth of the incorporated Ag could be tailored through changing the voltages [141]. Specifically, the surface obtained at lower voltage possess higher amount of superficial Ag, but the surface obtained at higher voltage shows deeper Ag penetration and better Ag leaching profile. Both surfaces remained the nanotubular morphology and exhibited satisfactory antibacterial activity against oral pathogens. But the biocompatibility of lower voltage surface is significantly compromised due to the large amount of accumulated superficial AgNPs. In addition to manipulate the surface content of AgNPs, controlling the release can also be recognized as an effective strategy to balance the two functions. For instance, AgNPs could be loaded onto Ti surface via chitosan-heparin polyelectrolyte layer-by-layer technique. The release of Ag+ remained in a sustained manner due to the multilayered coating structure, so that the surface exhibited non-cytotoxic effect and facilitated the adhesion, proliferation and stretch of HGFs, while maintaining long term antibacterial efficiency up to 28 days to fully cover the early dental implant soft tissue seal period [Fig. 10 e∼f] [129].

ZnO nanorods represented as another candidate to construct antibacterial surface, the main action mechanism could be ascribed to the disruption to the bacterial cell membrane and the release of H2O2 from ZnO [143]. A dual antibacterial effect was realized by fabricating a layer of ZnO nanospheres outside the ZnO nanorods, since the nanospheres showed rapid release of ZnO after sonication, while the nanorods showed sustained release of ZnO [143]. However, as-modified Ti surfaces exhibited minor cytotoxicity to HGFs. To further enhance the coating stability and cytocompatibility, a thin layer of zinc sulfide (ZnS) was deposited onto ZnO nanorod arrays via immersion of ZnO into thioacetamide aqueous solution to allow in situ deposition. The obtained ZnO@ZnS nanorod arrays realized the bactericidal effect through H2O2 production during the decomposition of ZnO, meanwhile the nanorods topography was considered to cause bacteria membrane damages as well, where the excessive coating of ZnS was designed to prolong the release of Zn2+ and reduce the toxicity towards tissue cells [144,171]. Moreover, due to the protective effect of ZnS, the release of Zn2+ was gentler than that of uncoated one, leading to an enhanced fibroblast adhesion and wound healing process.

Tantalum (Ta) has received increased attention for dental applications due to its corrosion resistance and excellent biocompatibility [172]. Abundant researches have demonstrated that Ta exhibited negligible cytotoxicity and would not inhibit local cell growth [173,174]. By applying magnetron sputtering technique with or without rapid thermal annealing treatment, amorphous and crystalline tantalum oxide (Ta2O5) coatings were deposited on Ti surfaces [145]. The amorphous Ta2O5 coating was tested to be more hydrophobic with lower surface energy, showing decreased colonization of both S. aureus and Actinobacillus actinomycetemcomitans, as well as compromised viability of human skin fibroblasts. On the other hand, the hydrophilic crystalline Ta2O5 coating exhibited improved viability of human skin fibroblasts compared with uncoated surface, however the colonization of bacteria was also accelerated, suggesting that further designs to inhibit bacterial colonization without hampering the viability of tissue cells.

Gallium ion (Ga3+) has emerged as a novel antibacterial candidate due to its similarity with ferric ion (Fe3+) in terms of charge, ionic radius, mass, electronic configuration, and coordination number, which makes it difficult for bacteria to distinguish Ga3+ from Fe3+ in iron utilization [175]. Once Ga3+ deceives bacterial ingestion system and involves in electron exchange reactions, iron-seizing bacteria would be starved to death since Ga3+ lacks the activity to transfer electrons like Fe3+ do with its potential interconversion with ferrous ion (Fe2+) [176]. Based on such antibacterial mechanism, a Ga3+-doped TiO2 nanopored Ti surface was prepared via gallium nitrate treatment, which showed enhanced antibacterial efficacy compared with pure TiO2 nanopored surfaces, while the bioactivity of HGFs was favorably maintained on as-modified surface [177]. Therefore, incorporating Ga3+ can be considered as a potential surface modification strategy to address the problems of poor soft tissue seal and bacterial infection. Further research should adopt in vivo to evaluate the antibacterial efficacy and the promoting effect in soft tissue seal to narrow the gap between the laboratory research and clinical practice.

The antibacterial ability of copper (Cu) has been widely recognized [178]. However, the Cu-coated surface showed negative influence on the viability and proliferation of HGFs mainly due to the Cu-induced cytotoxicity [179]. Since the biological effect of Cu is does-dependent [180], further efforts should be made to find out the ideal therapeutic dosage range between the antibacterial and cytotoxic effect.

6.3 Employing intrinsic antibacterial agents

The intrinsic antibacterial agents discussed here can be categorized into unique materials, natural polymers and synthesized polymers. Graphene oxide (GO), as an oxidized derivative of the two-dimensional carbon material graphene, is more applicable than graphene nanosheets and is the representative of unique materials [181]. GO possess satisfactory antibacterial activity, where the possible antibacterial mechanisms were suggested to be associated with physical damage theory [182], oxidative stress theory [183], or phospholipid molecular extraction damage theory [184]. Moreover, GO exhibits good biocompatibility and can facilitate stem cell proliferation and osteogenic differentiation [[185], [186], [187], [188]]. Based on the excellent antibacterial effect and bioactivity, GO coating was deposited on PEEK surface via the assistance of PDA self-polymerization [148]. The obtained GO coating could effectively kill P. gingivalis, F. nucleatum and S. mutans. Meanwhile, GO coating also promoted adhesion and proliferation of HGFs. To better regulate the coating thickness and morphology, electrophoretic deposition was employed to prepare GO coating with varied deposition voltages. GO coatings prepared at higher voltage possessed increased layered number, thickness, roughness and electropositivity. Increasing the layer number also resulted in an augmented ROS level, which enhanced the antibacterial effect in both monoculture and HGFs-bacteria coculture model [147].

