Biomolecule-functionalized dental implant surfaces: Towards augmenting soft tissue integration

Abstract

Dental implants are the primary solution for tooth replacement, providing both aesthetic and functional restoration. Their long-term success depends not only on osseointegration but also on robust peri-implant soft tissue integration (PSTI), particularly in the transmucosal region, where a stable epithelial seal is critical to preventing microbial infiltration and peri-implant inflammation. While surface topography modifications such as roughness, morphology, and porosity influence gingival cell behavior, passive surface modifications alone are often insufficient to promote rapid PSTI. This raises a fundamental question in dental implant design: How can implant surfaces be bioengineered to actively promote PSTI rather than passively relying on cellular responses? This review examines how biofunctionalization has emerged as a transformative strategy in implant surface engineering and critically analyses the latest biofunctionalization strategies for dental implants, with a particular focus on the underlying mechanisms that regulate biomolecule-implant interactions. It evaluates biomolecule incorporation via physical and covalent attachment, highlighting their distinct advantages in stability, efficiency, and scalability. We discuss approaches for functionalizing dental implant surfaces with bioactive molecules, such as proteins and peptides, and cells to replicate natural biological interactions, regulate immune responses, and enhance antimicrobial defense mechanisms. By addressing how bioengineered surfaces can be designed to actively engage with biological systems, this review provides a framework for developing next-generation implant technologies that achieve more effective and predictable PSTI, with strong potential for clinical translation.

Graphical abstract

Highlights

• •A critical and mechanistically focused review of biomolecule-functionalized dental implants, with specific emphasis on strategies to enhance soft tissue integration and epithelial sealing.
• •Systematically evaluates both physical and covalent biomolecule immobilization techniques, comparing their stability, scalability, and biological effectiveness for clinical translation.
• •Bridges biomaterials science and biology by dissecting the molecular mechanisms by which immobilized biomolecules modulate cell adhesion, immune responses, and inflammation resolution at the implant–tissue interface.
• •Identifies current limitations and overlooked challenges in long-term coating performance, biomolecule stability, and reproducibility of functionalization techniques under physiological conditions.
• •Outlines a roadmap for the development of next-generation bioactive dental implants with enhanced soft tissue integration, highlighting opportunities for multi-agent loading, stimuli-responsive surfaces, and scalable plasma-enabled fabrication strategies.

List of Abbreviations

Abbreviation	Definition
AA	Acorbic acid
AES	Auger Electron Spectroscopy
AFM	Atomic force microscopy
AI	Attachment index
AKT	Protein kinase B
Ala	Alanyl
AMBN	Ameloblastin
AMPs	Antimicrobial peptides
AO	Anodic deposition
AOD	Average optical density
A-PRF	Advanced platelet-rich fibrin
A-PRF+	Advanced Platelet-Rich Fibrin Plus
APTES	(3-Aminopropyl) triethoxysilane
Arg	Arginine
Ar-GDP	Glow discharge argon-based plasma treatment
ASCs	Adipose-derived stem cells
Asn	Asparagine
bFGF, FGF-2	basic Fibroblast growth factor
BMP-2	Bone Morphogenetic Protein type 2
BM-MSC	Bone marrow–derived mesenchymal stem cell(s)
BP180	Type XVII collagen
BP230	Bullous Pemphigoid Antigen 1
BSA	Bovine serum albumin
CAPs	Cell adhesion peptides
CDI	Carbonyldiimidazole
CD90	Cluster of differentiation 90
CD151	Cluster of differentiation 151
CD163	Cluster of differentiation 163
CD206+	Cluster of differentiation 206-positive
CFU	Colony-forming unit
CP	Commercially Pure
CSZP	Calcium–strontium–zinc–phosphate
C-terminal	Carboxyl-terminal
DAPI	Diamidino-2-Phenylindole
DEX	Dexamethasone
DNA	Deoxyribonucleic acid
DPSC(s)	Dental pulp stromal cell(s)
DSS	N,N′-disulphosuccinimidyl suberate
DPSCs	Dental Pulp stromal cells
ECM	Extracellular matrix
EDC	Ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride
EGF	Epidermal growth factor
EGFR	Epidermal growth factor receptor
ELISA	Electron Spectroscopy for Chemical Analysis
EMD	Enamel matrix derivative
ePTFE	expanded Polytetrafluoroethylene
ERK	Extracellular signal-regulated kinase
ERK1/2	Extracellular signal-regulated kinases 1 and 2
ESCA	Electron Spectroscopy for Chemical Analysis
F.I.R.S.T.	Fibrinogen-induced regeneration sealing technique
FAK	Focal adhesion kinase
FBS	Fetal bovine serum
FDA	Food and Drug Administration
Fg	Fibrinogen
FGF	Fibroblast growth factor
FN	Fibronectin
F-actin	Filamentous actin
GA	Glutaraldehyde
GBR	Guided bone regeneration
GelMA	Gelatin methacryloyl
GFBPt	Growth factor-binding peptides
GFOGER motif	Glycine-phenylalanine-hydroxyproline-glycine-glutamate-arginine
GFP	Green fluorescent protein
GFs	Growth factors
Gly	Glycyl
GMSCs	Gingival mesenchymal stem cells
GL13K	Gorr–L13K antimicrobial peptide
G2/M phase	Gap 2 and mitosis phase
H	Hydrogen
H2O2	Hydrogen peroxide
H2SO4	Sulfuric acid
HaCaT	Human epidermal keratinocytes
HBG	Horizontal bone gain
HBII	Heparin binding II
HD	Hemidesmosomes
HDFs	Human dermal fibroblasts
HEK293 cells	Immortalized human embryonic kidney cells
HFF	Human foreskin fibroblasts
HGF(s)	Human periodontal ligament fibroblasts
HPLF(s)	Human periodontal ligament fibroblast(s)
HSPGs	Heparan sulfate proteoglycans
hTGF-β	human Transforming growth factor-β
H&E	Hematoxylin and eosin
IKVAV	Ile-Lys-Val-Ala-Val
IL	Interleukin
IL - 1β	Interleukin - 1 beta
IL - 4	Interleukin - 4
IL - 6	Interleukin - 6
IL - 10	Interleukin - 10
IL - 11	Interleukin - 11
IL-12p40	Interleukin-12 subunit p40
IL - 13	Interleukin - 13
IL - 17	Interleukin - 17
IL -17A	Interleukin - 17A
IL - 23	Interleukin - 23
IL - 23R	Interleukin - 23 receptor
Ile	Isoleucyl
i-PRF	injectable Platelet-rich fibrin
ITAP	Porous titanium alloy flange
IκBα	inhibitor of kappa B alpha
K	Isoleucyl
KGF	Keratinocyte growth factor
KSL-W	KKVVFWVKFK
K23	KAFAKLAARLYRKALARQLGVAA
LAMA3	Laminin-332 α3 chain
LbL	Layer-by-layer
LDc	Liposome-DNA complexes
LG	Globular
Lip	Lipofectamine LTX
LN332	Laminin-332
Lys	Lysyl
MAPK	Mitogen-activated protein kinase
MC3T3-E1	MC3T3-E1 preosteoblast subclones
Mg2+	Magnesium(II) ion
MMP(s)	Matrix metalloproteinase(s)
MMP-1	Matrix metalloproteinase type 1
MMP-2	Matrix metalloproteinase type 2
MPF(s)	Multilayered polyelectrolyte film(s)
MSC(s)	Mesenchymal stem cell(s)
M1	pro-inflammatory
M2	pro-regenerative
N	Nitrogen
NHS	N-hydroxysuccinimide
NIH 3T3 fibroblasts	National Institutes of Health Swiss mouse embryonic fibroblast cell line clone 3T3
NOX4	NADPH oxidase 4
NP(s)	Nanoparticle(s)
OI	Orientation index
PBS	Phosphate-buffered saline
PCE	Polycarboxylates
PCL	Polycaprolactone
PDA	Polydopamine
PDGF	Platelet-derived growth factor
PDLLA	Poly‐d,l‐lactic acid
PEG	Polyethylene glycol
PEGylation	Polyethylene glycosylation
PGA	Propylene glycol alginate
PHSRN	Pro-His-Ser-Arg-Asn
PI3K-Akt	Phosphatidylinositol 3-kinase - protein kinase B
PIE	Peri-implant epithelium
PL	Platelet lysate
PLA	Polylactide
PLGA	Lactic-co-Glycolic Acid
PLL	Poly-L-lysin
PLLA	Poly-l-Lactic Acid
PR1P -	Proliferation-related ligand 1 peptide
prAMEL	Porcine recombinant amelogenin
PSTI	Peri-implant soft tissue integration
PTEN	Phosphatase and Tensin Homolog deleted on chromosome 10
PRDs	Platelet-rich derivatives
PRF	Platelet-rich fibrin
PRG4	Proteoglycan 4
PRHds	Platelet-rich hemoderivatives
PRP	Platelet-rich plasma
prTRAP	Porcine recombinant tyrosine-rich amelogenin peptide
RAD	Ac-(RADA)4-CONH2
RANTES	Regulated upon activation, normal t cell expressed and secreted
RGD	Arginine-glycine-aspartic acid
ROS	Reactive oxygen species
RT-PCR	Reverse transcription polymerase chain reaction
RT-qPCR	Reverse transcription quantitative polymerase chain reaction
SA	Sodium alendronate
SCAPs	Stem cells from apical papilla
SD	Sprague-Dawley
SDS	Sodium dodecyl sulfate
SEM	Scanning Electron Microscopy
Ser	Seryl
SF	Serum-free
SilMA	Methacrylate
SLA	Acid-etched
Sr2+	Strontium(II) ion
S phase	Synthesis phase
T cell(s)	Thymus-derived lymphocyte(s)
TCP	Tissue culture plastic
TGF-β	Transforming growth factor-β
TGF-β	Transforming growth factor-β
TGF-β1	Transforming growth factor-β1
TGF-β2	Transforming growth factor-β2
TGF-β3	Transforming growth factor-β3
Th17	T-helper 17 cells
Ti	Titanium
TiBP	Titanium-binding peptides
TIMPs	Tissue inhibitors of metalloproteinases
TIMP-1	Tissue Inhibitor of metalloproteinases-1
TIMP-2	Tissue Inhibitor of metalloproteinases-2
TiN	Titanium nitride
TiO2	Titanium oxide
TNF-α	Tumor necrosis factor-α
ToF-SIMS	Time-of-Flight Secondary Ion Mass Spectrometry
TRAP	Tyrosine-rich amelogenin peptide
Treg(s)	Regulatory T cell(s)
Tyr	Tyrosyl
Val	Valyl
VBG	Vertical bone gain
VEGF	Vascular endothelial growth factor
YIGSR	Tyr-Ile-Gly-Ser-Arg
ZA	Zoledronic acid
3D	Three dimensional
α-SMA	α-smooth muscle actin
–COOH	Carboxyl groups
–NH2	Amine groups
–OH	Hydroxyl groups
–SH	Sulfhydryl groups

1. Introduction

Dental implants are widely regarded as the gold standard for tooth replacement in partially or fully edentulous patients, offering functional durability and predictable aesthetic outcomes [1]. However, long-term implant success hinges on overcoming critical challenges, particularly in achieving robust hard and peri-implant soft tissue integration (PSTI) [2]. Hard tissue integration, or osseointegration, is essential to establishing a strong, stable interface with the surrounding bone, enabling the implant to endure mechanical forces without becoming loose over time. However, another equally critical objective is achieving peri-implant soft tissue integration (PSTI) to prevent bacterial infiltration and inflammation at the transmucosal region, where the implant traverses the gingival tissue [3,4] (Fig. 1a). Without effective soft tissue sealing, microbial pathogens and foreign particles can infiltrate the implant site, increasing the likelihood of peri-implant inflammation or infection (Fig. 1b). [5,6]. Consequently, recent advancements have prioritized strategies to enhance PSTI in the field of dental implantology.

Fig. 1.

PSTI .a) Dental implant and transmucosal region, b) Pre-impactites and implant failure resulting from poor PSTI, c) Surface biofunctionalization of dental implants using matrix proteins, peptides, cells, and GFs and d) Successful PSTI achieved on biofunctionalized surfaces (created using BioRender and Adobe Photoshop with permission).

PSTI is defined as the formation of a structurally stable and functionally competent seal between the peri-implant soft tissues and the transmucosal components of dental implants, most notably the abutment [4]. To improve the integration of soft tissues with hard tissue-interfacing implants, such as dental implants, various strategies have been implemented to modify the physical and/or chemical properties of the implant surfaces [[7], [8], [9], [10]]. Physical properties, including roughness [10], microstructure [7], and pore structure [11], are well-documented to influence the proliferation, shape, and adhesion strength of human gingival fibroblasts (HGFs). However, inconsistencies in the reported outcomes suggest that physical surface modifications alone may be insufficient to achieve optimal PSTI [10,12,13]. As such, modifying the chemical and/or biochemical features of the surface, or employing a combination of chemical, biochemical, and physical modifications, is increasingly recognized as a more effective approach to promote PSTI.

The surface chemistry of an implant is a key determinant of its interaction with surrounding soft tissues and can be tailored through various modification strategies [14,15]. The introduction of specific chemical functional groups or coatings to the surface can modulate interfacial properties such as surface charge and hydrophobicity, thereby influencing protein adsorption, cell receptor binding, and subsequent cellular responses [16,17]. Examples in this area include metals such as tantalum [18] and strontium [19], and polymers like chitosan and hyaluronic acid [20], which have all been employed in the form of coatings to enhance integration with surrounding soft tissues. Tantalum coatings on titanium (Ti) substrates, for instance, enhanced HGF adhesion, viability, proliferation, and migration, by promoting protein adsorption and activating cell-adhesive pathways mediated by integrin β1, vinculin, and focal adhesion kinase (FAK) [18]. A different approach to modifying surface chemistry involves implanting ions into the surface. For example, a recent study showed that magnesium and zinc ions, when introduced into Ti surfaces using plasma-immersion ion implantation technology, synergistically enhanced HGF motility and proliferation [21].

Surface biofunctionalization of dental implants has emerged as a distinct and promising strategy with the potential to transform PSTI [22]. Compared to conventional physical and chemical surface modification, biofunctionalization of implant surfaces increases functional versatility by attaching cells such as platelets [23], or bioactive molecules such as extracellular matrix proteins [9,24], growth factors (GFs) [25], and peptides [26] (Fig. 1c). The surface-bound molecules can be harnessed to promote natural biological processes at the implant surface to enhance PSTI, modulate immune responses, and mitigate inflammation; while certain agents, such as antimicrobial peptides (AMPs), provide inherent protection against bacterial colonization and biofilm formation, further supporting long-term implant success (Fig. 1d) [27,28]. An important advantage of biofunctionalization is its ability to mimic the native soft tissue microenvironment by presenting tissue-specific biochemical cues, spatially organized molecular signals, and matrix-like structural features [29,30]. Unlike traditional surface modifications that are primarily passive, biofunctionalized surfaces actively interact with cells by engaging specific receptors through precisely arranged biological ligands [31]. By offering spatiotemporal regulation of cell–surface interactions, bioactive coatings help direct consistent, pro-regenerative cellular responses, thereby enhancing the biological performance of implant surfaces.

The attachment of biological agents to implant surfaces can be achieved through simple physical adsorption [32], covalent attachment on chemically modified surfaces [33], or a combination of both [34]. Physical adsorption, such as dip-coating in a solution, is fast and cost-effective, but relies on weaker forces, such as electrostatic interactions and hydrophobic effects, which are prone to leaching and limit long-term stability. In contrast, covalent attachment enables the formation of robust and durable coatings of biomolecules, providing greater stability for PSTI [35]. The design and selection of appropriate technologies and protocols are particularly important for tailoring biofunctional dental implant surfaces to ensure compatibility with the implant material, biological agent, and intended application.

This review focuses on strategies and bioactive agents designed to enhance PSTI through the biofunctionalization of dental implant surfaces, highlighting both physical adsorption and covalent immobilization approaches. It discusses the role of matrix proteins, GFs, peptides, and cells in modulating implant-soft tissue interactions and assesses their potential for clinical translation. The concept of multi-agent loading is highlighted, emphasizing the synergistic effects of combining biomolecules to optimize biological responses. Furthermore, the review provides mechanistic insights into how these biomolecules promote attachment, cellular interactions, and epithelial sealing at the implant-soft tissue interface, offering pathways to more effective and durable integration. While these strategies have significantly advanced implant bioengineering, key challenges remain, including biomolecule stability, long-term coating performance, and achieving controlled biological responses at the implant interface. In this context, this review identifies major future research directions in the field with the potential to advance dental implant surface bioengineering. These include scalable fabrication technologies, advanced biomolecule delivery systems, and the integration of stimuli-responsive coatings, paving the way for next-generation dental implants capable of rapid and robust PSTI.

2. Cellular, structural, and molecular dynamics of peri-implant soft tissue healing: comparison with the natural periodontal interface

The surgical placement of a dental implant triggers a temporally regulated cascade of cellular events, tissue remodeling, and biological maturation that collectively guide PSTI. This complex process is driven by coordinated cellular and molecular mechanisms that respond to the initial surgical injury and promote wound healing.

Following implant placement, healing at the abutment–mucosa interface proceeds through a precisely coordinated sequence of cellular responses. Platelets arrive within minutes, forming a fibrin-rich clot that seals the wound and releases GFs such as Platelet-derived growth factor (PDGF), Transforming growth factor-β (TGF-β), and Vascular endothelial growth factor (VEGF) to initiate repair. Neutrophils infiltrate the site within hours, serving as the first line of defense by clearing pathogens and cellular debris. Around 24 h post-injury, monocytes are recruited and differentiate into macrophages, which initially adopt a pro-inflammatory (M1) phenotype before transitioning to a pro-regenerative (M2) state that facilitates tissue repair. Fibroblasts are recruited between days 2 and 4, synthesizing extracellular matrix proteins, particularly collagen, and contributing to granulation tissue formation. Between days 4 and 7, some fibroblasts transform into myofibroblasts, which promote wound contraction and matrix remodeling. Simultaneously, endothelial cells initiate angiogenesis between days 3 and 5, re-establishing vascular networks within the healing tissue. Keratinocytes begin migrating from the wound margins by days 2–3, with complete re-epithelialization typically achieved by day 7–14. Lymphocytes, including Thymus-derived lymphocytes (T cells) and Bone marrow-derived lymphocytes (B cells), appear later, modulating the immune response and supporting sustained tissue regeneration, particularly under microbial stress. By days 7–14, ongoing epithelial layering and matrix remodeling result in the formation of a mature mucosal barrier. The stratified squamous epithelium attaches to the implant surface through hemidesmosomal complexes and an internal basal lamina, supported by laminin and specific collagen isoforms. Beneath this layer, the connective tissue organizes into a dense collagen matrix that surrounds the implant abutment [4,36].