PDA and chitosan are natural polymers with intrinsic antibacterial properties. PDA is derived from mussels and possesses extraordinary adhesive ability to either organic or inorganic materials [189]. The coating of PDA can be easily prepared on Y-TZP surface via immersion with dopamine solution since the self-polymerization activity of dopamine [149]. And the PDA-coated Y-TZP surfaces were demonstrated to decrease adhesion of S. gordonii, S. mutans and P. gingivalis while promoting the adhesion of HGFs [125,149]. The antibacterial mechanism was not fully unveiled, with probable explanations that the negatively charged hydrophilic PDA coating reduced the negatively charged salivary proteins and the following bacterial adhesion [149]. On the other hand, chitosan is a polysaccharide with high deacetylation degree, which indicates the polycationic nature of it [190]. The antimicrobial activity of chitosan can be attributed to its interaction with anionic components of bacteria, which leads to the increased permeability of cell and results in the bacterial death [191]. As such, chitosan coatings were produced on Ti surface via catechol-based conjugation or polyethylene glycol crosslinking, exhibiting bactericidal effect on S. aureus and E. coli and cytocompatibility with human fibroblasts [150,151].

Synthesized polymers exert antibacterial activities through anti-fouling mechanism, contact-killing/release-killing mechanism. In terms of anti-fouling mechanism, a 2-hydroxylethyl methacrylate (HEMA) polymer coating was fabricated on Ti surfaces to inhibit bacterial adhesion, due to the zwitterionic nature of HEMA and the as-formed superhydrophilic hydration layer [152]. It was demonstrated that the neutral, watery HEMA surface can prevent bacterial attachment by maximally reducing the electrostatic and hydrophobic attractions between bacterial and the substrate [192,193]. Meanwhile, such superhydrophilic surface showed increased fibroblast adhesion compared with pure uncoated surface [152]. In addition to HEMA coating, the amphipathic diethyl phosphite (DEP) coating exhibited similar property in constructing a superhydrophilic surface and making the surface amphoteric to inhibit the adhesion of bacteria while supporting the attachment of fibroblasts [153].

In terms of contact-killing mechanism, dimethylaminohexadecyl methacrylate (DMAHDM), as a quaternary ammonium methacrylate, showed a broad antimicrobial spectrum due to its positive charges, like chitosan [194]. Under the assistance of PDA, DMAHDM and hydroxyapatite were coated on alkaili-heat treated Ti surface to deliver bactericidal effect against S. aureus, Streptococcus sanguinis and human saliva-derived biofilms, with the viability of human fibroblasts uncompromised [154]. Based on similar antimicrobial mechanism, a renewable antibacterial platform was produced on Ti with polyacrylic acid and sodium hypochlorite (NaClO). The as-obtained N-halamine polymeric coating exhibited lasting bactericidal effect through the contact-killing or release-killing of the active chlorine (Cl+) in the metastable N-Cl bond, without exerting negative influence on the soft tissue. In addition, the depleted N-Cl could be easily reloaded via peri-implant NaClO solution irrigation, which equipped it with promising applications in the clinical practice [155].

6.4 Developing stimuli-responsive bactericidal platforms

To achieve a real-time on-demand activation of antibacterial function, external stimuli were exerted upon specific responsive materials to deliver bactericidal effects. Photodynamic therapy (PDT) has drawn increasing attention due to its desired capability in eradicating oral pathogenic microorganisms and inhibiting biofilm formation through photo-induced generation of ROS (·OH, O2−, HO2−, and H2O2) [195]. In specific, under UV-C irradiation (200∼280 nm), TiO2 as a semiconductor, could induce the production of ROS on the surface [60]. However, the UV radiation can be harmful when directly used on viable tissues, therefore designs on how to fabricate a modified surface that can respond to safer visible light or infrared radiation were proposed. In particular, AuNPs were applied to generate ROS when visible light radiation (420∼480 nm) was given, where the as-obtained ROS could cause bacteria death [95]. Therefore, an on-demand bactericidal action triggered under viable tissue-friendly visible light radiation was proposed. It is worth noting that the generation of ROS should be controlled in a limited level, since excessive ROS would damage the surrounding tissue cells as well. In order to evaluate the biocompatibility of this ROS-generating system, an animal model was established and subsequent H&E staining was conducted, confirming that the light-responsive AuNPs system could effectively eliminate bacterial infection without causing significant inflammatory infiltration in the treated region [95]. However, there lacks the quantified examinations and control of the generated ROS in cells or tissue that received the ROS-driven antibacterial therapies, which are suggested to be included in future research.

In addition to PDT, near infrared (NIR)-induced photothermal therapy (PTT) had also been combined with adhesion-promoting surfaces to simultaneously achieve antibacterial effect as well as enhanced cell adhesion and proliferation. PDA have been widely used to construct PTT surface due to the extraordinary photothermal conversion performance and biocompatibility [196]. For example, PDA could be incorporated with gelatin and hydroxyapatite nanoparticles to construct a multifunctional coating, where gelatin, as a partial hydrolysate of collagen, could induce the expression of adhesion-related genes and proteins in HGFs, and the elevated temperature inhibited the growth of S. mutans under NIR irradiation [106]. In addition to PDA, GO also received a widespread recognition as photothermal material [197]. To optimize the balance between photothermal bactericidal effect and biocompatibility, a sandwich structured coating has been constructed by wrapping GO under collagen with the assistance of PDA. The GO nanosheet in the inner layer could generate heat under NIR irradiation to inhibit the growth of oral pathogens including S. sanguinis, F. nucleatum and P. gingivalis, while the collagen in the outer layer not only promoted the adhesion and proliferation of HGFs, but also effectively protected the cells and tissues from damage of GO [127]. TiO2 itself is also a applicable light-responsive medium for PTT [198]. A micro-porous TiO2-x layer was fabricated via plasma electrolytic oxidation and thermal reduction to perform excellent bactericidal photothermal effect based on the light multi-reflection inside the pores, while additional cerium (Ce) was doped to scavenge excessive ROS and facilitate soft tissue regeneration [157].

Electrostimulation therapy (EST) represents another on-demand therapeutic approach capable of preventing bacterial infection while improving soft tissue seal through a low level of current/voltage at the implanted site [[199], [200], [201]]. An in vitro study showed that the application of EST dosages (1.5 V for 5 min per day) onto a conductive Ti nanopore surface exhibited significant antibacterial efficacy towards human oral saliva biofilm [Fig. 10g∼h] [100]. The antibacterial mechanism could be ascribed to the combination of increased pH value and generation of free radicals at the electrode surface under the applied voltage [202]. Since EST has been proved to increase intercellular Ca2+ and promote growth factor secretion [203], the application of electrical stimuli onto the conductive Ti nanopore surface could also promote fibroblast attachment, proliferation and collagen secretion [100]. The therapeutic effect of EST was further evaluated via an in vivo infection model, where a rechargeable hydrazine hydrate-derived polyimide (HPI) coating has been deposited onto Ti surface to realize temporal-sequence regulation of antibacterial activity and soft tissue seal [158]. Specifically, after being charged with negative current, the post-charged coating could effectively prevent biofilm formation due to the contact-dependent violent electron transfer between bacteria and the surface, leading to the disruption of bacterial membrane, generation of intracellular ROS and inhibition of ATP synthesis. Later in the discharge stage, the attenuated discharging status could create a gentler electron-transfer microenvironment which was advantageous to fibroblast adhesion and growth, so that facilitating the establishment of soft tissue seal.

Sonodynamic therapy (SDT) seems to complete the puzzle of external stimuli-responsive platforms based on sound, light and electricity. Similar with PDT, SDT takes use of noninvasive ultrasound and sonosensitizer to produce ROS for antimicrobial applications [204]. Ultrasound can trigger the piezoelectric effect (charge transfer to the opposite surface of the material) in barium titanate (BaTiO3) [205], yet in a limited efficacy due to the large bandgap and high electron-hole recombination of BaTiO3 [206]. Therefore, a defective BaTiO3-x was produced via sodium borohydride reduction, and L-arginine was adopted to convert the generated ROS into reactive nitrogen species (RNS), which showed higher antibacterial activity than ROS [159]. Moreover, a nanorod array structure on Ti was designed to alleviate the inflammatory level induced by the produced ROS, thereby eliminating infection and protecting the soft tissue. Simpler than the design of BaTiO3-x/L-arginine nanorods, AuNPs-loaded TiO2 was proved to deliver sonodynamic bactericidal effect and create a repair-supportive microenvironment for soft tissue as well, probably due to the intrinsic immunomodulatory ability of AuNPs [160].

7 Modification strategies based on immunomodulation to facilitate soft tissue seal

Whether a robust peri-implant soft tissue seal could be built is influenced by the whole dynamic process of peri-implant wound healing. Specifically, the series of host immune responses occurring along with the insertion of the implant transmucosal region could be described as follow [207]: (1) bleeding and coagulation; (2) acute inflammatory response, mainly featuring in eliminating pathogens and clearing damaged tissue fragments; (3) attenuated inflammation period, featuring in an altered composition of immune cells and cytokines, preparing for the subsequent soft tissue regeneration stage; (4) granulation tissue stage, identified by soft tissue cells proliferation and neovascularization; (5) ECM remodeling stage, with collagen fibers becoming dense and organized while cells proliferating slow, finally ending in the establishment of soft tissue seal.

The afore-mentioned antibacterial methods mainly take effect in the acute inflammation stage (Stage 2) or after the seal establishment, while the primary adhesion-promoting properties play an essential role in the granulation tissue stage (Stage 4) and ECM modeling stage (Stage 5). Hence, surface modifications based on the initial coagulation stage (Stage 1) and the attenuated inflammation period (Stage 3) would be discussed in this section [Fig. 11].Fig. 11 Illustration showing the surface modification strategies based on immunomodulation to facilitate soft tissue seal, which include accelerating blood coagulation, and attenuating inflammation to create a repair-supportive immune microenvironment. PAR4-AP protease activated receptor 4-activating peptide, PDGF-B platelet-derived growth factor subunit B, TNF-α tumor necrosis factor α, IL-1β interleukin 1β, TGF-β transforming growth factor β, IL-10 interleukin 10, modSLA modified sandblasted, large-grit, acid-etched technique, IL-4 interleukin 4, CCN2 cellular communication network factor 2, IL-23 interleukin 23, iNOS inducible nitric oxide synthase, CLA conjugated linoleic acid, CO carbon monoxide, CeO2 ceria, Fe-NC iron-nitrogen-doped carbon single-atom.