Unlike natural teeth—where collagen fibers insert perpendicularly into the cementum, a mineralized layer covering the tooth root and serving as the anchoring site for periodontal ligament fibers—the soft tissue surrounding implants contains collagen fibers arranged parallel or circumferential to the implant surface (Fig. 2a). This architectural difference results in a weaker biological seal, limited vascular supply, and reduced immune surveillance, increasing the risk of microbial colonization, inflammation, and peri-implant disease. The absence of proprioceptive innervation further restricts the tissue's ability to adapt to occlusal forces. As a result, the peri-implant mucosa forms a biologically active but structurally constrained interface [36].

Fig. 2.

Differences in soft tissue behavior around teeth and implants, with associated molecular interactions. a) Schematic illustration of the PSTI interface in a natural tooth versus a dental implant. In natural teeth, collagen fibers insert perpendicularly into the root surface, forming a robust connective tissue attachment. In contrast, at the implant–mucosa interface, collagen fibers are oriented parallel to the implant surface and lack insertion, leading to a mechanically weaker interface. b) Magnified insets depict key molecular interactions at the epithelial and connective tissue levels. Epithelial attachment to the implant is mediated by hemidesmosomal complexes involving integrin α6β4, LN332, and collagen XVII. In the underlying connective tissue, fibroblasts interact with the ECM through integrins α3β1 and associated focal adhesion proteins (talin, vinculin, paxillin), without direct collagen anchorage to the implant surface.

At the molecular interface, epithelial adhesion is mediated by hemidesmosomal complexes and a specialized internal basal lamina as shown in Fig. 2b. Laminin-332 (LN332) binds to α6β4 integrins, initiating hemidesmosome assembly, which is anchored intracellularly by adaptor proteins such as BP180 (Type XVII collagen), BP230, and plectin [37]. Tetraspanin cluster of Differentiation 151 (CD151), also present at this interface, stabilizes α6β4 integrin clustering and facilitates hemidesmosome formation, further strengthening epithelial attachment [38,39]. In the underlying connective tissue, gingival fibroblasts produce extracellular matrix (ECM) components—including collagen types I, III, and V, fibronectin (FN), and laminin—and anchor to the implant surface through integrin-mediated focal adhesions (α3 and β1). These adhesions are supported by cytoskeletal linkers such as talin, paxillin, and vinculin, which activate PI3K-Akt and mitogen-activated protein kinase (MAPK) signaling pathways to regulate cell survival, ECM production, and cytoskeletal organization [38,40].

The Phosphatidylinositol 3-kinase - protein kinase B (PI3K-Akt) pathway promotes cell survival, migration, and ECM production by activating Akt (protein kinase B), which prevents apoptosis and supports cytoskeletal reorganization. PI3K-Akt pathway is essential for sustaining fibroblast activity and ensuring stable attachment to the implant surface [41,42]. The MAPK pathway is another key signaling cascade involved in tissue integration. It consists of three main subfamilies: ERK1/2 (extracellular signal-regulated kinases 1 and 2), JNK (c-Jun N-terminal kinase), and p38 MAPK [43]. Each responds to different environmental cues and regulates distinct aspects of cell behavior. ERK1/2 is primarily activated by GFs and promotes fibroblast proliferation and ECM synthesis [44]. In contrast, p38 MAPK and JNK are activated by stress and inflammatory signals, contributing to wound healing, inflammatory regulation, and adaptation to cellular stress [45]. The PI3K-Akt and MAPK pathways engage in complex cross-regulatory interactions that influence cellular responses in a context-dependent manner. Under certain conditions, activation of PI3K-Akt can enhance MAPK signaling, for example MAPK signaling, particularly via ERK1/2 or p38 MAPK, can inhibit PI3K-Akt activity through negative feedback loops or via activation of phosphatases such asphosphatase and tensin homolog deleted on chromosome 10 (PTEN) [[46], [47], [48]]. The bidirectional crosstalk enables cells to fine-tune key processes such as proliferation, survival, and differentiation in response to extracellular cues, and has important implications for tissue healing and implant integration.

3. Strategies for loading biological agents onto the surface

Surface biofunctionalization involves attaching biological agents to implant surfaces, using either physical methods, chemical methods, or a combination of both. Physical approaches rely on non-covalent interactions, such as electrostatic forces and surface entrapment, while chemical methods utilize covalent bonding to ensure stability and durability. This section reviews and discusses the key strategies for loading biomolecules, including direct physical attachment, direct covalent attachment, and indirect methods involving the encapsulation of biological agents.

3.1. Direct physical attachment of biomolecules

Physical attachment of biomolecules to implant surfaces relies on non-covalent interactions as well as their entrapment within surface pores and roughness features, as illustrated in Fig. 3 [49]. These interactions include van der Waals forces, electrostatic interactions, hydrogen bonding, ionic interactions, and hydrophobic interactions, each contributing to biomolecule adhesion with varying stability.

Fig. 3.

Physical, non-covalent biomolecule attachment to dental implants.a) Van der Waals forces, b) Electrostatic interactions, c) Hydrogen bonding, d) Ionic interactions, e) Hydrophobic interactions, and f) Physical entrapment within surface features (created using BioRender with permission).

Van der Waals forces are weak, short-range interactions that arise from transient dipoles between molecules (Fig. 3a) [50,51]. Electrostatic interactions (Fig. 3b) occur between oppositely-charged regions of biomolecules and surfaces [52]. Hydrogen bonds (Fig. 3c) form between hydrogen atoms and electronegative atoms, while ionic interactions (Fig. 3d) involve attractions between oppositely-charged ions or molecular domains. Hydrophobic interactions (Fig. 3e) drive the aggregation of non-polar biomolecular regions to minimize contact with water, promoting adsorption on hydrophobic surfaces.

The stability and efficacy of these physical interactions are highly dependent on environmental factors such as pH, ionic strength, and solvent composition. If not carefully controlled, variations in such conditions can disrupt biomolecule attachment, leading to desorption and reduced bioactivity at the implant interface [[53], [54], [55], [56], [57], [58], [59]].

Several techniques utilize physical forces to achieve surface attachment of biomolecules. One example is dip coating, a widely used and cost-effective method [[60], [61], [62]]. In this approach, the biomaterial is immersed in a solution containing biological reagents, enabling the formation of non-covalent bonds or the entrapment of biomolecules within surface features [9]. Another well-established physical method, layer-by-layer (LbL) assembly, involves sequentially adsorbing oppositely charged materials to create multi-layered coatings. This technique provides precise control over coating thickness and facilitates the controlled release of biological agents [63,64].

Pre-surface modification of dental implants using techniques such as sandblasting [65], acid etching [66,67], electrochemical anodization [68,69], and plasma treatments [[70], [71], [72]], enhance adsorption and adhesion by increasing surface roughness and enlarging the available surface area [15] (Fig. 3f). This creates more binding sites, improves mechanical interlocking, and supports the physical entrapment of biological agents, enhancing their stability and bioactivity [73,74]. Surface roughness, in particular, has been widely reported to influence protein adsorption by increasing the surface area and introducing topographical features for improved physical entrapment [75]. However, the effects of roughness extend beyond the surface area alone and vary depending on the properties of the proteins and biomolecule coatings. For instance, anisotropic proteins like fibrinogen (Fg), with elongated and uneven dimensions, may change their adsorption orientation on rough surfaces, shifting from a "side-on" (flat) to an "end-on" (upright) configuration, which allows for higher adsorption density [75]. In contrast, globular proteins like bovine serum albumin, which are compact and spherical, are less impacted by roughness as their adsorption primarily depends on surface chemistry rather than orientation [75]. Nanoscale grooves, meanwhile, can trap water molecules, forming a hydrated layer that further influences adsorption behavior [75]. Such a complex interplay between surface roughness and protein binding characteristics highlights the necessity of customizing surface modifications and employing alternative techniques to enhance biomolecule attachment efficiency and consistency.

Physical attachment methods offer the advantage of simplicity for rapid and cost-effective biomolecule functionalization of implant surfaces. Despite this advantage, however, they face significant challenges due to the inherent weaknesses of non-covalent interactions. In complex biological environments, the physical bonds are prone to displacement, particularly due to the Vroman effect, wherein initially adsorbed proteins are replaced by those with a higher affinity [76]. Such displacement can alter the surface composition over time, potentially reducing bioactivity and leading to unpredictable biological responses [58].

To address these limitations, advanced loading strategies have been developed. One such approach involves the entrapment of biomolecules within protective carriers, such as hydrogels, microbeads, or liposomes, which can subsequently be loaded onto biomaterial surfaces using either physical or chemical methods [[77], [78], [79], [80]]. This strategy is discussed in detail in Section 2.3. Alternatively, direct covalent attachment of biomolecules to implant surfaces provides a more stable and durable solution, as discussed in the following section.

3.2. Direct covalent attachment of biomolecules

Chemical covalent attachment addresses the limitations of non-covalent interactions—such as rapid degradation, leaching, and displacement—by establishing stable and durable bonds between biomolecules and implant surfaces.

This process typically involves introducing surface functional groups that form specific chemical bonds with complementary groups on the biomolecule. Covalent attachment strategies can be broadly classified into wet chemistry and dry plasma-based methods. Wet-chemistry methods rely on solution-based reactions to introduce functional groups to the implant surface [[81], [82], [83], [84]], while dry plasma approaches use plasma-based surface modifications to achieve covalent attachment by introducing chemical functional groups or surface-embedded reactive radicals [58,[85], [86], [87], [88], [89], [90]].

Wet chemical methods facilitate the covalent attachment of biomolecules to implant surfaces by introducing specific functional groups that enable stable bonding. For instance, silanization techniques utilize alkylsilane linkers to introduce amine groups that can subsequently react with biomolecules to form strong covalent linkages (Fig. 4a). Another widely used method involves forming amide bonds between carboxyl (–COOH) and amine (–NH₂) groups, often mediated by carbodiimide reagents such as Ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDC), with N-hydroxysuccinimide (NHS) enhancing reaction efficiency and stability Fig. 4b [58]. Polyethylene glycosylation (PEGylation) is another major method that specifically uses Polyethylene glycol (PEG) chains to bind biomolecules to the implant surfaces covalently. Typically, it involves surface activation through plasma or chemical treatments to introduce reactive groups such as hydroxyl (–OH) groups, then covalently bonding with plasma or chemical-activated PEG [91]. As a result, this method can lead to a stable formation of the PEG network, with potential crosslinking to enhance attachment on implant surfaces, reduce immunogenicity [[92], [93], [94], [95]], and provide a protective layer for therapeutic agents [96], as illustrated in Fig. 4c. Another approach is based on Schiff base reactions forming an imine bond through the aldehyde or ketone interaction with a primary amine, which can be stabilized via reductive amination (Fig. 4d) [97,98]. These reactions are advantageous in creating stable linkages between biological agents and implant surfaces. Thiol-based reactions, using sulfhydryl groups (–SH), can also be used to form strong covalent bonds, such as disulfide or sulfide bonds. The reaction process for these bonds is depicted in Fig. 4e [99,100]. These strategies are especially critical in implant surface modifications and promote better tissue integration [15,101]. Collectively, these wet chemical strategies, which are typically multi-step and substrate-dependent, enable the stable and functional integration of biomolecules onto implant surfaces for longer-term performance.

Fig. 4.

Chemical immobilization strategies for biomolecule covalent attachment to dental implant surfaces.a) Silanization, b) EDC/NHS coupling reaction, c) PEGylation, d) Schiff-base reaction, and e) Thiol-based conjugation (created using BioRender with permission).

Dry plasma techniques, which either utilize non-depositing gases such as oxygen, nitrogen, and argon (i.e., plasma treatment), or depositing gases or vapors such as acetylene, allylamine, and acrylic acid (i.e., plasma polymerization), offer transformative potential for surface biofunctionalization [[102], [103], [104], [105]]. The chemical functional groups which can be generated on implant surfaces through a plasma surface modification approach include, but are not limited to, hydroxyl [[106], [107], [108]], carboxyl [[109], [110], [111]], thiol [112,113], and amine groups [109,114]. Energetic ion bombardment during plasma processes can also be harnessed to generate long-lived radicals embedded onto the surfaces, serving as reactive sites for one-step, covalent biomolecule immobilization [85,108,[115], [116], [117], [118], [119]]. The simplicity of one-step, reagent-free covalent attachment, combined with its versatility, makes this approach particularly attractive for the surface biofunctionalization of a wide range of implantable materials.

Each technique for covalent attachment presents distinct advantages and limitations, shaped by its underlying mechanisms and operational requirements. For instance, wet chemistry methods can be carried out with relatively low capital costs; however, they often face significant challenges in scalability during fabrication [120]. Wet chemical methods require precise control over reaction conditions, including temperature, pH, and solvent composition, to achieve uniform surface coverage and avoid the formation of undesired by-products [121,122]. They also typically involve multiple steps, making them time-consuming and less suitable for large-scale applications. Another significant concern is the generation or retention of toxic agents on implants, particularly porous ones, as their removal is often infeasible, creating substantial hurdles for regulatory approval. Also, the generation of chemical waste during processing raises environmental concerns and increases costs [85,108,119,123,124]. In contrast, dry plasma-based technologies are more environmentally friendly, as they do not generate chemical waste and eliminate the risk of toxic residues on implant surfaces. Plasma-based methods also involve fewer steps, simplifying the process and increasing efficiency to scale up manufacturing. However, relatively high capital costs for the plasma processes may be challenging due to the requirement for large vacuum chambers in the case of low-pressure plasma processes, and the difficulty in achieving uniform plasma generation, especially for complex geometries. The capital costs, however, can be offset by low ongoing expenses, as plasma techniques require minimal reagents, generate almost zero chemical waste, and eliminate the need for waste management. The dry plasma processes can also be seamlessly integrated into existing manufacturing lines, allowing for automation and further reducing operational costs.

3.3. Indirect attachment using encapsulation of biological agents

Indirect attachment of biological agents involves encapsulating them within protective carriers, such as nanoparticles (NPs) [125], liposomes [126], and microparticles [127,128], which act as reservoirs and barriers for sensitive bioactive molecules [129]. This strategy protects the encapsulated molecules from adverse physiological conditions, including pH fluctuations, enzymatic degradation, and oxidative stress, ensuring their stability and functionality for eliciting the required biological responses.

One key benefit of encapsulation is that it allows for controlled and sustained release of biological agents. This is particularly advantageous for applications like tissue engineering and drug delivery, where prolonged and localized activity is essential for success [129]. Furthermore, encapsulation mitigates surface-related challenges often seen in direct attachment methods, such as uneven biomolecule distribution, competitive protein displacement (e.g., the Vroman effect) [130], and bioactivity loss due to non-native orientation. Encapsulation can also enhance biocompatibility by employing matrices such as NPs [129], microparticles [131], hydrogels [132], or liposomes [133,134] (Fig. 5), which can be engineered to minimize immune reactions and promote integration with surrounding tissues. Another advantage lies in its ability to maintain biomolecule activity in dynamic biological environments that contain potentially degradative bodily fluids, enzymes, and cells. For instance, functionalized microporous scaffolds embedded with growth factor-encapsulated NPs were specifically engineered to overcome the short half-life, limited stability, and enzymatic degradation of GFs in physiological conditions [12].

Fig. 5.

Indirect attachment of biomolecules using carrier-based delivery methods. NP loading, Liposome encapsulation, Microparticle embedding, and Hydrogel coating immobilization (created using BioRender and Adobe Photoshop with permission).

In addition to the above functions, encapsulation can integrate multiple biological agents —such as GFs, antibiotics, or signaling molecules—within a single carrier or distinct carriers to achieve sequential or multiple combined therapeutic effects [135]. For instance, a NP-embedded electrospun nanofiber scaffold was developed for the controlled dual delivery of Bone Morphogenetic Protein type 2 (BMP-2) and dexamethasone (DEX). BMP-2 was encapsulated in bovine serum albumin (BSA) NPs to maintain its bioactivity, while the scaffold was fabricated by co-electrospinning with a PCE copolymer. In vitro studies confirmed preserved bioactivity and sequential release, with DEX released within eight days and BMP-2 sustained for up to 35 days. Owing to the sustained and bioactive delivery of DEX and BMP-2, both in vitro and in vivo osteogenesis studies demonstrated significantly improved bone regeneration, with the dual-drug-loaded scaffold exhibiting markedly greater efficacy compared to the blank scaffold [135].

Hydrogels can also be utilized to encapsulate and load biomolecules onto implant surfaces. However, a significant challenge in this approach is achieving strong adhesion between soft hydrogels and hard implant materials. Strong adhesion can be achieved through physical or chemical modifications of the hydrogel, the implant surface, or both. For instance, a silica xerogel and chitosan hybrid coating has been applied to Ti surfaces to deliver fibroblast growth factor (FGF) [136]. Chitosan provided mechanical stability and controlled release of FGF, significantly enhancing osteoblast cell responses. In another study, fibrin-collagen hydrogels containing human adipose-derived stem cells (ASCs) were used to coat cochlear implant electrode arrays, showing promise for enhancing the bioelectric interface of CIs by improving cell viability and promoting neurotrophic factor release [137].

4. Protein-biofunctionalized surfaces

Proteins are macromolecules composed of amino acids arranged in specific sequences, which determine their structure and function. They play important roles in biological processes, including cellular adhesion, signaling, and tissue formation [138]. ECM proteins like collagen, FN, laminin, and Fg interact with cells through receptors such as integrins, facilitating cell attachment, migration, and differentiation [139].

Coating implant materials with mono- or multi-layers of proteins can be applied as a strategy to create a biomimetic surface, enhancing compatibility and integration with surrounding soft tissues [140]. Protein-functionalized implant surfaces can be designed to minimize immune rejection and inflammation, thereby enhancing implant integration and acceptance [141]. Such coatings are of particular interest for promoting fibroblast adhesion, an important factor for PSTI [142,143]. Cell adhesion is fundamental to tissue integration, providing the foundation for implant incorporation and mechanical stability. Insufficient cell adhesion compromises implant integration, increasing the risk of implant failure [144]. Beyond securing the implant, adhesion regulates intracellular signaling, promotes tissue repair, and establishes a protective barrier against infection—imperative factors for long-term implant success [145].

This section first examines the mechanisms regulating cell interactions with protein-functionalized surfaces, followed by an analysis of the proteins employed for surface functionalization to promote PSTI on dental implants.

4.1. Interactions between cells and ECM protein-coated implant surfaces

Integrins, transmembrane proteins found on the surface of cells like fibroblasts, primarily facilitate cell adhesion to ECM-coated implants. By binding to specific sequences, such as the Arginine-glycine-aspartic acid (RGD) [146] or the Glycine-phenylalanine-hydroxyproline-glycine-glutamate-arginine (GFOGER motifs) [147], integrins enable strong cell attachment, laying the foundation for successful tissue integration [148].