Fig. 11

7.1 Constructing a functional surface to accelerate blood coagulation

In the first stage, blood contacts with the transmucosal region surface, where platelets and fibrin adhere to the surface, initiating blood coagulation and secreting growth factors [208]. Based on the activation function of platelets, several surface modifications were conducted. Previously, a protease activated receptor 4-activating peptide (PAR4-AP) was proposed, which could mediate the platelet-fibrin clot formation [209]. The PAR4-AP was linked onto Ti or Y-TZP surfaces via phosphoric acid and EDC/NHS chemistry, resulting in a better adhesion of platelets, and promoting the epithelial cell adhesion [210,211]. In addition to clot formation, platelets also secrete various growth factors, aiding in create a repair-supportive immune microenvironment. Therefore, through loading cellulose with platelet lysate on a modified Ti surface, the M2 polarization of the incubated macrophages and improved fibroblast adhesion were successfully induced, since the platelet lysate contained various bioactive molecules, including growth factors, anti-inflammatory mediators, antimicrobial peptides, and scaffolding proteins [Fig. 12a∼d] [130]. Through the direct usage of platelet-derived growth factors, a polyethylenimine-plasmid DNA (pDNA) nanoplexes encoding platelet-derived growth factor subunit B (PDGF-B) was constructed on Ti surfaces to transfect the co-incubated HGFs with the pDNA, enabling the fibroblasts to display PDGF secretion and increased integrin 2 expression, thereby facilitating the adhesive interaction of HGFs to collagen and ECM [212].Fig. 12 Illustration showing the immunomodulatory activities of coagulation-related agent, growth factor, and nanozyme. (a) Atomic force microscope (AFM) topography image of an aligned platelet lysate (PL) sample on the border of the cellulose nanocrystals (CNC)-coated titanium surface. The arrow indicated the approximate local radial axis on that part of the surface. (b) Directionality histogram of CNC orientation at both center and border of the aligned surfaces. (c) Proteomic analysis of the PL proteins adsorbed on CNC surface. Proteins organized according to their protein class. Categories with a representation lower than 1 % were grouped in “other,” The percentage (%) corresponds to the fraction of protein hits in a category against the total number of process hits. (d) Macrophages incubated on titanium (Ti), CNC-coated Ti (Ti/CNC) surface, or CNC-coated PL aligned Ti (Ti/CNC/PL) surface were immunostained with anti-CCR7 (M1-like) or anti-CD206 (M2-like), which appeared in green, while nuclei and cytoskeleton appeared in blue and red, respectively. Adapted with permission from Ref. [130]. Copyright 2021 Wiley-VCH GmbH. (e)Schematic diagram of preparing a CCN2 bearing mesoporous silica nanoparticles (MSNs) on micro-arc oxidated (MAO) titanium substrates (CCN2@MSNs-Ti) to enhance peri-implant soft tissue seal. CCN2 cellular communication network factor 2 (f) Comparison of soft tissue around implants in Ti and CCN2@MSNs-Ti groups of rats at 4 w post implantation. (g) SEM images of the interfaces between the soft tissue and implant surface. The shapeless and rough Ti oxide layers formed by MAO treatment were visible on the surfaces of CCN2@MSNs-Ti implants, which were shown by a red dashed curve. (h) Van Gieson staining images of the rat tibias, wherein soft tissue was stained purple and new bone formation was stained red. The yellow dashed curve outlined the extent of epithelial downgrowth. I implant, S soft tissue, B bone. (i) Relative mRNA expression levels of inflammation-related factors (TNF-α and IL-1β). Data were presented as mean ± S.D. (n = 3, *p < 0.05, **p < 0.01, ****p < 0.0001). Adapted with permission from Ref. [221]. Copyright 2023 American Chemical Society. (j) Schematic showing the construction of nanoceria-coated Ti disks (Ti@Ce). (k) Intracellular ROS levels in HGFs as indicated by green signals under 500 μM H2O2 exposure. (l ∼ o) Heat map showing DEGs in Ti@Ce_500 μM, Ti_500 μM, Ti_0 μM, and Ti@Ce_0 μM from specific GO terms related to cell adhesion, antioxidant activity, apoptosis process, and cellular response to FGF stimulus. HGFs human gingival fibroblasts, FGF fibroblast growth factor. Adapted with permission from Ref. [239]. Copyright 2024 American Chemical Society.

Fig. 12

7.2 Constructing a bioactive surface to attenuate inflammation

In order to attenuate the inflammatory microenvironment and accelerate the wound healing around the transmucosal region, surface modifications based on physical surface property alterations, anti-inflammatory cytokine coatings and immunomodulatory agent incorporations were investigated.

7.2.1 Physical surface property alterations

The roughness, topography, and hydrophilicity of the surface not only influence the adhesion of soft tissue cells and the attachment of ECM, but regulate the behavior of immune cells (macrophages, dendritic cells, etc.) and the cytokine profile of the microenvironment as well [213,214]. The influence of surface roughness on macrophage polarization was investigated through producing four samples with different roughness on Ti substrates (Ra = 0.20 μm, 0.51 μm, 1.36 μm, and 2.60 μm, respectively). By comparing the expression level of M1 and M2 markers, it was found that the samples with the Ra ranging from 0.51 μm to 1.36 μm induced higher ratio of M2 polarization, while the smoother or rougher samples produced higher ratio of M1 polarization [215]. For surface topography, it was demonstrated that a hierarchical topography with both microgrooves and TiO2 nanopores significantly reduced the adhesion of macrophages while exerting little influence on the adhesion of HGFs [216]. Moreover, a honeycomb-like TiO2 topography with the diameter of approximately 90 nm were tested to remarkably induce M2 polarization of macrophages and produce a pro-regenerative microenvironment [217].

As for surface hydrophilicity, the hydrophobic SLA surface and hydrophilic modified SLA surface (modSLA surface) were used to examine the effect of surface hydrophilicity on the behavior of macrophages and dendritic cells. Specifically, through incubating macrophages on different surfaces, macrophages on the hydrophilic modSLA surface exhibited downregulated gene expression of pro-inflammatory interleukin 1β (IL-1β) and TNF-α, compared with that on the hydrophobic SLA surface [218]. Apart from macrophages, dendritic cells play also play a pivotal role in presenting antigen of the implanted biomaterial and mediating the later immune responses [213]. Through comparing the cytokine profile produced by dendritic cells incubated on SLA and modSLA surfaces, it was found that the supernatants obtained from the modSLA group included higher level of anti-inflammatory cytokines (TGF-β and IL-10), indicating a pro-regenerative cytokine profile [219].

Nevertheless, most of the studies mentioned in this subsection have stopped the research on the influence of surface properties on the behavior of immune cells, while some have extended the research to investigate the influence of the produced cytokine profile on bone regeneration, yet none of the studies have drawn the whole picture from the alteration in surface physical properties, to the changes of produced cytokine profile, finally to the influence on soft tissue seal. As such, in the following subsections, we have selected the studies which discussed the interplay between the immune cells and the soft tissue cells.

7.2.2 Anti-inflammatory cytokine coatings

In this paragraph, surface modifications based on soft tissue growth factors and anti-inflammatory interleukins were discussed. In particular, cellular communication network factor 2 (CCN2, also known as connective tissue growth factor CTGF) was recognized as a potent effector to promote fibroblast biological functions, where TGF-β being the potential upstream signal [220]. Recently, a long-acting CCN2 delivery system was constructed on Ti surfaces via micro-arc oxidation pretreatment and mesoporous silica nanoparticles loading, resulting in a decreased expression of inflammation-related factors including TNF-α, IL-1β, and interleukin 6 (IL-6) in soft tissue, and an intact in vivo soft tissue-Ti interface, indicating a satisfying peri-implant soft tissue seal [Fig. 12e∼i] [221].

Interleukins, on the other hand, were also explored for potential applications on surface modifications to alleviate inflammation and accelerate soft tissue healing. Interleukin 4 (IL-4), for example, as a typical stimuli to induce the M2 polarization of macrophages, was here utilized for obtaining a declined inflammatory level [222]. In specific, IL-4 was immobilized on Ti alloys surfaces via PDA coating, and the as-obtained surface was confirmed in vivo to successfully induce the M2 polarization of macrophages and boost the laminin 332 secretion in peri-implant junctional epithelium, predicting an enhanced sealing property [223,224]. In addition, interleukin 23 (IL-23) and related molecules were verified to regulate peri-implant inflammation as well. It was demonstrated that IL-23 was at a distinctively high level in the crevicular fluid from peri-implantitis sites [225], and IL-23 played a critical role in mediating autoimmune destruction [226]. As such, a nanocoating of IL-23 receptor antagonist (IL-23Ra) was fabricated on a Ti surface by alkai-heat pretreatment and silanization-mediated peptides immobilization, leading to the block of the IL-23/interleukin 17A (IL-17A) immune pathway in epithelial cells, which induced the M2 polarization of human macrophage-like cells and created a repair-supportive microenvironment [227].

7.2.3 Natural or synthesized anti-inflammatory agent incorporations

Besides, non-cytokine molecules were utilized in anti-inflammatory surface modifications as well. For instance, conjugated linoleic acid (CLA) was demonstrated to suppress the pro-inflammatory cytokines production of peripheral blood CD4+ and CD8+ T cells in patients with Crohn's disease [228]. Therefore, CLA was selected to be immobilized onto Ti surfaces to suppress the expression of M1 phenotype marker (inducible nitric oxide synthase, iNOS) of the murine macrophages, while promoting the expression of M2 phenotype marker (CD206), with the co-immobilized LamLG3 peptide capable of inducing an enhanced adhesion of epithelial cells, together improving the peri-implant soft tissue seal [118]. In addition to CLA, carbon monoxide (CO) has been gaining increasing attention as an immunomodulatory and antibacterial molecule [229]. It has been demonstrated that CO conducts immunomodulatory effects through the inhibition of a variety of pro-inflammatory signaling pathways [230], and the promotion on the macrophage polarization towards the pro-regenerative M2 phenotype [231]. Meanwhile, CO at a high concentration exhibits antibacterial activities via bacterial membrane permeabilization, ATP synthesis disruption, and cellular respiration inhibition [232,233]. Accordingly, a controllable on-demand CO releasing system was fabricated on Ti surface via coupling of pentaerythritol tetraenyl (3-mercaptopropionic acid), achieving burst CO release under NIR irradiation to eradicate bacteria, and slow CO release in the absence of NIR to regulate the immune microenvironment to a pro-regenerative condition [229].

Nanozymes capable of eliminating ROS and thereby exhibiting anti-inflammatory effect were extensively investigated over the decades [234]. Compared with traditional ROS-scavenging enzymes, nanozymes exhibited advanced chemical stability with lower cost [235]. Among others, Ce received widespread attention due to its ability to initiate Fenton reaction and scavenge ROS via Ce3+/Ce4+ conversion [236]. On this base, multiple surface modifications in form of cerium ions or oxides (CeO2) were developed on Ti/PEEK surfaces to confirm the altered level of cytokines (including inflammatory TNF-α, IL-1β and IL-6, and repair-related IL-10 and vascular endothelial growth factor, VEGF) in the soft tissue [Fig. 12j∼o] [[237], [238], [239]]. Apart from Ce, an iron-nitrogen-doped carbon single-atom (Fe-NC) nanozymes loaded TiO2 nanotubes was fabricated by electrochemical anodization and solvothermal method to effectively reduce the intracellular ROS level in macrophages and fibroblasts, due to the superoxide dismutase (SOD) and catalase (CAT) mimicking activities of Fe-NC nanozymes [240].