When cells interact with the ECM-coated surfaces, they first extend small projections called filopodia and lamellipodia, which help them sense their surrounding environment and establish a stable connection with the surface [149]. Filopodia are thin, spike-like projections that sense the environment and make the first points of contact with the surface, while lamellipodia are broad, sheet-like extensions that stabilize and strengthen cell attachment [150]. As filopodia and lamellipodia engage with ECM-derived coatings, they promote the formation of focal adhesions, which are intricate protein complexes that include integrins and other cytoskeletal components to anchor the cell firmly to the substrate [151]. The focal adhesions serve as anchor points that connect the integrins to the cell cytoskeleton, establishing a firm link between the cell and the surface coating. The focal adhesions can also initiate intracellular signaling pathways regulating cell survival, proliferation, and differentiation, which are important for optimal tissue integration [152,153]. The signaling cascade created in this way supports cellular functions, encouraging fibroblasts and other relevant cells to proliferate and spread across the implant surface, ultimately forming a cellular layer that enhances tissue integration.

Cells adhering to the ECM-coated surfaces can also upregulate endogenous production of ECM proteins, such as collagen and FN [154]. The ECM formed in this manner mirrors natural tissue architecture, establishing a stable connective interface around the implant and reinforcing the bond with the surrounding soft tissues [155]. Hence, through sequential processes, including filopodia and lamellipodia extension, focal adhesion formation, cellular proliferation, and ECM production—matrix protein-coated surfaces can create a bioactive environment that facilitates effective PSTI.

The proteins employed in the biofunctionalization of dental implants to promote PSTI are primarily ECM proteins, including collagen, FN, laminin, Fg and fibrin [156]. These proteins have been shown to support early wound healing and facilitate PSTI [156,157]. An overview of these proteins, along with their applications, substrate materials, and key findings, is presented in Table 1.

Table 1.

Summary of proteins used for surface modification of dental implants for enhanced PSTI.

Protein	Substrate	Method of immobilization	Key findings	Ref
Collagen	Ti and TiN	Dip coating or advanced methods such as anodic deposition (AO) and crosslinking.	Collagen dip coated on Ti and Titanium nitride (TiN) enhanced cell proliferation and adhesion, comparable or superior to RGD or no protein coating; TiN had no significant effect on HGF.	[32]
Collagen	Ti	Electrochemical method	No significant differences in soft tissue dimensions or connective tissue composition was observed compared to uncoated implants in dog model.	[158]
FN	Ti	Microgrooving through photolithography, acid etching, silanization, and FN coating	Enhanced fibroblast adhesion, spreading, and proliferation, demonstrating a fivefold increase in fibroblast proliferation compared to smooth Ti surfaces.	[159]
			Increased vinculin expression and enhanced cell alignment, spreading, and proliferation.	[160]
FN	Ti	Silanization, and UV-activated anthraquinone linker attachment	Improved keratinocyte adhesion and spreading, reduced platelet adhesion and clot formation, mitigated inflammation and infection risks.	[161]
FN	Y-TZP (Zirconia(ZrO2))	Silanization via Tresyl chloride–activated immobilization	Increased vertically oriented collagen fibbers on the implant surface, enhancing gingival fibber alignment and soft tissue response.	[162]
FN	Ti	Argon-based glow discharge plasma treatment followed by FN grafting	Significantly improved both hard and PSTI in vivo, with enhanced hydrophilicity and roughness facilitating better cell adhesion and differentiation.	[163]
Modified heparin binding II (HBII)-RGD domain of FN	Ti	(3-Aminopropyl)triethoxysilane (APTES)	Enhanced fibroblast adhesion, spreading, proliferation, migration, and activation.	[164]
FN	Porous Ti Alloy Flange (ITAP)	The FN was combined with the hydroxyapatite during the electrochemical coating process.	Reduced epithelial downgrowth and increased soft-tissue integration compared with drilled flanges, but no additional benefit from coatings.	[165]
LN332	Ti	LBL-like assembly using PDLLA as a carrier for controlled protein release	Improved keratinocyte adhesion, migration, and proliferation through PI3K-Akt pathway activation and promoted gingival mesenchymal stem cells (GMSC) epithelial differentiation for enhanced soft tissue sealing.	[37]
LN332	Ti	Chemical activation with piranha solution, silanization (APTES + Glutaraldehyde (GA)), and covalent bonding of LN332	Improved keratinocyte adhesion, cytokine secretion, and tissue repair-related gene expression without cytotoxicity.	[34]
LN332	Ti-6Al-4V	Passivation with ultrapure nitric acid followed by laminin coating	Significantly improved keratinocyte attachment, spreading, and hemidesmosome formation due to changes in laminin conformation, enhanced cell-binding domain accessibility, and reduced ion release via a stable Titanium oxide (TiO2) TiO2 layer.	[166]
Fg	Ti	Polydopamine (PDA) pre-coating followed by covalent attachment of Fg	Improved both soft and hard tissue cell adhesion, reduced epithelial downgrowth, minimized fibrous capsule formation, and enhanced angiogenesis.	[167]

4.2. Collagen-biofunctionalized surfaces

Collagen, the predominant structural protein in the ECM of connective tissues, is widely distributed throughout the body [168]. Its natural biocompatibility and abundance make collagen ideal for medical applications, including implants and tissue engineering scaffolds [169].

Collagen coatings can be applied on Ti dental implants to enhance PSTI. For example, in one of the early works in this area, it was shown that when human HGFs were cultured on both collagen-coated and uncoated Ti substrates, the number of adherent cells on collagen-coated Ti was significantly greater than on non-coated Ti (Fig. 6a). Although collagen coatings had no effect on cell adhesion on polystyrene (∼570 cells/cm²), it nearly doubled cell adhesion on Ti from ∼450 to ∼930 cells/cm², highlighting its effectiveness on Ti surfaces and demonstrating its substrate dependency. Collagen coatings can also influence the shape of attached cells, promoting a flattened and spread morphology with well-defined lamellipodia and filopodia (Fig. 6b and c) [24].

Fig. 6.

Collagen biofunctionalized surfaces for PSTI. a) Number of cells adhered to non-coated or collagen-coated Ti and tissue-culture polystyrene. A significant increase in adherent HGFs on collagen-coated Ti compared to non-coated Ti was observed. b, c) SEM images of HGFs on uncoated and collagen-coated Ti show that HGFs on collagen-coated Ti exhibit a more spread morphology with well-defined lamellipodia and filopodia, indicating improved cell attachment and interaction with the surface [24], d) HGFs morphology on thick and thin collagen-coated Ti and TiN surfaces (scale bar: 20 μm) shows enhanced adhesion, focal adhesions, and alignment with thinner coatings, while thicker coatings reduce cell alignment and substrate interaction [32], e) Cell viability and proliferation assay on days 3 and 7 shows enhanced proliferation on collagen and RGD-coated Ti compared to unmodified Ti. Reproduced from Ref. [32], with permission from Springer Nature. © 2016, Springer-Verlag Berlin Heidelberg. f) Left: Ground section of a control implant and the surrounding soft and hard tissues after 8 weeks of healing, original magnification × 25. Right: Enlarged view of the left image under polarized light, showing collagen fibers oriented in different directions, original magnification × 200. g) Left: Ground section of a test implant and the surrounding soft and hard tissues after 8 weeks of healing, original magnification × 25. Right: Enlarged view of the left image under polarized light, showing collagen fibers oriented toward the implant surface, original magnification × 200. Reproduced from Ref. [158] with permission from John Wiley and Sons. © John Wiley & Sons A/S. Published by John Wiley & Sons Ltd.

The thickness of collagen coatings is an important factor in modulating soft tissue cellular responses to implant surfaces and must be carefully optimized [32]. Ritz et al. investigated how human HGFs responded to surface modifications of Ti and Ti nitride (TiN) alloys coated with type I collagen [32]. The study aimed to compare the effects of simple dip coating of collagen with other methods, such as anodic deposition (AO) and crosslinking, focusing on the role of coating thickness in modulating cellular behavior. The results revealed that thinner collagen coatings substantially enhanced soft tissue cellular interactions, with HGFs exhibiting stronger adhesion, more distinct focal adhesions, and alignment closely following the surface microstructure (Fig. 6d). Conversely, thicker collagen coatings diminished cell alignment and reduced interaction with the underlying substrate (Fig. 6d). For cell viability and proliferation, all modifications, including collagen and RGD coatings, demonstrated improved proliferation rates compared to unmodified controls, with the highest increases observed at days 3 and 7 (Fig. 6e) [32].

The authors attributed the observed cell responses, including altered orientation and adhesion on thicker collagen coatings, to the masking effect of the dense collagen layers, which conceal the underlying surface topography. The masking effect reduces the exposure of surface microstructures, leading to less effective cell adhesion and altered cell behavior. In contrast, thinner collagen coatings can preserve these features, promoting better cell adhesion and orientation [32]. However, another potential reason, not discussed in the paper, could include steric hindrance [170]. Steric hindrance in cell attachment is a physical barrier that blocks cells from adhering to a surface by preventing integrin receptors from reaching binding sites [170]. Large or densely packed molecules or layers on the surface can create this obstruction, hindering stable cell contact. In contrast, thinner collagen coatings may provide an optimal density of attachment sites (such as RGD peptides), balancing cell adhesion without overcrowding ligands [171]. Excessive attachment points can disrupt cell signaling, resulting in less efficient cell attachment and proliferation. Therefore, accurately measuring coating thickness using techniques such as as Atomic force microscopy (AFM) or ellipsometry can provide detailed insights into the coating's physical dimensions and surface characteristics, allowing optimization of ideal coating thickness for target cells.

The varying responses of HGFs to different collagen-coating thicknesses may also provide insight into why collagen-coated implants failed to enhance PSTI in a dog model [158]. In this study, type I collagen-coated Ti implants were surgically placed in the mandibular regions of dogs, but no significant differences were observed in soft tissue dimensions or connective tissue composition between collagen-coated and uncoated implants at 4- and 8-weeks post-implantation (Fig. 6f and g). Both groups exhibited similar tissue characteristics, including low blood vessel density and parallel collagen fiber orientation, with no evidence of improved attachment for the collagen-coated implants [158]. This lack of improvement may be attributed to the use of a single collagen coating thickness, applied using an electrochemical method, without evaluating the effects of varying thicknesses. Exploring a broader range of collagen coating thicknesses may yield different and more favorable outcomes.

Although the role of collagen-biofunctionalized surfaces in promoting PSTI has not been comprehensively elucidated, emerging evidence suggests that their efficacy extends beyond simple considerations of coating thickness. Instead, key determinants include the specific type of collagen employed, its molecular conformation, and the immobilization strategy used to anchor it to the substrate. For example, type I collagen, the most prevalent form in the human body, is abundantly expressed in skin, ligaments, tendons, and teeth, where it provides tensile strength and contributes to ECM stability. Owing to its robust structural properties and biocompatibility, it is widely used in wound dressings and implant coatings [172]. By contrast, type IV collagen is a principal component of basement membranes in epithelial tissues, where it mediates cell adhesion and barrier function [173]. Given these distinct physiological roles, selective incorporation of different collagen types into implant coatings may differentially modulate cellular behavior.

Beyond collagen type, the molecular conformation of collagen is critical in determining its biological performance. The optimal conformation for biomaterial coatings depends on the intended application and desired cellular responses. However, preserving the native triple-helical structure is generally beneficial, as it enhances cell adhesion, proliferation, and tissue integration. For instance, when electrospun polymer fibers were coated with native triple-helical collagen and its acidic derivatives via a LBL method, the resulting surface facilitated strong attachment and spreading of NIH 3T3 fibroblasts (National Institutes of Health 3-day transfer, inoculum 3 × 10⁵ cells), underscoring the importance of maintaining the triple-helical conformation to ensure optimal bioactivity [174].

Equally important is the method of collagen immobilization. Although yet not been studied for PSTI, comparative studies have demonstrated that covalent attachment of collagen to biomaterial surfaces yields superior outcomes relative to physical adsorption. In one study, collagen was immobilized onto hydroxyapatite (HA) discs either covalently (HA-C) or via adsorption (HA-hC). Covalently bonded collagen resulted in greater surface retention, higher alkaline phosphatase activity, and enhanced osteoblastic differentiation of MC3T3-E1 preosteoblast subclones (MC3T3-E1) cells, while physically adsorbed coatings showed limited efficacy [175]. In a more complex system, Müller et al. engineered biofunctional stainless steel substrates by first applying a corrosion-resistant tantalum oxide layer, followed by covalent immobilization of type I collagen. This was achieved through surface functionalization with 3-aminopropyltriethoxysilane (APTES), crosslinking with N,N′-disulphosuccinimidyl suberate (DSS), and further stabilization using carbodiimide-induced crosslinking. Physically adsorbed collagen coatings served as controls. The covalently immobilized layers exhibited superior mechanical integrity and enzymatic resistance. Subsequent in vitro culture with human mesenchymal stem cells and in vivo subcutaneous implantation in nude mice confirmed enhanced cytocompatibility and tissue integration relative to physically adsorbed counterparts [176].

Together, these findings highlight the importance of integrating collagen type, molecular conformation, and immobilization chemistry when designing biofunctionalized surfaces for dental implants. A rational, multi-parameter approach is required to optimize the bioactivity, stability, and integration capacity of collagen-based coatings for soft tissue applications.

4.3. Fibronectin (FN)- biofunctionalized surfaces

FN is a large glycoprotein in the ECM, integral to processes such as cell adhesion, migration, and wound healing [177]. With its multiple binding sites, FN interacts with integrins on cell surfaces, facilitating cell attachment and initiating signaling pathways [178]. In biomedical applications, FN is frequently applied to implant surfaces to enhance tissue integration [179]. By promoting cell adhesion and activating key signaling mechanisms, FN-coated surfaces support the attachment and proliferation of cells such as fibroblasts and osteoblasts, accelerating both soft and hard tissue integration, reducing healing times, and improving implant durability [[179], [180], [181]].

The potential of FN-biofunctionalized surfaces to enhance PSTI in dental implants has been demonstrated in several studies [[159], [160], [161],163]. Kim et al. investigated FN-immobilized micro-grooved surfaces and their effects on HGF proliferation, gene expression, and protein expression [159]. Ti surfaces were prepared by microgrooving through photolithography, followed by acid etching, silanization, and FN coating resulting in FN-immobilized micro-grooved Ti. The rationale behind combining microgrooving and FN functionalization lies in their complementary ability to improve cell behavior on Ti surfaces. Microgrooves provide a structured topography that directs fibroblast alignment and movement through contact guidance, while FN functionalization introduces biomechanical signals that simulate the ECM, enhancing cell adhesion and signaling. The study revealed that the most pronounced adhesion and spreading of fibroblasts were observed on microgrooved Ti with FN immobilization (E60/10FN in Fig. 7a) in comparison to smooth (NE0), and FN-immobilized smooth Ti (NE0FN). Furthermore, the proliferation rate on micro-grooved Ti (E60/10) without FN was approximately twofold higher than that on smooth Ti (NE0 in Fig. 7b). However, FN-immobilized micro-grooved Ti (E60/10FN) further enhanced fibroblast proliferation, demonstrating a fivefold increase after 96 h compared to the smooth Ti control, which shows the significant synergistic effect of combining physical and biochemical surface modifications in promoting cellular responses [159].

Fig. 7.

FN biofunctionalized surfaces for PSTI. a) Confocal images of HGFs after 16 h on NE0, NE0FN, and E60/10FN surfaces, showing vinculin (green) and nuclei (blue) for each surface type. Scale bar: 50 μm. b) HGFs proliferation on NE0, E0, E60/10, NE0FN, E0FN, E60/10FN surfaces at 48, 72, and 96 h. E60/10FN showed the highest adhesion and a fivefold proliferation increase at 96 h, confirming the synergy of topography and biochemical modification. Reproduced fromRef. [159]with permission from Springer Nature. © 2018, The Korean Tissue Engineering and Regenerative Medicine Society and Springer Science Business Media B.V., part of Springer Nature. c) Immunofluorescence staining shows FN attachment on different Ti surfaces, with the highest intensity on silanized microgrooved Ti (M + APTES), indicating enhanced FN loading via covalent bonding. d) Vinculin staining and quantitative analysis show that HDFs on microgrooved surfaces align along grooves, while covalent FN loading via silanization further enhances cell adhesion and proliferation, confirming the synergistic effects of surface topography and biochemical modification. Scale bar: 100 μm [160]. e) Keratinocyte adhesion and spreading after 2 h on untreated Ti (left) and anthraquinone/GOPTS-linked FN-coated Ti (right), visualized with fluorescein diacetate staining. The coating doubled keratinocyte adhesion and spreading. f) Platelet receptor GPIIb/IIIa measured via CD41 absorbance after 60-min PRP incubation. The FN coating reduced platelet adhesion and clot formation, minimizing inflammation compared to unmodified Ti. Reproduced from Ref. [161] with permission from Elsevier. © 2006 Academy of Dental Materials. Published by Elsevier Ltd. All rights reserved.

In another study, the impact of covalent or physical loading of FN on smooth and microgrooved Ti surfaces was examined, with a particular emphasis on the influence of silanization and its effects on HGF behavior [160]. The samples were microgrooved, treated with H₂SO₄ for hydroxyl activation, silanized with APTES, and coated with FN (M + APTES + FN). The comparison of FN adhesion in Fig. 7c showed that microgrooved surfaces, whether silanized or non-silanized, exhibited superior FN adhesion compared to smooth counterparts (S and S + APTES), highlighting the higher loading capacity of microgrooved surfaces due to their increased surface area. Furthermore, silanization increased FN adhesion on both smooth and microgrooved surfaces, demonstrating that covalent loading via silanization enhances the loading capacity for any surface topography. The silanized microgrooved surface showed the most effective adsorption, reflecting the synergistic effect of both physical structure and covalent bonding on loading capacity.

In this study, human dermal fibroblasts (HDFs) cultured on smooth surfaces exhibited spindle-shaped or polygonal morphology, while those on microgrooved surfaces appeared elongated and aligned along the grooves, regardless of whether FN was physically or covalently loaded (Fig. 7d). This indicates that microgrooves guide cell alignment independent of the binding nature of FN. However, covalent bonding via silanization significantly increased vinculin fluorescence intensity compared to non-covalently loaded FN coatings, highlighting the superiority of covalent protein loading over physical loading, particularly on microgrooved surfaces [160]. The highest fluorescence intensity was observed on silanized microgrooved surfaces (M + APTES + FN), reflecting the synergistic effects of microgrooves and covalent loading of FN. A similar trend was observed in the proliferation assay across different groups [160]. These findings are consistent with previous studies demonstrating the synergistic effects of microgrooves and FN in enhancing cell behavior [159], and highlight that covalent FN bonding to the surfaces results in superior cell adhesion and proliferation compared to physical loading [160].

Complementing these findings, another study highlighted the role of FN coatings in promoting keratinocyte adhesion and spreading. FN covalently coupled to Ti surfaces resulted in a twofold improvement in keratinocyte adhesion and spreading compared to unmodified Ti surfaces, supporting PSTI by forming an "epithelial seal" (Fig. 7e) [161]. To stabilize FN and enhance its loading capacity, Ti discs were polished, silanized, and treated with an anthraquinone linker (FN-anthraquinone-coated Ti) for UV-activated covalent attachment. The photocoupling method nearly doubled FN immobilization compared to dip-coating, creating a robust and stable coating [161].

Beyond enhancing cell adhesion [161], FN-coated surfaces have been reported to reduce inflammation by minimizing platelet adhesion. Platelet adhesion on implants triggers the release of cytokines and chemokines, attracting immune cells such as neutrophils and macrophages, which form a fibrin-platelet matrix and amplify the inflammatory response [182]. It has been shown that FN coatings significantly reduce platelet adhesion and clot formation on Ti surfaces, as demonstrated in Fig. 7f [161]. While unmodified Ti showed moderate platelet adhesion, anthraquinone-coated Ti had the highest thrombogenicity. In contrast, FN-anthraquinone-coated Ti reduced Fg adsorption, minimizing platelet adhesion and clot formation.

In addition to mitigating inflammation, FN-functionalized surfaces can enhance PSTI by reducing infection risk through reversible pellicle adsorption [161]—a saliva-derived protein layer that typically facilitates bacterial biofilm development [183]. Unlike unmodified Ti, which forms stable pellicle layers supporting biofilms, FN-coated surfaces disrupt bacterial attachment, lowering infection risk and improving implant longevity [161].

FN biofunctionalization also benefits hard tissue integration in dental implants [163,184]. Enhancing hard tissue responses simultaneously can improve PSTI, as hard tissue fixation stabilizes the implant and prevents movement and micromotions [185]. Chang et al. demonstrated the effectiveness of FN biofunctionalization in improving hard tissue integration [163]. In this work, FN was grafted onto Ti surfaces through a series of surface modifications. First, the implants underwent low-discharge plasma treatment with argon gas (Ar-GDP), followed by exposure to allylamine gas in the GDP reactor and immersion in GA solution for functionalization. Finally, the implants were immersed in a FN solution, and a tris-phosphate buffer (pH 7.4) was used to interrupt FN crosslinking. Although the authors did not explicitly specify whether the bonding was physical or covalent, FN molecules were likely covalently bound to the amine-functionalized surface via GA mediation. The treated implants, Ar-GDP and FN-grafted Ti implants (GDP-fib) were evaluated in vivo using a beagle dog model and were placed in the mandibular premolar area of the dogs for 2–8 weeks to assess their impact on peri-implant tissue response. The findings showed that GDP-fib implants showed significantly higher bone maturation in comparison to Ar-GDP implants. The enhanced outcomes for GDP-fib implants were attributed to FN's role in improving surface properties such as hydrophilicity and roughness, which facilitated better cell adhesion and differentiation. Although the study noted that FN increased the surface roughness on GDP-treated surfaces, the thickness of the FN coating was not measured, leaving the mechanism behind this effect unclear.

Nevertheless, a key challenge in functionalizing Ti surfaces with FN is its sensitivity to degradation, limiting its clinical application [164]. To address this challenge, a recent study investigated functionalizing Ti with a recombinant HBII domain of FN, modified with an Arg-Gly-Asp (RGD) sequence. The findings demonstrated that engineered FN fragments can enhance fibroblast adhesion, proliferation, migration, and activation, achieving effects comparable to full-length FN [164].

An additional advantage of FN functionalization on dental implants lies in its emerging immunomodulatory capacity. While not yet explored within the context of PSTI, FN has been shown to exert potent regulatory effects on immune cell differentiation when incorporated into collagen-based matrices. For example, as demonstrated by Logie et al. [186] FN-functionalized 3D collagen networks were shown to modulate the differentiation of primary human monocytes into a tolerant macrophage phenotype, distinct from classical pro-inflammatory M1 or anti-inflammatory M2 polarization states [186]. This tolerant phenotype is characterized by suppressed expression of inflammatory cytokines under basal conditions, yet retains the ability to mount a protective response upon pathogenic challenge [186]. FN demonstrates the ability to promote immune homeostasis without impairing host defense. Leveraging this property at the transmucosal region of dental implants may reduce chronic inflammation and improve long-term mucosal integration.

4.4. Laminin-biofunctionalized surfaces

Laminin, a high-molecular-weight glycoprotein, is a principal component of basement membranes within the ECM. It provides structural support within the basement membrane and regulates critical cellular processes, including adhesion, migration, and differentiation [187,188]. Its structure offers multiple binding sites for cell receptors, such as integrins, facilitating cell-matrix interactions [[188], [189], [190]]. This complex protein family includes various types, each with unique combinations of α, β, and γ chains, tailored to support specific tissue functions [191]. The cell-binding activity of laminin is not primarily due to an RGD motif [190]. Instead, the mechanism primarily relies on its globular laminin G-like (LG) domains, particularly at the carboxyl-terminal (C-terminal) ends of the α, β, and γ chains, to engage integrins and cell surface receptors [190].

Laminin coatings on implants and scaffolds have shown improved cell adhesion, proliferation, and differentiation for enhanced PSTI [166,192,193]. Key types of laminin investigated in this field include LN332 (laminin-5) [37], laminin-511 [194], laminin-411 [195], and laminin-521 [196]. Among these, LN332, also known as laminin-5, has been more widely used than others for PSTI of dental implants.

LN332, is a component of the basement membrane in epithelial tissues, especially in areas exposed to high mechanical stress [197]. Produced by epithelial cells, LN332 is abundant in the skin, anchors epidermal cells to the underlying dermis, and in the gingival epithelium, aids attachment around teeth [192]. Recognized for its ability to promote epithelial cell adhesion and migration, it is ideally suited for applications in PSTI. In this context, a few studies have investigated the potential of LN332-functionalized Ti surfaces for enhancing PSTI of dental implants [34,37].

In a study by Liu et al., the non-covalent loading of LN332 onto Ti substrates to enhance PSTI of dental implants was investigated [37]. They developed a poly(D,L-lactide) (PDLLA)- LN332 (TiPLN) composite coating through a LBL process, with LN332 binding to PDLLA-coated Ti via non-covalent interactions, such as hydrophobic forces, ionic interactions, and polyelectrolyte absorption [37]. TiPLN surfaces significantly improved human epidermal keratinocyte (HaCaT) adhesion, spreading, migration, and proliferation (Fig. 8a–d). HaCaT cells cultured on TiPLN exhibited more rapid and extensive spreading, with larger cell areas and more developed filopodia than cells on uncoated Ti (Ti in Fig. 8a) and PDLLA-coated Ti (TiP in Fig. 8b) surfaces. Proliferation steadily increased over seven days, highlighting the coating's benefits for HaCaT growth. It was shown that laminin-coated substrates activated the PI3K-Akt pathway, which is fundamental for cellular adhesion, migration, and proliferation. Laminin-coated substrates also upregulated hemidesmosome-related proteins (integrin α6, LN332, integrin β4, and plectin), which are important in stable epithelial attachment and reduced bacterial invasion [37]. In the same study, the laminin coating also enhanced the adhesion, spreading and proliferation of gingival mesenchymal stem cells (GMSCs) and promoted their transdifferentiation into epithelial-like cells [37]. Such a cell-instructive, surface bio-engineering approach can be harnessed to form a tissue seal around the implant, by mimicking the natural epithelial barrier.

Fig. 8.

Laminin biofunctionalized surfaces for PSTIa) SEM images of HaCaT cells on Ti, TiP, and TiPLN substrates at 6, 12, and 24 h show enhanced adhesion, spreading, and filopodia development on TiPLN surfaces. b) Confocal images of HaCaT cytoskeleton at 2, 6, and 24 h show improved cell organization and structure on TiPLN, supporting stable epithelial attachment. c) Scratch-wound assay shows faster HaCaT migration on TiPLN. d) Proliferation curves indicate increased HaCaT growth on TiPLN [37]. e) Green fluorescent protein (GFP) expression efficiency of HEK293 cells on CS-(HA-LDc)n coatings peaks at five layers, indicating optimal DNA delivery and expression. f) DNA loading efficiency increases with more layers but peaks at five before declining due to excess negative charge. g) Fluorescence assay shows enhanced HEK293 proliferation and attachment on CS-(HA-LDc)5, supporting improved ECM integration [198]. LN332 biofunctionalization significantly upregulated tissue repair genes in oral keratinocytes cultured for 4, 24, and 48 h on dental implant fragments biofunctionalized with LN332, assessed using reverse transcription quantitative polymerase chain reaction (RT-qPCR). h) TGF-β gene expression. i) FGF gene expression. j) Epidermal Growth Factor (EGF) gene expression. k) VEGF gene expression. GAPDH was used as the housekeeping gene (control) [34].

Considering the effectiveness of the LBL technique for non-covalently loading laminin onto surfaces to biofunctionalized Ti, optimizing the number of layers could further enhance its performance. While not specifically optimized for laminin, this method has been adapted to load laminin γ2 Deoxyribonucleic acid (DNA), a subunit of laminin-5, onto Ti substrates. This was achieved by alternating layers of positively charged chitosan and negatively charged hyaluronic acid, combined with cationic lipid-DNA complexes [198]. The LBL approach created a stable, hydrophilic multilayer coating through electrostatic interactions, with configurations tested at one, three, five, seven, and nine layers. A key finding was the optimization of five coating layers, which demonstrated the highest transfection efficiency—a process where foreign DNA is introduced into cells—allowing cells to produce laminin-5 (Fig. 8e). Moreover, compared to other configurations, five layers provided an ideal release profile, ensuring timely DNA delivery for effective expression. Although increasing the number of layers initially improved DNA availability (Fig. 8f), additional layers beyond five introduced excess negative charge from HA, which competed with cellular uptake and ultimately reduced transfection efficiency [198]. The study also demonstrated that laminin γ2 DNA coatings with five layers (CS-HA lip 5) enhanced laminin-5 expression by promoting its incorporation into the ECM and interaction with integrin α6β4, strengthening epithelial attachment, hemidesmosome formation, and proliferation (Fig. 8g).

Covalent immobilization of LN332 onto Ti surfaces has been investigated to improve coating stability and bioactivity [34]. In a recent study, laminin coatings on Ti surfaces were created using a silanization method, where the surfaces were treated with APTES to introduce amino groups, followed by GA crosslinking to covalently attach LN332. LN332 biofunctionalized surfaces demonstrated excellent biocompatibility, supporting high oral keratinocyte viability, enhanced adhesion, spreading, proliferation, and upregulation of key tissue regeneration genes such as TGF-β, VEGF, FGF, and EGF (Fig. 8h–k) [34]. While it has been suggested that a multi-step chemical modification process, including silanization and GA cross-linking, was employed to covalently bind LN332 to the Ti surface, the covalent nature of this attachment remains unverified. Analytical techniques such as energy-dispersive X-ray spectroscopy (EDS) and scanning Electron Microscopy (SEM) confirmed the presence of LN332 on the surface but did not distinguish between covalent bonding and simple adsorption. To validate covalent attachment, a washing step with sodium dodecyl sulfate (SDS), a detergent capable of removing non-covalently bound proteins while retaining covalently immobilized ones, could have been employed [199,200]. Furthermore, assessing the long-term stability of the coating would provide important insights into its durability and potential clinical performance. Evaluating its resistance to hydrolytic degradation under physiological conditions, such as incubation in buffer solutions at varying pH levels over time, could help determine robustness for long-term applications.

The surface conformation of LN332 is critical in mediating soft tissue cell responses to implants. This was demonstrated in a study where HaCaT keratinocytes exhibited greater attachment and spreading on LN332-coated passivated Ti surfaces compared to unpassivated counterparts [166]. Ti-6Al-4V disks were passivated via acid immersion prior to laminin coating, a process that, while not significantly altering protein adsorption levels, influenced protein conformation, as informed by a combination of AFM, electron spectroscopy for chemical analysis (ESCA), and time-of-flight secondary ion mass spectrometry (ToF-SIMS) analyses. The conformation changes in proteins were attributed to the formation of a thicker and more homogeneous Ti oxide (TiO₂) layer, as calculated using Auger Electron Spectroscopy (AES). The oxide layer thickness increased from 33.3 Å in unpassivated samples to 93.9 Å in passivated samples.

The authors reported that keratinocytes spread extensively on LN332-coated passivated surfaces, forming significantly more hemidesmosomes than on unpassivated surfaces. They attributed this cellular behavior to the altered conformation of LN332 on passivated surfaces, which they suggested enhances the exposure of cell-binding domains, thereby facilitating stronger cell attachment. The study also proposed that a possible contributing factor, though not directly tested, could be the reduced ion release due to the formation of a thicker passive oxide layer, which may provide a more stable environment for protein adsorption and cell interactions. Additionally, AFM images qualitatively indicated surface topography differences between LN332-coated passivated and unpassivated samples, which the authors suggested as another potential factor in influencing cell behavior [166].

Enhancing PSTI requires more than the mere presence of bioactive proteins such as LN332; it demands precise control over the physicochemical properties of the substrate to modulate protein conformation and biofunctionality. Evidence suggests that passivated surfaces can markedly improve the biological presentation of LN332, despite similar total protein adsorption, highlighting the critical role of subtle surface cues—such as oxide layer composition and chemical uniformity—in directing cellular responses [166]. In parallel, the long-term retention and functional stability of protein coatings under physiological conditions necessitate robust immobilization strategies. Covalent binding of LN332 to the implant surface provides a durable interface that preserves protein activity over time [34]. Taken together, these findings advocate for a synergistic design approach that integrates surface passivation with covalent functionalization to engineer stable, biologically responsive implant surfaces capable of supporting sustained soft tissue attachment and homeostasis.

4.5. Fibrinogen (Fg)/fibrin-biofunctionalized surfaces

Fg and fibrin are components of the blood clotting process, each with distinct but interconnected roles [201]. Fg is a soluble glycoprotein produced in the liver and present in blood plasma, serving as the precursor to fibrin [202]. Fibrin, conversely, is a protein formed from Fg during the clotting process, creating the structural framework of a blood clot [203]. When a blood vessel is damaged, thrombin, an enzyme in the clotting cascade, cleaves Fg, releasing fibrinopeptides and transforming it into fibrin. Fibrin is an insoluble protein that polymerizes to form a stable mesh, acting as a structural backbone for the blood clot [204]. The fibrin mesh strengthens the clot, prevents further blood loss, and supports tissue repair [205]. Fg can provide biochemical cues for cellular behavior such as adhesion, migration, and proliferation, especially during healing and integration processes [206]. Fg can significantly enhance soft and hard tissue cell responses when applied as a coating on implant surfaces [167].

A study has shown that Ti surfaces pre-coated with PDA, followed by covalent attachment of Fg, formed a stable bioactive layer [167] that facilitated integrin-mediated interactions, leading to well-spread fibroblasts and enhanced keratinocyte adhesion. Furthermore, the coating supported the early attachment of cells resident in hard tissue, such as preosteoblasts (Ti—PDA-Fg in Fig. 9a). Early attachment of preosteoblasts is important for creating a stable implant interface, which in turn promotes soft tissue healing by reducing implant movement or displacement. This study also demonstrated that Fg-coated Ti screws reduce epithelial downgrowth in a rat tibial percutaneous implantation model (Fig. 9b). Epithelial downgrowth, characterized by the migration of outer skin or mucosal cells along the implant surface, can hinder soft tissue attachment and increase the likelihood of inflammation or infection. Managing this challenge affects long-term implant success and integration, and Fg-coated implants have shown significant efficacy in overcoming this issue. Fg coatings in this study also addressed issues such as fibrous capsule formation and limited angiogenesis by reducing fibrous thickness and enhancing vascularization (Fig. 9c and d). A thick fibrous capsule can isolate the implant from surrounding tissues, impeding direct attachment and integration of soft tissue. Also, insufficient angiogenesis—or reduced blood vessel formation—further hampers PSTI by limiting nutrient and oxygen supply [167].

Fig. 9.

Fg biofunctionalized surfaces for PSTI.a) Fluorescence micrographs showing keratinocytes (HaCaT), human foreskin fibroblasts (HFF), and preosteoblasts (MC3T3-E1) cultured for 12 h on Ti and Ti—PDA-Fg surfaces demonstrated improved cell adhesion and spreading on the Ti—PDA-Fg surface compared to unmodified Ti. b) HE-stained images of Ti and Ti–PDA–Fg surfaces at 5 months revealed reduced epithelial downgrowth in the Ti–PDA–Fg group. c) Statistical analysis of normalized fibrous capsule thickness showed a significant reduction in fibrous capsule formation around Ti–PDA–Fg implants compared to Ti. d) Quantification of new vessels per mm²demonstrated enhanced angiogenesis around Ti–PDA–Fg implants. Reproduced with permission from Wang X. et al. [167], Biomaterials Science, 2021, 9(15), 5192–5208, © Royal Society of Chemistry. e) Epifluorescence images of HDFs on nonmodified and fibrin-coated PLA (PLA F) at days 1 and 3 show improved adhesion and spreading on PLA F. f) Confocal images of β1-integrin in fibroblasts on day 3 show enhanced cell-matrix adhesion on PLA F. g) Mitochondrial activity of HDFs on nonmodified and fibrin-coated PLA F shows higher metabolism, especially with AA, indicating enhanced fibroblast proliferation. Reproduced with permission fromRef. [210], Int J Nanomedicine, 2016;11:771–789. © Dove Medical Press. CC BY-NC 3.0.

Similarly to Fg, fibrin can also be employed to functionalize the surface of dental implants, enhancing PSTI and sealing. The integrin-binding sites in fibrin facilitate cell attachment and signaling, thereby promoting healing and tissue remodeling [207]. In addition, the three-dimensional fibrous network offers distinct advantages by enhancing cell adhesion, migration, and proliferation, contributing to the development of a robust soft tissue seal and enhanced angiogenesis [208,209]. Although fibrin coatings have not yet been applied to dental implants for improving PSTI, their potential to enhance soft tissue cell response was demonstrated in a study where electrospun polylactide (PLA) nanofibrous membranes were coated with a fibrin nanolayer [210]. The fibrin coating was created through the thrombin activation of adsorbed Fg, resulting in a uniform nanocoating that remained stable in culture conditions. The fibrin layer significantly improved HDF spreading and adhesion by promoting β1-integrin-mediated interactions and stimulating the production of collagen I (Fig. 9e and f). Moreover, fibroblast proliferation was significantly higher on fibrin-coated membranes compared to nonmodified PLA, particularly in the presence of acorbic acid (AA)-supplemented media, which further enhanced metabolic activity and cell growth (Fig. 9g).