7.3 Balancing the immunomodulatory activity and antibacterial effect

The immunomodulatory surface designs aim to attenuate the inflammation and create a repair-supportive microenvironment, while the antibacterial surface modifications sometimes induce aggravated inflammation to eliminate pathogenic microorganisms. Herein, two main strategies can be summarized to design dual-functional surfaces.

One of the strategies is the design of sequential release of antibacterial/immunomodulatory agents. For instance, a smart temporal-sequence functioning system based on antibacterial calcium superoxide (CaO2) and anti-inflammatory IL-4 was constructed [241]. Such hierarchical structure exhibited incipient bactericidal effect via the excessive ROS generation of outer CaO2-embedded methacrylated gelatin, and the subsequent inflammation-alleviating effect via the release from inner IL-4-loaded TiO2 nanotubes, thus realizing the combined anti-infection/anti-inflammation function of the surface. This TiO2 nanotubes/hydrogel structure is inspiring for sequential release of agents, while other drug delivery systems like mesoporous silica nanoparticles (MSNs) and extracellular vesicles (EVs) can also be incorporated with micro/nano surface patterns to realize the sequential delivery of antibacterial and immunomodulatory agents [242,243].

Another strategy is constructing immunomodulatory topographies. For example, a sulfonated PEEK substrate with Cu coating was produced via sulfuric acid solution immersion and magnetron sputtering [244]. Through copper ion (Cu2+) release, the surface showed satisfactory bactericidal activity against methicillin-resistant S. aureus, while the sulfonated PEEK surface induced the macrophages to secrete increased level of anti-inflammatory IL-4 compared with PEEK surface. However, such Cu-related surface modification should be further managed to improve its biocompatibility with soft tissue before being employed as a multifunctional surface for enhancing soft tissue seal. Similar immunomodulatory topography was also applied in the afore-mentioned ultrasound-responsive BaTiO3-x/L-arginine coating [159], where a nanorod topography facilitated the M2 polarization of macrophages while the ultrasound stimuli triggered on-demand RNS generation exerts bactericidal effect.

8 Modifications for concurrent soft tissue seal and osseointegration

Among the included surface modifications, several designs have successfully facilitated both superficial soft tissue seal and underlying osseointegration of dental implants, which is consistent with the idea of integrated regeneration of soft/hard composite tissues [69,106,156]. On one hand, the soft tissue seal exists as a barrier for the underlying bone tissue during the formation of initial osseointegration. After implantation, the disrupted peri-implant mucosa would quickly heal to form a seal around the implant (the initial epithelial seal establishes in about 2 week) [41], while the formation of the underlying osseointegration takes months. The established soft tissue seal helps the implant-bone interface to shield outside contaminants and avoid the unexpected transfer of occlusal force, thereby facilitating osseointegration. On the other hand, the quality of osseointegration determines the attaching position of soft tissue, since there is a constant distance of 3∼4 mm from the margin of the peri-implant mucosa to the first bone-to-implant contact or the stabilized top of the adjacent bone [20]. This means undesired epithelial downgrowth would probably occur if the osseointegration was poor. Herein, the surface modifications accomplishing concurrent soft tissue seal and osseointegration were summarized in Table 5.Table 5 Surface modifications for concurrent soft tissue seal and osseointegration.

Table 5Substrates	Coatings	Soft tissue cells responses	Bone-related cells responses	In vivo tests	Ref.	
Ti	PDA/HA-embedded gelatin coating	The cell proliferation of HGFs was maintained, while the cell adhesion and the genes expression of ITGA3, ITGB1, VCL, COL-I were promoted.	The cell proliferation of BMSCs was maintained, while the cell adhesion, the genes expression of RUNX-2, BMP-2, OCN, OPN, ALP activity, and ARS staining were improved	–	[156]	
Y-TZP	Akermanite coating	The cell proliferation of HGFs was maintained, while the spreading, genes expression of ITGA5, ITGB1, COL-I were promoted.	The cell proliferation of BMSCs was maintained, while the cell migration, genes expression of RUNX-2, BSP, OCN, COL-I, ALP activity, and ARS staining were improved.	Female rabbit: Femoral bone implantation model, 6 w and 12 w new bone generation and bone-implant contact were increased.	[106]	
PEEK	SrTiO3 coating and UV-C irradiation	The cell viability, adhesion and spreading of GE1 and NIH3T3 cells were improved. The protein expression of integrin β4 in GE1 cells and vinculin in NIH3T3 cells were promoted.	The cell viability, adhesion, spreading, ALP activity and ARS staining of MC3T3-E1 cells were improved.	–	[69]	
Ti titanium, Y-TZP yttria-stabilized tetragonal zirconia polycrystals, PEEK poly-ether-ether-ketone, PDA polydopamine, HA hydroxyapatite, SrTiO3 strontium titanate, UV-C ultraviolet C, HGFs human gingival fibroblasts, ITGA3 integrin α3, ITGB1 integrin β1, VCL vinculin, COL-I collagen I, ITGA5 integrin α5, GE1 cells gingival epithelial cells, NIH3T3 cells fibroblast, BMSCs bone marrow mesenchymal stem cells, RUNX-2 runt-related transcription factor 2, BMP-2 bone morphogenetic protein 2, OCN osteocalcin, OPN osteopontin, ALP alkaline phosphatase, ARS staining Alizarin Red S staining., BSP bone sialoprotein, MC3T3-E1 cells osteoblast-like cells.