The potential of fibrin-based functionalization of dental implants to improve PSTI could be further demonstrated by its broad applications in contact with other soft tissue in the form of fibrin glue [211,212]. Various fibrin-based glues—a combination of Fg and thrombin—have been developed for diverse soft tissue applications, with their properties tailored by adjusting Fg and thrombin concentrations. Two Food and Drug Administration (FDA)-approved fibrin glues, Tisseel (Baxter, Westlake Village, CA) [213] and Evicel (Ethicon/J&J, Somerville, NJ) [214], highlight this versatility. Tisseel, containing 85 mg/mL Fg and 500 IU/mL thrombin, is suited for procedures requiring moderate clotting speed and enhanced stability, such as gastrointestinal sealing, cardiovascular surgeries, and skin graft fixation. In contrast, Evicel, with 70 mg/mL Fg and 1000 IU/mL thrombin, promotes rapid clot formation, making it ideal for soft tissue surgeries, including laparoscopic and vascular procedures, where quick hemostasis is needed. Various concentrations of Fg and thrombin can be explored to optimize fibrin glue for biofunctionalizing dental implant surfaces to enhance PSTI.

While fibrin glues have not yet been specifically investigated or utilized for PSTI in dental implants, promising outcomes have been reported in bone integration through the Fg-induced regeneration sealing technique (F.I.R.S.T.) [215]. This advanced approach incorporated fibrin-based sealants to enhance the effectiveness of guided bone regeneration (GBR). GBR is a surgical technique that utilizes bone grafts and barrier membranes to address minor defects around dental implants [216]. In a study involving 62 patients with 105 implants monitored over 3–7 years, Fibrinogen-induced regeneration sealing technique (F.I.R.S.T.) achieved a cumulative implant survival rate of over 94 %. Significant horizontal and vertical bone gains were also observed [215]. Although the study primarily focused on bone regeneration and did not quantitatively analyze the role of fibrin sealants in PSTI, consistent reports of healthy soft tissue outcomes in cases involving these sealants suggest their potential effectiveness. Performing a quantitative analysis of PSTI could offer a more comprehensive understanding of their role and efficacy in this context.

4.6. Anti-inflammatory protein–biofunctionalized surfaces

Beyond their established roles in promoting cellular adhesion and tissue integration, certain proteins contribute to PSTI through intrinsic anti-inflammatory activity. These anti-inflammatory proteins actively suppress or resolve inflammatory responses, thereby fostering a more regenerative microenvironment at the tissue–implant interface [217]. Examples of anti-inflammatory proteins include cytokines [218] and enzyme inhibitors [219], each fulfilling distinct roles in modulating immune responses. However, in the context of biofunctionalized biomaterial surfaces, particularly for PSTI, cytokines have garnered particular attention [217,220]. This section discusses the role of immunomodulation in promoting PSTI and highlights the potential of anti-inflammatory proteins as biofunctional components in surface modification strategies aimed at enhancing PSTI.

4.6.1. Immunomodulation for PSTI: impact of anti-inflammatory biofunctionalization

Following dental implant placement, inflammation serves as an essential component of the early wound healing process, coordinating the immune response to surgical injury and biomaterial insertion, as described in Section 4.6.1 [221]. Tissue disruption elicits the rapid release of pro-inflammatory cytokines and chemokines, which recruit neutrophils and monocytes to the implant site [222]. Neutrophils facilitate early microbial defense and debris clearance, while monocyte-derived macrophages initiate broader immune regulation and tissue remodeling [223]. Although inflammation is a natural component of the healing process, excessive or prolonged inflammatory responses can disrupt tissue repair and integration [224].

An important factor influencing this outcome is the polarization state of macrophages. In response to local microenvironmental cues, macrophages exhibit remarkable functional plasticity, adopting distinct phenotypes that include the undifferentiated M0 state, the pro-inflammatory M1 phenotype, and the anti-inflammatory, pro-regenerative M2 phenotype [225,226]. M0 macrophages represent the undifferentiated or resting state of macrophages, serving as the baseline phenotype from which they can be polarized into either the pro-inflammatory M1 or anti-inflammatory M2 states in response to environmental cues [225].

Pro-inflammatory M1 macrophages impair tissue repair by releasing cytokines such as tumor necrosis factor-α (TNF-α) and interleukin-1 beta (IL-1β), which inhibit fibroblast proliferation, suppress collagen synthesis, and disrupt epithelial adhesion. In contrast, anti-inflammatory M2 macrophages create a regenerative environment by secreting interleukin-10 (IL-10), TGF-β, and interleukin-4 (IL-4) [217]. These mediators enhance ECM deposition and remodeling by fibroblasts and promote the expression of integrins and basement membrane components, thereby facilitating hemidesmosome formation and reinforcing epithelial attachment at the implant–tissue interface [227]. Hence, by promoting a shift toward the M2 phenotype, anti-inflammatory agents not only reduce local inflammation but also enhance the regenerative activities of fibroblasts and epithelial cells [217,228,229].

Therefore, promoting stable PSTI may be achieved by modulating the local immune environment to direct macrophage polarization toward the pro-regenerative M2 phenotype [217]. A range of immunomodulatory strategies have been explored to induce macrophage polarization from the pro-inflammatory M1 state to the pro-regenerative M2 phenotype. These approaches include the rational design of surface topographies (e.g., grooves or TiO₂ nanotubes), modulation of surface wettability, incorporation of immunoregulatory ions such as magnesium(II) ion (Mg²⁺) or strontium(II) ion (Sr²⁺), and the immobilization of bioactive agents [217]. Among these strategies, immobilizing bioactive agents, such as anti-inflammatory cytokines [230] and anti-inflammatory peptides [231,232], offers a distinct advantage by enabling the localized, sustained delivery of immunomodulatory signals that directly engage immune and stromal cells at the tissue–implant interface, thereby closely replicating the natural wound healing environment.

Biomolecules commonly employed for surface biofunctionalization primarily include cytokines (typically protein-based) and anti-inflammatory peptides. In this section, we review biofunctionalized surfaces incorporating anti-inflammatory cytokines aimed at PSTI. Approaches involving anti-inflammatory peptides will be addressed in a subsequent section.

4.6.2. Cytokines-biofunctionalized surfaces

Cytokines are small, soluble proteins that serve as signaling molecules within the immune system, orchestrating a wide range of biological processes including intercellular communication, inflammation, and tissue repair [233]. Secreted by both immune and non-immune cells, cytokines exert their effects through interaction with specific surface receptors, activating tightly regulated intracellular signaling cascades [233,234]. In the context of immune modulation, cytokines exhibit dual functionality: while pro-inflammatory cytokines such as TNF-α, interleukin-1β (IL-1β), interleukin-6 (IL-6), and interleukin-17 (IL-17) initiate and amplify immune responses essential for pathogen clearance and early wound healing, their prolonged expression can lead to chronic inflammation and tissue degradation [235]. Conversely, anti-inflammatory cytokines—including IL-4 and IL-10—promote immune resolution, suppress excessive inflammation, and support tissue regeneration via promoting the polarization of macrophages towards the M2 phenotype [236].

In the context of PSTI, cytokines such as IL-4, IL-10, and interleukin-13 (IL-13) have demonstrated potent immunomodulatory effects, effectively reprogramming macrophages toward a pro-regenerative phenotype and fostering a microenvironment conducive for PSTI [217,220]. For instance, Liu et al. [229]. engineered "immune-smart" Ti implants by first coating the surface with PDA and subsequently conjugating IL-4. In a rat transgingival implantation model, IL-4/PDA-coated Ti implants were inserted through the gingival mucosa following maxillary molar extraction (Fig. 10a). Histological analysis performed two weeks post-implantation revealed that while all groups developed an epithelial layer parallel to the implant surface, only the IL-4/PDA-coated implants demonstrated a continuous, well-organized epithelial structure (Fig. 10b) [229]. Immunohistochemical staining further confirmed significantly elevated cluster of differentiation 206-positive (CD206⁺) macrophage infiltration—indicative of M2 polarization—and enhanced linear expression of laminin 5α3, a key protein involved in hemidesmosome formation and epithelial attachment, at the soft tissue–implant interface (Fig. 10c and d). In contrast, control and PDA-only groups exhibited disorganized laminin distribution and fewer M2 macrophages, correlating with impaired epithelial sealing and heightened inflammation [229]. These findings underscore the efficacy of IL-4-functionalized surfaces in reprogramming the local immune microenvironment to promote stable epithelial integration.

Fig. 10.

Immunomodulatory effects of cytokine (IL-4) for promoting soft and hard tissue integration at the dental implant interface. a) Schematic of the transgingival implantation model used by Liu et al., in which Ti mini-implants were inserted through the gingival mucosa following maxillary molar extraction in rats. b) Histological staining revealed that IL-4/PDA-coated implants supported the formation of a continuous and well-organized epithelial layer, in contrast to control and PDA-only groups. c–d) Immunohistochemical analysis showed increased infiltration of CD206⁺ M2 macrophages and enhanced linear expression of laminin 5α3 at the IL-4-functionalized interface, indicating improved epithelial sealing and reduced inflammation. Reproduced from Ref. [229] with permission from Wiley-VCH. © 2020 Wiley-VCH GmbH. e) RT-PCR analysis from rat subcutaneous implantation model showing that IL-4 downregulated M1-associated genes (CCR7, IL-1β) and upregulated M2 and osteogenic markers (CD206, BMP2) in peri-implant tissue. f) Immunofluorescence staining confirmed macrophage phenotype switching, with fewer iNOS⁺ (M1) cells and more CD206⁺ (M2) cells in the IL-4-treated groups, while H&E staining demonstrated reduced fibrous capsule thickness, reflecting attenuated chronic inflammation. Reproduced from Ref. [237] with permission from Elsevier. © 2020 Elsevier B.V.

Beyond its established role in PSTI in dental implants, IL-4 has also been investigated for enhancing hard tissue integration under inflammatory conditions. For example, Zhao et al. [237] developed calcium–strontium–zinc–phosphate (CSZP) coatings to improve osseointegration by increasing surface wettability and releasing osteogenic ions. However, these coatings also elicited a pro-inflammatory response, marked by M1 macrophage polarization. In both in vitro and in vivo models, IL-4 supplementation successfully counteracted this effect, promoting a shift toward the anti-inflammatory M2 phenotype. In a rat subcutaneous implantation model, CSZP-coated implants pre-seeded with bone marrow–derived mesenchymal stem cells (BM-MSCs) were implanted, and IL-4 was locally administered from day 3–7 post-implantation. Reverse transcription polymerase chain reaction (RT-PCR) analysis (Fig. 10e) revealed that IL-4 reduced expression of M1-associated genes (CCR7, IL-1β) while upregulating M2 and osteogenic markers (CD206, BMP2). Immunofluorescence staining (Fig. 10f) confirmed this phenotypic shift, showing decreased inducible nitric oxide synthase-positive (iNOS⁺) (M1) and increased CD206⁺ (M2) macrophages. H&E staining further demonstrated a reduction in fibrous capsule thickness following IL-4 treatment, indicating diminished chronic inflammation. Hence, these findings highlight IL-4's capacity to modulate the immune response and foster a regenerative peri-implant environment conducive to hard tissue integration [237].

Despite these promising results, the clinical application of IL-4 remains limited due to its instability and rapid degradation in physiological environments. To overcome these challenges, a recent study developed a stabilized delivery platform incorporating IL-4 and bovine serum albumin (BSA) within Poly(lactic-co-glycolic acid) films [238]. Coupled with the topographical guidance of electrospun poly-L-lactic acid (PLLA) fibers, the platform enabled sustained cytokine release while preserving IL-4 bioactivity. The BSA-stabilized IL-4 significantly enhanced macrophage elongation, upregulated arginase-1 and IL-10/interleukin-12 subunit p40 (IL-12p40) gene expression, and suppressed the secretion of IL-12p40 and regulated upon activation, normal t cell expressed and secreted (RANTES). The findings highlight the critical role of cytokine stabilization within biomaterial platforms in enabling effective immunomodulation, offering a compelling strategy to resolve inflammation and facilitate tissue regeneration.

Collectively, the reviewed studies illustrate the potential of cytokine-functionalized surfaces—particularly those leveraging IL-4—to modulate the immune landscape at the tissue–implant interface, guiding macrophage behavior, reducing inflammation, and enhancing both soft and hard tissue integration. Such strategies represent a compelling direction for next-generation implant designs in dental, craniofacial, and orthopedic applications.

5. Peptide-biofunctionalized surfaces

Peptides are short chains of 2–50 amino acids linked by peptide bonds. Naturally occurring in the body, peptides are essential for various biological functions, including cell signaling, immune responses, and tissue repair. In PSTI, peptides play key roles in cell adhesion, proliferation, and migration, promoting tissue regeneration [239]. They can also modulate immune responses, reduce inflammation, and prevent bacterial invasion around the implant, improving tissue integration [240,241]. Peptides that have been or could potentially be utilized to enhance the PSTI of dental implants are typically categorized into five groups; cell adhesion peptides (CAPs), mimetic peptides, growth factor-binding peptides, anti-inflammatory peptides, and AMPs. Each of these peptide groups has distinct structures and functions, which are detailed in this section. Table 2 summarizes their immobilization strategies, substrate types, and key functional roles.

Table 2.

A summary of literature in which peptides were used for PSTI in dental implants, including underlying substrates, methods of immobilization and main findings.

Peptide	Dental implant material	Method of immobilization	Main findings	Reference
RGD CAPs with either linear or cyclic sequences	ZrO2 abutment	Direct covalent immobilization via a PDA film pre-coated on the surface	Enhanced HGF attachment, spreading, proliferation, and activity, while reducing bacterial colonization on ZrO2 without compromising PDA's antimicrobial properties	[26]
RGD CAPs	Commercially pure (CP) Ti substrates	Direct covalent immobilization by Carbonyldiimidazole (CDI) activation	Significant increase in the in-vitro cell adhesion and proliferation of HGFs and epithelial cells	[242]
Laminin-5-derived peptide	Ti substrates	Direct covalent immobilization using a Poly-l-lysin (PLL)/Poly(l-glutamic) acid multilayered polyelectrolyte film pre-coated on the surface	Enhanced human oral epithelial cell adhesion and proliferation, with in vivo results showing seamless tissue integration of biofunctionalized Ti implants	[243]
Interleukin – 23 (IL-23) receptor antagonist peptide	Ti substrates	Direct covalent immobilization via surface pre-silanization	Reduced the expression of pro-inflammatory IL-23 cytokine and polarized macrophages toward a pro-regenerative M2-like phenotype that can facilitate soft tissue (e.g., gingival) regeneration	[232]
Leucine-leucine 37 (LL-37) antimicrobial peptide (AMP)	Polystyrene tissue culture plates	Non-covalent immobilization via physisorption	LL-37 AMP bacteria-induced inflammation and degradation in HGFs, enhancing their viability, proliferation, nutrient uptake, and ECM production	[27]

5.1. Cell adhesion peptides (CAPs)

CAPs mimic the binding sites on ECM components such as laminin, FN, and collagen. They attach to cell receptors (e.g., integrins) and improve their adhesion to the surface [244]. There are various types of CAPs with distinct peptide sequences and biological functions. The primary CAPs for soft tissue applications include RGD, PHSRN, laminin-derived and ameloblastin-derived peptides. RGD peptides effectively mimic the binding sites of FN and vitronectin, facilitating cell attachment and signaling [245]. Integrin receptors recognize and bind to the RGD sequences to trigger cellular signaling pathways, enhancing cell adhesion and proliferation [5]. This mechanism resembles the cell adhesion process to the ECM-coated surfaces discussed in Section 3.

RGD peptides with either linear (KGGRGDSP) or cyclic (c(RGDfK)) sequences have been covalently immobilized onto ZrO₂ abutment surfaces pre-coated with PDA films [26]. Linear RGD peptides, with an open peptide chain containing the RGD motif and additional amino acids (KGGRGDSP), exhibit a flexible structure that is less resistant to enzymatic degradation compared to the cyclic forms. In contrast, the cyclic RGD peptide is stabilized by disulfide bonds linking the peptide chain termini to form a loop. RGD-biofunctionalized ZrO₂ surfaces, particularly those modified with cyclic peptides, demonstrated significantly enhanced cell adhesion, proliferation, and spreading of HGFs, as shown in Fig. 11a–e [26]. The reported superior effectiveness of cyclic RGD peptides can be attributed to their 20-100-fold stronger affinity for integrins [246], and their greater stability and resistance to proteolysis [247] compared to the linear RGD peptides. The RGD-functionalized ZrO₂ substrates also maintained the antimicrobial properties of PDA, effectively reducing colonization by Streptococcus mutans and Porphyromonas gingivalis, indicating their potential to improve implant integration and minimizing infection [26]. This study [26] provides important findings on the biofunctionalization of implants with RGD CAPs for improved PSTI. However, two key limitations affect clinical applicability. Firstly, while the study evaluated the initial effectiveness of RGD-functionalization in reducing bacterial colonization, it did not address the long-term antibacterial efficacy over time. Moreover, the stability of the RGD peptides on the biofunctionalized surface under dynamic physiological conditions remains unexplored, which is crucial for maintaining functionality in vivo.

Fig. 11.

Cell-adhesion peptide biofunctionalized surfaces for PSTI. Confocal laser scanning microscopy images of HGFs after a) 3 h and b) 24 h of cell culture (scale bar = 20 μm for high magnification and scale bar = 100 μm for low magnification), and quantitative results for spreading areas, c) perimeters, d) and numbers e) of HGFs cultured for 3 and 24 h on pristine ZrO₂, PDA-coated (ZrO₂-P), coated ZrO₂with PDA and liner cell adhesion peptides (ZrO₂-P/L), and coated ZrO2 with PDA and cycle peptides (ZrO₂-P/C), showing that biofunctionalization with RGD peptides, especially in their cyclic forms, can enhance the adhesion, spreading, and proliferation of HGFs [26], f) Immunofluorescent (scale bar = 100 μm) and SEM images (scale bar = 50 μm) of human oral epithelial cell (HOEC) on uncoated Ti with smooth and porous surface, coated porous Ti with PLL/PGA, and coated porous Ti with PLL/PGA as well as laminin 5-derived peptide, showing that laminin 5-derived peptide increases the tendency of HOECs to spread, proliferate, and colonize the porous Ti substrate. g) Images of the stained transversal sections of smooth and porous Ti implants before and after functionalization with laminin 5-derived peptide after 6 months of implantation in dogs, exhibiting an enhanced PSTI to the porous Ti implant when biofunctionalized with the peptide. Reproduced with permission fromRef. [243]., Biomaterials, 2009;30(12):2291–2301. © 2009 Elsevier Ltd. All rights reserved.

The effectiveness of RGD peptides in promoting PSTI was further demonstrated in another study that utilized carbonyldiimidazole (CDI) activation to attach these peptides to the surface of commercially pure (CP) Ti substrates [242]. In the CDI activation, 1,1′-carbonyl diimidazole activates carboxyl or hydroxyl groups on the substrate surface, generating reactive sites that facilitate peptide binding [242]. The results indicated a marked increase in HGF and epithelial cell adhesion and proliferation on the RGD-functionalized surface over a 72-h culture period [242]. Nevertheless, a more comprehensive optimization of peptide concentration on the surface could be beneficial for optimizing cell adhesion without overstimulating inflammatory responses [248]. The degradation profile of the attached peptides could also be assessed to predict long-term performance of the biofunctionalized surface.

The combination of different CAPs can generate synergistic effects that significantly enhance PSTI. For example, co-administering RGD and PHSRN peptides has been shown to enhance integrin-mediated cell adhesion in human mammary fibroblasts [249]. PHSRN adopts a specific conformation that aligns with integrin receptors, facilitating cell adhesion and influencing cellular behavior. Both RGD and Pro-His-Ser-Arg-Asn (PHSRN) peptides replicate natural binding domains in FN, a key ECM protein [249]. When applied together, these peptides increase the binding affinity of RGD for integrin receptors, while promoting integrin clustering and downstream signaling [249]. This interaction leads to stronger cell adhesion, ultimately supporting improved tissue integration.

Laminin-derived peptides, another type of CAPs, mimic specific laminin regions that interact with cell surface receptors, including integrins. Key examples of the laminin-derived peptides include Tyr-Ile-Gly-Ser-Arg (YIGSR) and Ile-Lys-Val-Ala-Val (IKVAV), which have demonstrated efficacy in promoting epithelial cell attachment and spreading [250], which is critical for PSTI in dental implants. Laminin-derived peptides facilitate cell-matrix adhesion by engaging receptors such as integrins and syndecans, thus activating intracellular signaling pathways [251]. Werner, S. et al. covalently grafted laminin-5-derived peptide onto porous Ti substrates pre-coated with native or cross-linked PLL/PGA multilayered polyelectrolyte films (MPFs) [243]. As observed from Fig. 11f, functionalizing the substrates with this peptide significantly enhanced the spreading and proliferation of human oral epithelial cells by mimicking the natural interactions of laminin with cell surface receptors. Furthermore, preliminary in vivo studies conducted in dogs revealed that epithelial cells adhered and proliferated effectively on the porous Ti implants biofunctionalized with the laminin-5-derived peptide. These implants also demonstrated seamless integration into surrounding tissues, with soft tissue cells colonizing the pores, as shown in Fig. 11g.

Ameloblastin-derived peptides are also classified as CAPS. They are derived from a tooth-specific protein in the hard, outermost layer of teeth (known as enamel), which participates in tissue regeneration when the periodontium is disturbed [252,253]. While ameloblastin-derived peptides have not yet been used alone for PSTI in dental implants; they were used in combination with a LN332-derived peptide to enhance the peri-implant mucosal seal in a study by Koidou et al. [254]. In this work, ameloblastin- and LN332-derived peptides were covalently immobilized on Ti substrates via pre-silanization [254]. The peptides were expected to have complementary functions in promoting PSTI. Ameloblastin-derived peptides can support tissue regeneration and adhesion [253], while LN332 can enhance epithelial attachment by facilitating hemidesmosome formation [255]. The in-vitro results fulfilled these expectations, showing higher effectiveness of the combined peptide coating in improving keratinocyte proliferation compared to control surfaces and monopeptide coatings, particularly within the first 24 h of culture. Keratinocytes are the primary cells of the epidermis that can form a protective barrier and attach to the implant surface via hemidesmosomes [256]. This attachment stabilizes epithelial integration and prevents bacterial invasion [256]. In addition, the combination of ameloblastin- and LN332-derived peptides significantly increased the in vitro formation of hemidesmosomes, which are critical for stable epithelial attachment [254]. Based on these results, co-immobilization of laminin- and ameloblastin-derived peptides has the potential to improve the per-mucosal seal around dental implants, ultimately preventing peri-implantitis. However, multiple cell studies involving fibroblasts, immune cells, and oral microbiota should be complemented to evaluate holistic tissue responses. Comprehensive animal and clinical experiments, which fully replicate the complex in vivo environment around the dental implants, are also required to validate the findings in a physiological context.

5.2. Mimetic peptides

Mimetic peptides are synthetic peptides engineered to replicate the functional properties of natural proteins, allowing them to interact with cell receptors and activate specific biological pathways, including cell proliferation, migration, differentiation, and ECM production [257]. By mimicking GFs and other signaling proteins, mimetic peptides can enhance tissue repair and regeneration, making them valuable in regenerative medicine, wound healing, and implant integration. Mimetic peptides are especially useful in tissue engineering, providing targeted and controlled stimulation without the risks associated with high doses of natural GFs. Peptides have been developed to mimic transforming growth factor-beta1 (TGF-β1) and BMP-2, which carry key functions in PSTI.

TGF-β1-mimetic peptides replicate the structure and functions of the naturally occurring TGF-β1, stimulating fibroblast proliferation and collagen synthesis for tissue repair [258]. To date, TGF-β1-mimicking peptides have seen limited application in the biofunctionalization of dental implants for PSTI. However, their effectiveness in cartilage repair [259] suggests promising potential for improving PSTI. These peptides stimulate fibroblast or epithelial cell proliferation and ECM synthesis. Additionally, their anti-inflammatory properties can mitigate chronic inflammation and fibrotic responses, promoting improved healing and integration with soft tissue [258]. For example, TGF-β1-mimicking peptides were shown to promote in situ cartilage regeneration when incorporated into a hydrogel composite scaffold made from the self-assembling Ac-(RADA)4-CONH2 (RAD) peptide [260]. The scaffolds stimulated chondrocyte proliferation and ECM production, both essential for cartilage formation and maintenance. TGF-β1-mimetic peptides also enhanced the migration of mesenchymal stem cells (MSCs) to the injury site and their differentiation to chondrocytes, further enhancing cartilage repair [260]. Suppressing inflammation around the damaged area was also accomplished by TGF-β1-mimicking peptides, preventing the tissue from further degradation and creating a favorable condition for regeneration [258].

BMP-2 mimetic peptides, which replicate the active binding domains of BMP-2, represent another subclass of mimetic peptides with promising potential for tissue integration [261]. To the best of our knowledge, the application of BMP-2 mimetic peptides in PSTI, particularly for dental implants, has not been explored. However, given their proven efficacy in promoting bone/cartilage regeneration and hard tissue integration [[261], [262], [263]], BMP-2 mimetic peptides represent a promising avenue for future research in soft tissue healing. Studies have shown that these peptides enhance the proliferation, differentiation, and migration of MSCs to the injury site while promoting collagen synthesis [261,262]. BMP-2 mimetic peptides bind to BMP-2 receptors, activating signaling pathways that regulate expression of genes related to growth and development [262]. These receptor interactions also regulate MSC proliferation, differentiation to osteoblasts, and migration toward the injury site, increasing the population of regenerative cells in the damaged area [262]. Furthermore, the induction of collagen synthesis by BMP-2 mimetic peptides enhances the structural integrity of the healing tissue and stimulates the attachment and growth of new cells [262].

5.3. Growth factor-binding peptides

Growth factor-binding peptides are short amino acid sequences, typically 5 to 20 residues long, designed to bind to specific GFs and support biological activities such as tissue repair and regeneration [264]. By enhancing the local concentration, stability, and half-life of GFs, these peptides promote targeted cellular responses and amplify growth factor activity [264]. They can be tailored to bind selectively to GFs like VEGF or FGFs, influencing processes such as tissue repair, angiogenesis, and immune response regulation. In the context of PSTI, the most prominent examples include peptides targeting TGF-β1 and VEGF.

TGF-β1-binding peptides, for instance, specifically bind to TGF-β1—a growth factor crucial for soft tissue regeneration—facilitating its interaction with cell receptors and triggering cell signaling pathways involved in proliferation, migration, and differentiation of cells such as fibroblasts. One study immobilized TGF-β1-binding peptides on polycaprolactone (PCL) films through pre-silanization and NHS/EDC chemical treatment [265]. The peptide-coated films showed homogeneous TGF-β1 accumulation on the surface, as shown in Fig. 12a–c, maintaining its biological activity and leading to a 2.5-fold increase in collagen matrix production in human hamstring cells. In vivo tests in rats further demonstrated TGF-β1 accumulation and enhanced vascularization around the peptide-functionalized implant by day 7, exhibited in Fig. 12d [265]. Although the effect of TGF-β1-binding peptides on the PSTI of dental implants has not been directly studied, their success in enhancing muscle cell behavior suggests similar potential for PSTI of dental implants. However, while TGF-β1 promotes tissue regeneration, excessive activation can lead to fibrosis (the formation of excess fibrous connective tissue) by stimulating fibroblasts to differentiate into myofibroblasts [266]. Therefore, studies should assess the risk of TGF-β1 overstimulation and the potential fibrotic response on the biofunctionalized surfaces with TGF-β1-binding peptides.

Fig. 12.

Surface-attached GF-biding peptides for PSTI; a) Fluorescence quantification of the immunochemistry assay against human Transforming growth factor-β1 (hTGF-β1) on the PCL films either in native forms or biofunctionalized with TGF-β1-binding peptides and incubated either with or without hTGF-β1 (immunochemistry performed with primary and secondary antibodies), exhibiting a significant increase in the GF (hTGF-β1) binding on the peptide-biofunctionalized film incubated with hTGF-β1, b) Fluorescence image of the peptide-functionalized PCL film without hTGF-β1 incubation (left) vs with hTGF-β1 incubation (right) (bar: 1000 μm), c) Fluorescence image of the native film with hTGF-β1 incubation (left) vs the peptide-functionalized film with hTGF-β1 incubation (right) (bar: 1000 μm), all demonstrating a homogenous distribution of the attached GF (hTGF-β1) within the films, d) Histological evaluation of native and the peptide-functionalized PCL film sections after 3 and 7 days of implantation, showing an enhanced fibrogenic response and vascularization around the biofunctionalized implant with TGF-β1-binding peptides. Reproduced with permission fromRef. [265], Acta Biomater., 2017;54:239–249. © 2017 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

Another study incorporated TGF-β1-binding peptides into gelatin methacryloyl (GelMA) hydrogel scaffolds, demonstrating enhanced efficacy in in-vitro cartilage and bone regeneration by promoting the adhesion, proliferation, and osteogenic differentiation of MSCs [267]. Improved osteochondral regeneration was also observed in an in-vivo rat model, where the peptide-incorporated scaffolds were implanted into the defected femoral trochlea of Sprague-Dawley (SD) rats [267]. The finding of this work can be applied to soft tissue-encountering implants where the TGF-β1-binding peptides can enhance the cellular interactions with fibroblasts, ECs, or other relevant soft tissue cells at the implant interface. The hydrogel used in this study mimics the natural tissue environment due to its soft and flexible structure. This simultaneous use of such materials (hydrogels) with TGF-β1-binding peptides could facilitate PSTI by providing a synergistic effect from the biomaterial and peptide.

VEGF-binding peptides, another category of growth factor-binding peptides, can enhance PSTI by promoting angiogenesis in healing tissues [268,269]. Peptides such as proliferation-related ligand 1 peptide (PR1P) bind to VEGF, stabilizing and prolonging its activity, thereby increasing its local concentration [268]. This makes VEGF more available to interact with endothelial cells, activating signaling pathways that promote EC proliferation, migration, and new blood vessel formation. These newly formed blood vessels supply oxygen and nutrients to the injury site, creating an optimal environment for soft tissue regeneration and integration.

5.4. Anti-inflammatory peptides

Anti-inflammatory peptides represent a class of bioactive molecules capable of modulating the immune response to favor regenerative healing at the soft tissue–implant interface [231,270,271]. Analogous to anti-inflammatory cytokines discussed previously, anti-inflammatory peptides can be immobilized onto implant surfaces to establish a localized immunomodulatory niche that promotes PSTI. By selectively driving macrophage polarization toward the pro-regenerative M2 phenotype, these peptides suppress local inflammation while simultaneously enhancing fibroblast-mediated ECM deposition and epithelial adhesion [228,229]. The dual immunoregulatory and regenerative capability highlights their translational potential as surface biofunctionalization agents for improving PSTI.

For instance, although not originally designed for PSTI, the anti-inflammatory peptide K23 (KAFAKLAARLYRKALARQLGVAA) was incorporated by Zhou et al. [231] into an inflammation-responsive Ti coating (DOPA-P1@P2) to modulate early immune responses through the controlled release of K23. In this design, the K23 peptide was tethered to the surface using a cleavable linker sequence (PVGLIG) that is sensitive to matrix metalloproteinases (MMP-2/9)—enzymes typically upregulated during early inflammation. This configuration enabled K23 to be released in a controlled manner in response to the local inflammatory microenvironment, thereby initiating timely immunomodulation by promoting M2 macrophage polarization and suppressing pro-inflammatory signaling. While primarily explored in the context of bone regeneration and inflammation resolution, its targeted immunomodulatory effects highlight the potential of K23 for future applications in enhancing soft tissue healing around percutaneous and transmucosal implants [231].

Beyond the conventional M1–M2 macrophage polarization framework, persistent inflammation at the tissue–implant interface may also arise from alternative immune pathways, notably the IL-23/T-helper 17 cells (Th17) [272]. IL-23 is a potent pro-inflammatory cytokine that sustains Th17 cell activity and promotes the secretion of IL-17, leading to prolonged neutrophil recruitment, heightened inflammatory signaling, and ECM degradation [273,274]. This pathway represents a distinct immunological challenge and effective immunomodulatory strategies must extend beyond macrophage polarization to also mitigate IL-23-driven chronic inflammation in order to support stable PSTI.

In this context, IL-23 receptor antagonist peptides have emerged as a promising class of bioactive molecules for targeting alternative inflammatory pathways. They act by competitively binding to the IL-23 receptor (IL-23R), which is expressed on immune cells such as Th17 cells and macrophages. By blocking the interaction between IL-23 and its receptor, these peptides inhibit downstream signaling that drives the production of pro-inflammatory mediators and the recruitment of immune cells to the peri-implant region [275]. For example, Pizarek et al. [232] functionalized Ti implant surfaces with nanocoatings of a noncompetitive IL-23 receptor antagonist peptide via silanization. Keratinocytes cultured on these coated surfaces exhibited a significant reduction in IL-23 and interleukin-17A (IL-17A) expression, including under inflammatory stimulation with lipopolysaccharide. Lipopolysaccharide, a bacterial endotoxin derived from the outer membrane of Gram-negative bacteria, is commonly used in vitro to simulate infection-driven inflammation [276]. In addition, macrophages exposed to conditioned media from these keratinocytes displayed increased expression of cluster of differentiation 163 (CD163) and CD206, along with decreased inducible nitric oxide synthase (iNOS) levels, indicating polarization toward a pro-regenerative M2-like phenotype, highlighting their dual role in immunomodulation [232]. These results suggest that IL-23R antagonist peptide nanocoatings can effectively modulate epithelial and immune responses to support soft tissue healing and mitigate the risk of peri-implantitis.

In addition, these anti-inflammatory peptides can promote PSTI by reducing the excessive formation of fibrosis, which can otherwise disrupt healing [277]. Excessive fibrosis results in dense scar tissue that interferes with natural tissue structure, flexibility, and cell alignment, ultimately hindering the seamless integration of implants [278]. By controlling inflammation, anti-inflammatory peptides lower the risk of fibrosis, ensuring smoother healing and improved integration of biomaterials with the surrounding soft tissue [277]. As such, the biofunctionalization of implantable biomaterials with anti-inflammatory peptides can offer a promising approach to enhancing PSTI via minimizing fibrosis and reducing inflammation.

5.5. Antimicrobial peptides (AMPs)

The oral cavity presents a complex and microbially rich environment where persistent exposure to pathogenic bacteria poses a significant risk to the integrity of peri-implant tissues. Such microbial challenges frequently contribute to the onset of peri-implantitis, a major cause of dental implant failure [279,280]. To mitigate early bacterial colonization and reinforce the mucosal barrier, increasing efforts have focused on integrating antimicrobial agents—such as NPs and antibiotics—into implant surface coatings [281,282]. While effective, these strategies face critical limitations, including NPs-associated cytotoxicity [283] and the global escalation of antibiotic resistance [284].

As such, AMPs, which mimic natural host-defense mechanisms [285], have emerged as promising multifunctional agents for dental implant surface modification [286,287]. AMPs are small, naturally occurring molecules that are essential to the innate immune system. Found across a wide range of organisms, including humans, animals, plants, and bacteria, AMPs play a pivotal role in defending against microbial infections [288]. These peptides possess broad-spectrum activity, effectively targets a wide range of oral pathogens, such as Streptococcus mutans, Escherichia coli (E. coli), and Staphylococcus aureus (S. aureus), while their membrane-disruptive mechanisms confer a low risk of inducing bacterial resistance [[289], [290], [291]]. In addition to their antimicrobial role, AMPs inhibit bacterial adhesion and early biofilm formation, which is critical for maintaining a pathogen-free interface during early healing [289,290]. AMPs primarily act by disrupting microbial cell membranes by binding to their negatively-charged components, such as lipopolysaccharides in Gram-negative bacteria or teichoic acids in Gram-positive bacteria [286]. Such interactions compromise membrane integrity, causing permeabilization, leakage of cellular contents, and ultimately, cell death [286].

One of the AMPs explored for preventing dental implant-associated infections is Gorr–L13K antimicrobial peptide (GL13K), a synthetic peptide derived from the human salivary protein BPIFA2 (also known as parotid secretory protein, PSP). GL13K has attracted interest due to its broad-spectrum antibacterial activity, particularly in dental and biomedical contexts [[292], [293], [294]]. In a study by Mutreja et al. [295], a cysteine-modified variant of this peptide, Cys-GL13K, was developed to enable covalent immobilization onto Ti implant surfaces via thiol–maleimide click chemistry. Dental implants coated with Cys-GL13K exhibited strong antibacterial effects against Streptococcus gordonii and Streptococcus mutans, two key pathogens involved in oral biofilm formation and peri-implant infections. Notably, the covalent attachment provided excellent coating stability, withstanding extended ultrasonication without significant peptide loss—highlighting its mechanical durability and promising clinical applicability.