On Ti substrate, a PDA/hydroxyapatite-embedded gelatin coating was fabricated through alkaili-heat pretreatment, solution immersion of dopamine and hydroxyapatite nanoparticles, and gelatin setting [156]. The coated surface displayed adhesion-promoting effect on HGFs, facilitated the osteogenic differentiation of bone marrow mesenchymal stem cells (BMSCs) and accelerated the mineralization of the new bone matrix. The promotion on osseointegration is explained by the introduction of hydroxyapatite, which is a recognized osteoconductive material with widespread applications in Ti implants [245]. For enhancing the soft tissue seal, gelatin, as a partial hydrolysis product of collagen, was employed to provide a biomimetic scaffold [246].

On Y-TZP substrate, an akermanite coating by negative pressure sol infiltration and sintering was presented, where different infiltration time produced diversified surface morphologies [106]. In particular, the group received 15 min of akermanite sol infiltration exhibited highest hydrophilicity and moderate number of akermanite crystals within zirconia crystals. In essence, akermanite is Ca2MgSi2O7, and the effect of Ca, Mg and silicon (Si) on bone regeneration was identified [247]. As for the positive influence on soft tissue seal, a combination effect of surface morphology, hydrophilicity and active ion release was suggested.

On PEEK substrate, a SrTiO3 coating was achieved via magnetron sputtering [69]. From the viewpoint of designing, strontium (Sr) is an intrinsic osteogenic element human body [248], while SrTiO3 is a photocatalytic material [249], which make the SrTiO3 coating a UV-responsive osteogenic modification. In addition, coatings with different thickness (100 nm or 200 nm) were prepared, where the most sustained strontium ion (Sr2+) release was proved to be from the specimen with 200 nm thick coating. As such, the specimen with 200 nm thick coating of SrTiO3 was selected for further UV-C photofunctionalization for 30 min, which displayed improved surface wettability without morphology alteration than UV-C treated bare PEEK, thereby facilitating the soft tissue seal.

However, as indicated in Table 5, limited studies have investigated the concurrent soft tissue seal and osseointegration on Ti/Y-TZP/PEEK, and some of the listed studies still lack higher evidence of in vivo tests. The integrated regeneration of periodontal soft/hard tissues should be further emphasized.

9 Conclusions and future perspectives

An ideal transmucosal region surface is expected to facilitate peri-implant soft tissue seal which acts as a protective barrier against bacterial invasion and guarantees the long-term functioning of implants. Some of the previously mentioned surface modifications were tested in clinical trials and presented promising future applications. For example, compared with a conventionally-polished zirconia abutment, an ultra-polished zirconia abutment was confirmed to result in a smaller probing depth at 1st month after the final restoration, indicating an increased sealing capacity [250]. However, after one-year follow-up, the probing depth around the ultra-polished abutment displayed no significant difference with the control abutment, suggesting that the ultra-polished surface only produced short-term improvement of the sealing capacity. On the other hand, a series of clinical tests had been carried out to evaluate the influence of plasma pretreatment on the sealing ability of Ti abutments. It was proved that the application of argon plasma pretreatment could result in an increased collagen density and more obliquely oriented fibers in peri-implant connective tissue after two week of healing, as well as promoting the keratinization and maturation of peri-implant epithelium at 2nd month, supporting the establishment of a robust soft tissue seal [[251], [252], [253]].

Though ultra-polishing and plasma pretreatment were confirmed via clinical trials to be able to improve the soft tissue attachment around the transmucosal region, further antibacterial effect and immunomodulatory ability are also expected. For the realization of long-term peri-implant soft tissue health, the present review reported the recent advances versatile in promoting soft tissue attachment, preventing bacterial infection and creating a repair-supportive immune microenvironment. Firstly, improving surface hydrophilicity, or constructing specific micro/nano textures (grooves or tubes), or coating with bioactive components (metallic ions and ECM proteins) could facilitate epithelial sealing and connective tissue attachment. Secondly, applying antibacterial peptides (lactoferrin-derived amyloid) and agents (AgNPs or Ga3+), with or without external stimuli-responsive properties (photo, electrical, or ultrasound stimuli), could prevent bacterial infection while encouraging soft tissue cells adhesion. Thirdly, employing immunomodulatory coatings (IL-23Ra or cerium) is capable of creating a repair-supportive immune microenvironment towards accelerated soft tissue healing. Nevertheless, there is still a gap from bench to bedside for these elaborately designed surfaces, where evidence of higher grade is demanded.

To this end, animals models (including rats, rabbits, dogs and pigs) were established, yet the obtained results could not be directly translated in human circumstances, which calls for further pre-clinical trials [254]. Previously, some clever human biopsy protocols were designed. For instance, the application of a thinner test abutment realized the obtaining of a complete abutment-soft tissue biopsy, and the subsequent placement of a healing abutment with normal diameter was able to produce the patient with a satisfying restorative outcome [42]. More effective and harmless protocols for clinical trials should be developed, thereby advanced experimental evidence can be obtained to apply the novel transmucosal region surface modifications to the clinic.

Meanwhile, there are still obstacles in entirely fulfilling the clinical demands in terms of adhesion-promoting, antibacterial and immunomodulatory properties. Based on the clinical requirements, the future research orientations in this field could be proposed as follow [Fig. 13].Fig. 13 Illustration showing the future perspectives for enhancing peri-implant soft tissue seal, where constructing bioinspired surfaces, hybrid surfaces, multifunctional surfaces, and intelligent surfaces could be considered. OEpSCs odontogenic epithelial stem cells, DFCs dental follicle cells, ROS reactive oxygen species.