In another study, Shi et al. engineered a multilayer antimicrobial coating on Ti via a LBL assembly strategy, incorporating the synthetic peptide Tet213 covalently conjugated to collagen IV (AMPCol) [286]. Alternating layers of AMPCol and hyaluronic acid were deposited over a chitosan base to create a tunable thin film with enhanced hydrophilicity and surface uniformity (CS-(HA-AMPCol)₁₀). The coating exhibited favorable cytocompatibility with HaCaT keratinocytes and negligible hemolytic activity, confirming its in vitro biocompatibility. Notably, sustained antimicrobial activity was observed against both S. aureus and Porphyromonas gingivalis for up to 29 days, with near-complete inhibition under continuous exposure conditions (Fig. 13a and b). In parallel, early S. aureus biofilm formation was significantly reduced on coatings comprising seven or ten bilayers [286]. While the coating's impact on soft or hard tissue integration was not directly assessed, the inclusion of collagen IV and the resulting biointerface properties suggest a potential to support epithelial attachment and soft tissue compatibility.

Fig. 13.

AMP-biofunctionalized surfaces for PSTI. a-b) Sustained antimicrobial activity demonstrating that P. gingivalis (a) and S. aureus (b) were effectively inhibited by CS-(HA-AMPCol)₁₀ coatings for up to one month (29 Days). Adapted form [286], Scientific Reports 5, 16336 (2015), licensed under CC BY 4.0. c-h) Antibacterial and in vivo anti-inflammatory performance of PDLLA- and PLGA-based HHC-36 sustained-release coatings on Ti nanotubes. c) CFU assays at various bacterial concentrations show reduced colony formation in both PDLLA + HHC-36 and PLGA + HHC-36 groups, confirming in vitro antibacterial efficacy. d) Quantification reveals significantly lower CFU counts in treated groups, with PDLLA + HHC-36 showing superior inhibition at higher bacterial loads. e) In a murine implant model, peri-implant soft tissue response scores were significantly lower in both AMP-coated groups, indicating reduced inflammation. f) Representative implant site images show visible inflammation in the control group, while coated groups display improved soft tissue healing. g, h) Histological and immunohistochemical analysis reveal reduced inflammatory cell infiltration (g) and lower IL-6 expression (h) in AMP-treated groups (∗p < 0.05 vs. control), demonstrating effective immunomodulatory effects. Reprinted with permission from Ref. [298] under CC BY license.

Beyond their antimicrobial role, AMPs actively modulate host immune responses by promoting anti-inflammatory macrophage phenotypes and attenuating excessive inflammation, thereby fostering a pro-regenerative microenvironment. Additionally, certain AMPs have been shown to enhance angiogenesis, improving vascularization and nutrient delivery essential for tissue maturation [296,297]. From a soft tissue perspective, AMPs directly support the adhesion, migration, and proliferation of cells such as gingival fibroblasts, contributing to the formation of a stable epithelial seal [28,298].

For instance, Miao et al. [298]. engineered a dual-functional surface for percutaneous implants by integrating the synthetic cationic antimicrobial peptide HHC-36 into a biodegradable sustained-release matrix composed of poly(D,L-lactic acid) (PDLLA) and PLGA, deposited onto Ti dioxide nanotube substrates. This strategy aimed to overcome two persistent challenges in transcutaneous device integration: bacterial colonization at the epidermal and dermal layers directly interfacing with the implant, and inadequate PSTI—both of which critically compromise long-term implant stability. The coated surfaces exhibited potent antibacterial activity, with dose-dependent inhibition of S. aureus growth, as seen in colony formation assays (Fig. 13c), and quantitatively confirmed by CFU counting (Fig. 13d). Beyond infection control, the coatings demonstrated excellent in vivo biocompatibility. Implants functionalized with HHC-36-loaded PDLLA or PLGA significantly attenuated local inflammatory responses, as evidenced by improved peri-implant soft tissue response scores (Fig. 13e), reduced visible inflammation around implant sites (Fig. 13f), lower inflammatory cell infiltration (Fig. 13g), and a marked decrease in IL-6 positive cells (Fig. 13h). Notably, PDLLA-based coatings elicited a more pronounced anti-inflammatory effect, likely reflecting differences in polymer degradation kinetics and early-phase peptide availability [298].

Another antibacterial peptide with promising applications in PSTI is KSL-W, a synthetic decapeptide (KKVVFWVKFK) originally developed for its antimicrobial stability in the oral environment [28]. In addition to its broad-spectrum antibacterial activity, KSL-W has been shown to exert direct pro-regenerative effects on HGFs in vitro. As reported by Park et al. [28], fibroblasts seeded onto glass slides and treated with KSL-W exhibited enhanced adhesion, which was associated with increased filamentous actin (F-actin) organization. In addition, KSL-W promoted cell cycle progression by increasing the proportion of cells in the S (synthesis phase) and G2/M (gap 2 and mitosis) phases. The peptide also modulated ECM remodeling by upregulating the secretion of MMP-1 and MMP-2 alongside their endogenous inhibitors (tissue Inhibitor of metalloproteinases-1 (TIMP-1) and tissue Inhibitor of metalloproteinases-2 (TIMP-2)), maintaining proteolytic balance. Notably, KSL-W significantly accelerated fibroblast migration in wound closure assays and increased collagen gel contraction, correlating with elevated expression of α-smooth muscle actin (α-SMA), a hallmark of myofibroblast differentiation. Together, these findings suggest that KSL-W not only provides antimicrobial protection but also actively supports key cellular processes involved in soft tissue repair, positioning it as a promising dual-function coating component for transmucosal implants [28].

LL-37 is another AMP that has been reported to enhance PSTI. This peptide, with 37 amino acids and an alpha-helical structure, exhibits a broad-spectrum antimicrobial activity against pathogens like Candida albicans, S. aureus, and E. coli [299]. LL-37 was also reported to reduce bacterial-induced pro-inflammatory cytokine expression in gingival fibroblasts by directly neutralizing bacteria and preventing their binding to fibroblast receptors, thus minimizing inflammation [27,300]. LL-37 can also promote hepatocyte growth factor production by fibroblasts, attributed to its stimulation of MAPK and Extracellular signal-regulated kinase (ERK) signaling pathways [27,301,302]. In a study where LL-37 was coated on polystyrene tissue culture plates via physisorption, it reduced bacterial lipopolysaccharide-induced inflammatory cytokine production and inhibitor of kappa B alpha (IκBα) degradation in HGFs [27]. LL-37 AMP also enhanced HGF viability and proliferation at certain concentrations, activating cell signaling pathways and reducing apoptosis [27]. It also facilitated nutrient absorption and promoted ECM production, supporting cell growth and viability [27].

While these findings highlight the potential of LL-37 AMP in promoting oral health and soft tissue healing, its reliance on physical adsorption may limit long-term stability and efficacy. The durability of LL-37, which may be compromised by enzymatic degradation, mechanical stress, or environmental factors [303], has not been assessed. Furthermore, since LL-37 can influence fibroblast activity and may contribute to excessive collagen deposition or fibrosis [304], its impact on long-term tissue responses remains unclear.

6. GFs-loaded surfaces

GFs are naturally occurring proteins that are essential signaling molecules within the body. They regulate cellular processes, including proliferation, differentiation, and tissue regeneration. GFs exert their effects by binding to specific cell surface receptors, activating intracellular signaling pathways similar to those described for ECM proteins in Section 3.

In the context of PSTI, GFs regulate key biological processes, including wound healing [305], angiogenesis [306], ECM production [307], and immune modulation [308]. The surface biofunctionalization of dental implants using GFs has thus emerged as a promising strategy to enhance their integration with the surrounding soft tissue [309]. The GF-coated surfaces promote re-epithelialization and fibroblast activity. Major GFs involved in PSTI include TGF-β [310], VEGF [311], PDGFs [312], FGFs [313], and hepatocyte GFs [301,311,314]. TGF-β supports fibroblast proliferation and differentiation, enhancing ECM production and regulating the immune response [310]. VEGF promotes angiogenesis, ensuring a steady blood supply to the healing tissues [315]. PDGF recruits repair cells to the injury site and stimulates ECM remodeling [312], while FGF drives fibroblast proliferation, and migration, and facilitates angiogenesis [316]. Hepatocyte growth factor aids epithelial and fibroblast activity, promoting tissue regeneration and re-epithelialization [314]. This section discusses the characteristics of these GFs and their reported or potential roles in facilitating PSTI.

6.1. Platelet-derived GFs (PDGFs)

PDGFs, as a group of dimeric glycoproteins, consist of two polypeptide chains linked via disulfide bonds with two distinct domains: a core domain and a specific receptor-binding domain [317]. There are four types of PDGFs—PDGF-A, PDGF-B, PDGF-C, and PDGF-D—distinguished by their polypeptide chains, which form either homodimers (e.g., PDGF-AA, PDGF-BB) or heterodimers (e.g., PDGF-AB) [317]. PDGF-BB, particularly, has been shown to promote periodontal tissue regeneration in both animal studies and clinical trials [[318], [319], [320], [321]], and the FDA has approved PDGF-BB for periodontal therapy to treat intrabony defects, furcation lesions, and gingival recessions [322,323]. It has been shown that PDGFs contribute to PSTI and wound healing through multiple mechanisms, including stimulating the formation of new blood vessels [324], promoting collagen synthesis [312], and recruiting and expanding fibroblasts [23,323].

PDGFs induce angiogenesis, ensuring the delivery of essential nutrients and oxygen to healing tissues. They enhance angiogenesis by activating signaling pathways that promote EC proliferation and migration; while maintaining the structural integrity of the newly formed vessels by recruiting pericytes, the cells that stabilize and support blood vessels [324].

Enhancing collagen production and alignment during wound healing are other important roles of PDGFs that can result in PSTI [312,325]. PDGFs bind to the receptors on fibroblasts—the primary cells responsible for collagen synthesis—triggering intracellular signaling cascades that promote their proliferation and migration to the injury site. Such increased collagen production at the injury site can enhance the strength and mechanical properties of the growing tissue.

To investigate this effect, recombinant human PDGF-BB (rhPDGF-BB) was adsorbed onto Vicryl sutures via a dip-coating method at varying concentrations (i.e., 0.3, 1, and 10 mg/ml) and applied in a rat Achilles tendon transection model [312]. As shown in Fig. 14a, an initial burst release of PDGF-BB was observed within the first hour of incubation, followed by sustained, dose-dependent release over 48 h. Fig. 14b further illustrates the cumulative release profile, demonstrating that higher coating concentrations led to greater overall release. Histological analysis revealed superior collagen fibre alignment in tendons treated with PDGF-BB-coated sutures compared to controls, as evidenced by polarized light microscopy (Fig. 14c). In vivo assessments confirmed improved biomechanical properties, including greater maximum load and stiffness, in PDGF-BB-treated groups [312]. These findings highlight the therapeutic potential of PDGF-BB-functionalized biomaterials not only for tendon repair but also for enhancing PSTI in applications such as gingival tissue regeneration around dental implants.

Fig. 14.

Biofunctionalization of sutures and abutment surfaces with PDGF-BB and TGF-β2 to enhance PSTI. a) Quantitative release profile of recombinant human PDGF-BB (rhPDGF-BB) from coated sutures at various time points, normalized by suture length (ng/cm). b) Cumulative release of rhPDGF-BB over time, plotted on a semi-logarithmic scale. c) Representative polarized light microscopy images depicting collagen fiber alignment in tendon repair zones using pristine sutures and those coated with increasing concentrations of PDGF-BB (0.3, 1, and 10 mg/mL). Brighter regions indicate enhanced collagen alignment along the load-bearing axis (scale bar = 500 μm), demonstrating superior matrix organization with PDGF-BB-coated sutures [312]. d) Morphological assessment of HGFs surrounding Ti abutments with and without TGF-ß2treatment. TGF-ß2stimulation promotes an elongated, organized fibroblast arrangement perpendicular to the surface, indicative of enhanced cellular orientation. e) Immunofluorescence images showing increased expression of basement membrane and ECM proteins—laminin, FN, and collagen IV—in HGFs treated with TGF-ß2 compared to untreated controls. Reproduced with permission from Ref. [337], J Periodontol, 2001; 72(12):1505–1511. © 2001 American Academy of Periodontology. Published by John Wiley and Sons.

PDGFs can also promote the recruitment and proliferation of fibroblasts through several key mechanisms that involve binding to PDGF receptors and activating signaling pathways [326,327]. Upon receptor binding, PDGFs stimulate tyrosine kinase activity, initiating a cascade of signals that drive fibroblast proliferation and migration [23,326,327], and enhance cell survival, growth, and mobility [23]. Furthermore, PDGFs can serve as a chemotactic cue for fibroblasts, favoring their migration from the surrounding tissues into the injury/implant site where they can proliferate and synthesize ECM components [23]. By harnessing these effects, PDGFs can be strategically utilized to drive soft tissue regeneration and improve epithelial attachment at the implant interface.

6.2. Transforming growth factor-beta (TGF-β)

TGF-β is a homodimeric cytokine consisting of two identical polypeptide chains, each containing around 110 to 140 amino acids that are linked via disulfide bonds, stabilizing the cytokine structure and maintaining its integrity. The three-dimensional configuration of TGF-β with conserved cysteine residues also allow it to have precise interactions with its cognate receptors. These structural features make TGF-β a multifunctional cytokine with beneficial functions in regulating various cellular processes such as proliferation, differentiation, migration, and ECM production [328]. In humans, TGF-β exists in three primary isoforms —TGF-β1, TGF-β2, and TGF-β3— each playing distinct yet overlapping roles in tissue development, repair [328], homeostasis [328], inflammation [329], and fibrosis [328].

TGF-β contributes mainly to PSTI via fibroblast activation and ECM production. It promotes the activation of fibroblasts and the synthesis of essential ECM proteins such as collagen and FN, enhancing the structural integrity and strength of the healing soft tissues through well-defined mechanisms [330]. TGF-β binding to fibroblasts has been shown to promote their proliferation and differentiation into myofibroblasts-cells specialized for producing ECM components. These ECM proteins form the structural foundation of the granulation tissue, which is essential for wound contraction and epithelial cell migration [331]. Granulation tissue forms a favorable matrix for epithelial cells to adhere and migrate due to its richness in ECM components like collagen and FN [332]. In addition, cells within granulation tissue secrete GFs such as TGF-β and EGF, which can further promote epithelial cell proliferation and migration to the injury site [332]. By orchestrating fibroblast activation, ECM synthesis, and epithelial migration, TGF-β plays a central role in reinforcing the peri-implant soft tissue barrier [333].

TGF-β can also support angiogenesis [[334], [335], [336]], by inducing EC proliferation and migration. Furthermore, TGF-β was reported to modulate immune responses by balancing pro-inflammatory and anti-inflammatory signals [310]. It suppresses the activity of pro-inflammatory immune cells, including pro-inflammatory M1 macrophages, neutrophils, and T-helper 1 (Th1) cells, thereby reducing the release of pro-inflammatory cytokines [310]. Simultaneously, TGF-β stimulates the differentiation of regulatory T cells (Tregs), which secrete anti-inflammatory cytokines like IL-10 to prevent excessive inflammation [310]. In addition, TGF-β promotes the polarization of macrophages to their M2 phenotype [310], which is known to minimize excessive inflammation [226].

Although TGF-β has not been extensively used for biofunctionalization of dental implants, its potential to enhance PSTI and wound healing was investigated in a study where HGFs were cultured on Ti abutments [337]. The treated fibroblasts with different TGF-β isoforms shifted from a disorganized sunburst pattern to a structured, elongated alignment perpendicular to the Ti surface (Fig. 14d), reflecting improved organization and attachment. Furthermore, TGF-β significantly enhanced ECM production, notably increasing FN, laminin, and collagen IV expression in HGFs (Fig. 14e). These findings highlight the potential of TGF-β in promoting cell adhesion and ECM production, strengthening the biological interface between HGFs and the biofunctionalized surface.

6.3. Epidermal growth factor (EGF)

EGF is a single peptide chain of 53 amino acids with a three-dimensional structure stabilized by three intramolecular disulfide bridges [338]. It can regulate cell proliferation and differentiation by interacting with the epidermal growth factor receptor (EGFR) via its receptor-binding domain [339]. Upon binding to EGFR, EGF induces receptor dimerization, leading to phosphorylation of the intracellular tyrosine kinase domain and activation of downstream signaling pathways [339].

EGF can contribute to wound healing and PSTI by stimulating cellular activities, including adhesion, migration, and proliferation [340]. It was reported to accelerate the proliferation of keratinocytes and fibroblasts by activating cell-cycle regulators [340,341]. EGF has also been investigated as a bioactive agent for implant coatings to improve epithelial attachment and integration, potentially reducing infection risks and peri-implant complications [342].

To assess the potential of EGF in enhancing PSTI, Pansani et al. [343] investigated the effects of EGF-coated Ti surfaces on HGFs. Using a PEG-based film-forming system, EGF was immobilized onto machined and pre-sanded Ti discs. The EGF-coated Ti surfaces exhibited uniform coverage, as confirmed by fluorescence imaging (Fig. 15a), indicating successful and homogeneous application across the disc surfaces. To assess cellular uptake, gingival fibroblasts seeded on the EGF-coated Ti surfaces displayed notable incorporation of the growth factor within 1 h of contact, and the interaction remained evident at 24 h, with green fluorescence indicating binding to cell membrane receptors (Fig. 15c). Release kinetics analysis showed a burst release profile, with the highest EGF concentration detected immediately after immersion, followed by a gradual reduction over 72 h (Fig. 15b), aligning with the critical timeframe for early cell attachment and signaling activation. Cell viability results showed significantly higher metabolic activity in HGFs cultured on EGF-coated discs compared to uncoated and PEG-only controls, suggesting a stimulatory effect of EGF on cellular function. To more accurately replicate the peri-implant soft tissue environment, the researchers employed a 3D collagen matrix model. HGFs were embedded within a type I collagen gel and cultured on both EGF-coated and uncoated Ti discs. This approach allowed for a more physiologically relevant simulation of the connective tissue surrounding dental implants, enabling the evaluation of cellular behavior in a matrix that closely mimics the native extracellular environment. Confocal microscopy confirmed a uniform EGF film on the Ti surface, and within the 3D model, EGF was internalized by fibroblasts embedded in the matrix (Fig. 15d). Cell distribution analysis revealed enhanced fibroblast density, alignment, and contraction in the collagen matrices on EGF-coated discs compared to controls (Fig. 15e), indicating stronger cellular organization and interaction with the ECM [343].

Fig. 15.

EGF-functionalized surfaces for PSTI. a) Fluorescence microscopy confirmed uniform coverage of EGF on the Ti disc surface using a PEG-based film-forming system. b) The release profile of EGF demonstrated a burst release of EGF into the surrounding medium upon exposure to an aqueous environment. c) Fluorescence photomicrographs showing accelerated adhesion and EGF uptake by HGFs within 1 and 24 h of seeding on EGF-coated Ti. d) Confocal microscopy images of a 3D collagen matrix cultured on EGF-coated Ti discs revealed successful internalization of EGF into the cytoplasm of HGFs. e) The 3D collagen model demonstrated denser fibroblast distribution and greater matrix contraction on EGF-coated surfaces, indicating enhanced fibroblast activity, migration, and ECM remodeling. Reproduced fromRef. [343]with permission from Elsevier. © 2019 Elsevier Ltd. All rights reserved.