Fig. 13

(1) Construction of bioinspired surfaces to mimic natural dental tissues. From a biomimetic perspective, the optimal surface for epithelial attachment is enamel, and the ideal surface for collagen protrusion is cementum. Therefore, an enhanced peri-implant soft tissue seal could be obtained if there was enamel/cementum deposition on the material transmucosal region, simulating natural tooth surfaces. Since in vitro artificial deposition of enamel and cementum could be challenging, dental stem cell-engineered surfaces being able to biologically deposit enamel or cementum could be constructed. Over the decades, various dental stem cells of diversified histological origins were isolated and confirmed to be able to regenerate tooth tissues [255], where dental follicle cells (DFCs) could differentiate into cementoblasts and synthesize cementum [256], and odontogenic epithelial stem cells (OEpSCs) were capable of generating enamel, though it would take efforts to detect OEpSCs that hid in the postnatal dental lamina, the epithelial rests of Malassez or the reduced enamel epithelium [257].

(2) Construction of hybrid surfaces for soft/hard tissues integrated regeneration.

From a clinical perspective, the repairing of bone and its attached soft tissues is always needed. And this has led to the idea of interface tissue engineering, which aims to repair or fill the damaged or diseased zones between different tissue types (e.g. bone and soft tissue) [258,259]. It is reported that the repairing of bone-attached tendon or ligament can be realized through applying scaffold materials and mesh materials [[260], [261], [262]], since these materials could form a structure mimicking the biological bone-ligament interface. Besides, bioprinting systems can also be introduced to develop sequential deposition of the materials [263]. In the field of dental implant, osseointegration and soft tissue seal have been widely investigated respectively, however the integrated repairing of the soft/hard composite tissues received insufficient attention. More efforts should be made to realize the concurrent regeneration of peri-implant soft/hard tissues.(3) Construction of multifunctional surfaces to balance antibacterial effect and immunomodulation. Introducing antibacterial materials could lead to aggravated inflammation, since the antibacterial materials often exert bactericidal effects through producing excessive ROS or releasing foreign body particles. Both of uncontrollable microbial proliferation or lasting inflammatory response will compromise the outcome of soft tissue seal. A series of ROS-scavenging biomaterials (polyphenols, meta or carbon-based artificial nanozymes) can be combined with the antibacterial strategies to eliminate the excessive ROS in the immune microenvironment [234], thereby creating a repair-supportive immune microenvironment. For example, a chlorin e6 (Ce6)-coated nanoceria composite was synthesized to perform versatile functions under the excitation of red-light irradiation [264]. In specific, the photosensitizer Ce6 conducted responsive antibacterial PDT by overwhelming ROS generation, followed by the antioxidant activity of nanoceria to scavenge the residual ROS and downregulate the M1/M2 ratio of macrophages and create a repair-supportive microenvironment.

(4) Construction of intelligent surfaces combining diagnosis and treatment. Maintaining peri-implant soft tissue seal is a long-term career, and the seal is in a constantly dynamic environment. Up to now, barely no protocols can conveniently monitor the changing situations around the modified transmucosal region, failing to reflect the full-course interaction between the implanted surface and the surrounding tissues. Therefore, it would be helpful to develop a platform to provide timely reports of the sealing condition, the bacterial invasion and the immune microenvironment profile [265,266]. In addition to monitoring the development of the pathological conditions, it would be more satisfactory to enable the surface to deliver therapeutic agents or effects in response to the pathological conditions [267,268], for example, applying pH-responsive materials like metal-organic frameworks (MOFs) or layered double hydroxides (LDHs) which precisely respond to the acidic microenvironment produced by bacterial metabolism [[269], [270], [271]]. Upon realizing this concept, an intelligent surface with timely monitoring and self-repairing abilities can be built. Managing peri-implant health requires persistent efforts, where the idea of integration of diagnosis and treatment ought to be encouraged and incorporated in transmucosal surface constructions, thus alleviating the burden of peri-implant maintenance therapy.

Ethics approval and consent to participate

Not applicable.

CRediT authorship contribution statement

Siqi Jin: Writing – review & editing, Writing – original draft, Visualization, Validation, Project administration, Conceptualization. Yameng Yu: Writing – review & editing, Writing – original draft, Visualization, Conceptualization. Ting Zhang: Validation, Software, Resources, Formal analysis, Data curation. Daping Xie: Writing – review & editing, Software, Resources, Funding acquisition, Formal analysis, Data curation. Yufeng Zheng: Writing – review & editing, Supervision, Project administration, Conceptualization. Chunming Wang: Writing – review & editing, Supervision, Project administration, Conceptualization. Yunsong Liu: Writing – review & editing, Writing – original draft, Supervision. Dandan Xia: Writing – review & editing, Supervision, Project administration, Conceptualization.

Declaration of competing interest

Yufeng Zheng is an editor-in-chief for Bioactive Materials and was not involved in the editorial review or the decision to publish this article. Dandan Xia and Chunming Wang are editorial board members for Bioactive Materials and were not involved in the editorial review or the decision to publish this article. All authors declare that there are no competing interests.

Acknowledgments

This study was supported by the 10.13039/501100012166 National Key Research and Development Program of China (2023YFC2412600 ), the 10.13039/501100001809 National Natural Science Foundation of China (52271243 , 52171233 , 82370924 ), the 10.13039/501100002261 NSFC-RFBR Joint Research Scheme (82361138575 ), the 10.13039/501100005090 Beijing Nova Program (20230484459 ), the Beijing Natural Science Foundation (7242173 ), and the 10.13039/501100007937 Clinical Medicine Plus X-Young Scholars Project of Peking University , the Fundamental Research Funds for the Central Universities (PKU2024LCXQ014 ).

Peer review under responsibility of KeAi Communications Co., Ltd.
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