EGF-coated Ti surfaces also have shown significant potential in improving the biological response of oral mucosa cells, particularly under the adverse effects of nitrogen-containing bisphosphonates [25]. Bisphosphonates, such as zoledronic acid (ZA) and sodium alendronate (SA), are used to treat conditions like osteoporosis and bone metastases by reducing bone resorption and increasing bone density. However, they can have cytotoxic effects on oral mucosa cells, impairing wound healing, angiogenesis, and cell adhesion [344]. Such effects compromise biological sealing (soft tissue attachment to the implant) and tissue integration, both critical for implant stability and protection against infection [345].

To investigate the protective role of EGF against bisphosphonates, Pansani et al. seeded human keratinocytes and HGFs on Ti surfaces with or without EGF coating. The cells were then exposed to SA or ZA at concentrations of 0.5, 1, and 5 μM [25]. EGF-enhanced cell adhesion, spreading, and viability are shown by increased cytoskeletal labeling of keratinocytes and fibroblasts on the EGF-coated Ti surfaces, even when exposed to ZA and SA bisphosphonates (Fig. 16a–c). This improvement is attributed to EGF stimulating cell anchorage and cytoskeletal organization, which can mitigate the cytotoxic effects of ZA and SA and reduce cell death. HGFs on the EGF-coated surfaces also exhibited enhanced VEGF synthesis, even under ZA exposure (Fig. 16d), which is critical for angiogenesis and tissue healing. Further, EGF downregulated the secretion of MMP-2 in HGFs (Fig. 16e) [25], a regulatory effect with important implications for soft tissue healing and peri-implant tissue stability. MMP-2 facilitates wound healing by degrading ECM components, thereby enabling cell migration, angiogenesis, and tissue remodeling [346]. Although essential in the initial stages of tissue repair, prolonged or elevated MMP-2 activity can disrupt ECM integrity—particularly through degradation of type IV collagen—ultimately compromising tissue stability [347,348]. By suppressing MMP-2 expression, EGF may help preserve ECM architecture and support more stable PSTI.

Fig. 16.

The effects of EGF on oral cell adhesion, viability, and activity. Fluorescence images of a) human keratinocytes and b) HGFs treated with different concentrations of SA and ZA bisphosphonates and cultured on Ti disks with or without EGF coating (original magnification × 100), which show enhanced adhesion of both cell lines on the biofunctionalized surface with EGF, even when exposed to ZA and SA, c) Viability of HGFs, d) VEGF synthesis, and e) MMP-2 expression by HGFs, cultured on Ti discs coated with and without EGF and treated with different concentrations of SA and ZA, demonstrating improved viability of HGFs, their higher VEGF expression, and downregulated synthesis of MMP-2 on the EGF-coated surfaces. Reproduced from Ref. [25] with permission from Springer Nature. © 2021 The Author(s), under exclusive license to Springer-Verlag GmbH Germany, part of Springer Nature.

EGF can also enhance cell migration to the injury site by stimulating cytoskeletal reorganization and cell adhesion dynamics. Such enhancement in the migration of tissue-forming cells to the damaged areas or implant sites can improve PSTI [341]. EGF has been shown to facilitate epithelialization—an essential step in wound healing—by promoting the migration and proliferation of keratinocytes [341,349]. As a result, the scar formation was minimized and the closure of acute oral soft tissue wounds was enhanced, as shown in a canine oral mucosa defect model [349]. Beyond cellular migration, EGF can stimulate the production of VEGF, which supports vascularization by promoting the formation of new blood vessels, delivering essential oxygen and nutrients for soft tissue regeneration and integration [350]. EGF has also been reported to modulate inflammatory responses by suppressing pro-inflammatory cytokines, preventing excessive inflammation, and fostering a regenerative environment conducive to stable soft tissue attachment [229].

6.4. Fibroblast GFs (FGFs)

FGFs, as signaling proteins, consist of 120–140 amino acids and contain heparin-binding domains, which stabilize their interaction with heparan sulfate proteoglycans (HSPGs) on the cells surface, protecting them from degradation [351,352]. They are involved in various biological processes, including embryonic development, tissue repair, angiogenesis, and metabolic regulation [353]. The FGF family in humans consists of 22 members, categorized into paracrine, endocrine, and intracellular FGFs based on their mechanisms of action [354,355]. Among these, basic FGF (bFGF or FGF-2) is one of the most widely studied isoforms of FGF. Distinguished by its basic isoelectric point, FGF-2 is integral to angiogenesis, wound healing, and tissue regeneration, making it a focus of research in regenerative medicine and biomaterials [356].

Similarly to EGF, FGF can stimulate VEGF production, facilitating vascularization critical for soft tissue survival and nutrient supply [311,357,358]. In a study by Gorin et al., FGF-2 was incorporated into a fibrin hydrogel and placed between human tooth slices before being subcutaneously implanted in mice [311]. The findings highlighted the ability of the FGF-2-loaded hydrogel to induce VEGF secretion and enhance angiogenesis.

FGF can also improve PSTI and the sealing of dental implants via its ability to stimulate cellular activities [359]. FGF-2, in particular, was reported to enhance the proliferation and differentiation of various cell types, including fibroblasts [358], epithelial [360], endothelial [361], and mesenchymal [362,363] cells, highlighting its potential in promoting a controlled tissue repair. By promoting the growth and differentiation of gingival epithelial cells and fibroblasts, FGF-2 enhances tissue regeneration and strengthens the biological seal around implants, reducing the risk of microbial infiltration and peri-implant complications [359]. Moreover, its interaction with heparan sulfate proteoglycans in the ECM and on cell surfaces localizes and stabilizes FGF-2, preventing rapid diffusion and enabling sustained signaling for wound healing and soft tissue attachment [359].

ECM production, which involves the generation of ECM components such as collagen, elastin, and FN, is another function of FGFs that contributes to the strength and elasticity of the healing soft tissue [316]. Appropriate ECM secretion ensures long-term functionality of the integrated soft tissue, especially around implants that continually experience a dynamic environment like in joints or oral mucosa. FGFs also facilitate ECM remodeling, enabling tissues to adapt and bind effectively to the surrounding implant—key steps in successful tissue integration [364]. FGFs, particularly FGF7, can also prevent fibrosis and excessive scar formation by promoting epithelial regeneration and controlling fibroblast activity [313,365,366]. For instance, in a fetal wound model, exogenous FGF7 reduced fibrosis and enhanced epithelialization by stimulating epidermis formation and the organization of dense collagen fibers at the wound site [313].

7. Cell-functionalized surfaces for PSTI

Pre-biofunctionalization, or the pre-coating of biomaterials with cells before implantation, is an innovative technique designed to enhance implant integration and functionality [367,368]. This method involves seeding live cells onto biomaterial surfaces before introducing them into the body. The pre-seeded cells create an active biological interface that promotes better interaction with surrounding tissues, modulates immune responses, and accelerates tissue regeneration. By secreting bioactive molecules, such as GFs and ECM components, pre-biofunctionalization guides tissue repair and fosters an immune-modulating environment, reducing inflammation and enhancing healing outcomes [368]. Moreover, this approach is adaptable to various applications as it has been studied in various applications such as vascular grafts and orthopedics materials [369,370]. For instance in orthopedics, pre-biofunctionalization of implants with osteoblasts could accelerate bone healing and improve osseointegration, leading to greater implant stability [370]. A recent study on mandibular bone augmentation compared MSC-seeded bone microparticles with acellular counterparts in a rat model. The MSC-seeded biomaterials demonstrated significantly enhanced bone volume formation, faster remodeling, and improved osteogenesis. By 8 weeks, nearly all the MSC-seeded microparticles had been replaced with new bone, whereas the acellular group retained unincorporated particles [370].

Although the pre-coating of implants with cells has not yet been directly studied to enhance the PSTI of dental implants, a study explored the effect of MSC delivery to promote tissue sealing and healing around dental implants in a rat oral model [371]. The study compared systemic delivery through tail vein injection with local injection into the gingivobuccal fold as illustrated in Fig. 17a [371]. Systemic delivery via intravenous injection led to enhanced epithelial sealing, attributed to greater MSC accumulation around the implant compared to local injection into the gingivobuccal fold. The systemic MSC group showed uniform LN332 staining across the implant-peri-implant epithelium interface with a higher quantity in comparison to the local injection group (Fig. 17b and c). In contrast, the local and control groups had limited LN332 expression, mostly confined to the apical region. Systemically delivered MSCs quickly migrated to the peri-implant tissue and efficiently homed to the apical region, whereas locally delivered MSCs showed delayed migration and remained near the injection site for several days (Fig. 17d and e). Poor vascular access further impedes MSC distribution and the high cell density at the injection site increases apoptosis [371].

Fig. 17.

Cell delivery for PSTI enhancement of dental implants.a) Photograph showing the administration of MSCs into the model rat with the experimental implant. b) Ln-332 distribution in the PIE after MSCs injection. Left: Micrographs show Ln-332 (red) in gingiva around implants in control, local, and systemic groups at 4 weeks. White arrowheads mark areas lacking staining. Blue: diamidino-2-Phenylindole (DAPI) (scale bar = 200 μm). c) Quantification of Ln-332 expression in PIE shows higher levels in the systemic MSC group, suggesting enhanced epithelial sealing. d) Schematic illustration of the tissue structure around the implant, highlighting the buccal and palatal regions. Gray boxes indicate the specific tissue areas analyzed in panels (e, a–d). e) Immunofluorescence images and corresponding graphs showing CD90 (red)/GFP (green) double-positive MSCs in peri-implant tissues following systemic and local administration. White arrowheads denote double-positive cells. Scale bars = 100 μm. (a) Alveololingual sulcus adjacent to the local administration site. (b) Buccal side of the peri-implant mucosa. (c) Palatal side of the peri-implant mucosa (d) Palatal mucosa [371]. Systemic delivery led to faster MSC migration to the peri-implant tissue, while local injection delayed migration. f). Schematic of in vivo experimental design: tooth extraction, implant placement, and in situ injection of GMSCs-laden hydrogels, photo-crosslinked with 405 nm light. g) Hematoxylin and eosin (H&E) staining showing thicker PIE in the P/S6+GMSC group at 4 weeks. Graph quantifies PIE thickness at 2 and 4 weeks. h) Immunofluorescence staining of LAMA3 at the implant-PIE interface shows increased expression in the hydrogel group, supporting enhanced epithelial attachment. The graph provides quantitative data at 2 and 4 weeks. i) BP180 staining at week 4 reveals stronger hemidesmosomal adhesion in the hydrogel-treated group, with quantitative analysis confirming increased expression at 2 and 4 weeks. j) Fluorescence imaging shows gradual degradation of hydrogels over 5 weeks, ensuring sustained cell delivery and aligning with tissue healing progression. Reproduced fromRef. [376]with permission from Elsevier. © 2022 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

Although systemic delivery could enhance epithelial sealing, it presents considerable risks, including off-target cell migration, immune responses, and pulmonary trapping, where cells become immobilized in lung capillaries [372]. Additional concerns include thrombotic events, low targeting efficiency, and short cell survival due to immune clearance or apoptosis. There is also a potential risk of cancer-related complications, as inadequately characterized or prepared stem cells may undergo tumorigenic transformation, leading to uncontrolled proliferation or malignancies [[373], [374], [375]]. Comprehensive long-term studies are necessary to ensure their safety when studying systemic delivery.

To optimize cell delivery to enhance PSTI and address the inherent limitations of systemic and local gingivobuccal fold injections, a promising strategy involves delivering cells in situ around the dental implants after implantation, specifically into the subgingival area, as illustrated in Fig. 17f. Li et al. explored this concept by developing injectable, photo-cross-linkable porous hydrogels using GelMA and silk fibroin methacrylate (SilMA) as carriers for GMSCs [376]. GMSCs were encapsulated within the hydrogels, which were subsequently injected around implants in a rat oral model and photo-crosslinked to localize the delivery. In vivo, GMSCs-laden hydrogels significantly improved peri-implant PSTI and biological sealing. Histology results showed a thick peri-implant epithelium (PIE) (Fig. 17g), while immunofluorescence data at 2 and 4 weeks (Fig. 17h and i) demonstrated increased hemidesmosome-related proteins (laminin-332 α3 chain (LAMA3), BP180) at the implant-PIE interface, strengthening epithelial attachment. Fluorescence imaging (Fig. 17j) confirmed their localized retention around the implants, gradual degradation over time, aligning with the tissue healing timeline and ensuring sustained delivery of therapeutic cells during the tissue repair phase [376].

However, pre-seeding the implants before implantation may provide increased benefits than in situ delivery because the stem cells can be primed toward specific tissue lineages, such as soft tissue cells, before implantation. For example, in a study by Baba et al. [377], three groups were evaluated: (1) MSCs pre-seeded on a 3D woven scaffold and cultured in an osteo-inductive medium, (2) MSCs injected into the scaffold during surgery, and (3) a control group with the scaffold alone. The results showed that pre-seeding PLLA scaffolds with MSCs significantly improved bone regeneration compared to in situ delivery methods. Pre-seeding allowed MSCs to adhere, proliferate, and differentiate into osteogenic cells on the scaffold, facilitating bone matrix deposition prior to implantation. This approach led to more extensive and organized bone formation and could be studied to enhance PSTI of dental implants. Although not originally designed to investigate PSTI, in another study conducted by Laino et al. [378], dental pulp stem cells (DPSCs) were pre-seeded onto Ti implant surfaces, resulting in enhanced cell adhesion, early osteogenic differentiation, and robust matrix mineralization. These findings indicate the potential of cell-functionalized surfaces to modulate biological responses at the tissue–implant interface, offering a promising direction for future strategies aimed at improving PSTI.

8. Multi-agent biofunctionalized surfaces

Multi-agent biofunctionalized surfaces represent a next-generation strategy for enhancing PSTI, offering the capacity to simultaneously present multiple biological signals at the implant–tissue interface. These systems can be broadly classified into two categories: surfaces enriched with platelet-rich hemoderivatives (PRHds), and customizable multi-agent constructs. PRHd-based coatings utilize autologous platelet-derived formulations to deliver a naturally complex milieu of GFs and cytokines [379]; however, their composition is biologically variable and lacks tunability. In contrast, customizable platforms are strategically engineered to integrate defined combinations of biomolecules with tailored functionalities and precisely controlled dosages. These include synergistic combinations—such as multiple pro-proliferative peptide to enhance cellular responses [380]—or complementary pairings, such as coupling AMPs with regenerative factors to simultaneously promote tissue healing and prevent infection [381]. This conceptual framework marks a transition from the empirical application of biologically complex extracts toward the rational design of multifunctional surfaces with controlled and predictable biological outcomes In the following section, both PRHd-enriched and customizable multi-agent strategies will be critically reviewed, highlighting their respective design principles, biological effects, and potential for enhancing PSTI.

8.1. PRHd-enriched surfaces as multi-agent biofunctionalized surfaces

Platelets, or thrombocytes, are small blood cell fragments rich in bioactive molecules such as VEGF and PDGF, important for cellular adhesion, proliferation, and tissue regeneration [382]. Granules within platelets are also stored with cytokines and chemokines, which contribute to their functions in regulating inflammation and immune responses [182]. Platelets can also enhance wound healing and tissue repair by adhering to the damaged area, forming a platelet plug, and releasing GFs and cytokines [383].

To leverage the natural regenerative properties of platelets, platelet-rich hemoderivatives (PRHds) have been identified as advanced biological materials to support tissue regeneration and repair. The primary types of PRHds include platelet-rich plasma (PRP), platelet-rich fibrin (PRF), and platelet lysate (PL), each offering unique advantages for clinical use [384,385]. PRP is a plasma-based solution with a high concentration of platelets, which release GFs to stimulate cell proliferation and promote angiogenesis. PRF, on the other hand, is a fibrin matrix enriched with platelets and leukocytes, enabling a sustained release of GFs while serving as a natural scaffold for cell migration and tissue integration. PL, obtained by lysing platelets, provides a bioavailable supply of soluble GFs to activate cellular regeneration processes. The PRHds are particularly effective in enhancing PSTI [386], accelerating wound healing [387], and boosting fibroblast activity [388], making them valuable tools in implantology. Due to their diverse biological components such as GFs and cytokines, PRHds are considered as multi-agent systems, and surfaces treated with them are classified as multi-agent-loaded surfaces, offering synergistic benefits for tissue repair and regeneration [389]. This section discusses the characteristics and functions of PRHds, highlighting their role in enhancing PSTI.

8.1.1. PRP-functionalized surfaces

PRP is an autologous blood-derived product obtained by centrifugation of whole blood collected in the presence of anticoagulants [390]. This process facilitates the separation of blood into distinct layers, yielding a platelet-rich fraction concentrated with GFs, cytokines, and adhesive plasma proteins such as Fg, FN, and vitronectin [391]. PRP can be applied in either liquid or gel form, depending on the intended application. Liquid PRP is commonly used as a coating to enhance localized tissue regeneration, such as in dental implants [392]. Gel PRP is produced by activating liquid PRP with agents such as calcium chloride or bovine thrombin, resulting in the formation of a fibrin matrix [393]. The gel form of PRP is particularly suited for scaffolding applications and long-term delivery of GFs.

PRP-enriched surfaces can facilitate PSTI by enhancing cell attachment and growth [394,395]. PRP is rich in GFs such as PDGF, TGF-β, and VEGF that can be released from the coated surface [394]. As discussed in Section 5, these GFs enhance PSTI through various mechanisms, including the proliferation of epithelial and fibroblast cells, as well as the production and remodeling of ECM. Despite its advantages, PRP has seen limited investigation for PSTI in dental implants.

In one of only a few studies, PRP was used to coat polymeric expanded polytetrafluoroethylene (ePTFE) membranes and other polymeric substrates to create a platelet-enriched surface [396]. In this study, light microscopy images (Fig. 18a) showed an enhanced attachment of human periodontal ligament fibroblasts (HPLFs) to the PRP-coated membranes. The PRP-coated surfaces also exhibited clusters of closely interconnected HPLFs primarily located on smooth surfaces, whereas non-coated membranes showed sparse or no cellular attachment, as observed in the SEM images (Fig. 18b). On the PRP-coated surfaces, HPLFs demonstrated multiple extensions and were firmly wrapped around the membrane fibbers (Fig. 18c). Similarly, HGFs exhibited improved attachment to the PRP-coated membranes, extending thread-like projections and embedding into the fibrous network. These improvements in cell adhesion and proliferation on the PRP-coated surfaces were attributed to the release of autologous GFs, such as PDGF and TGF-β, from PRP, which can enhance cellular interactions and create a supportive environment for PSTI [396]. While these findings shed light on the potential of PRP in dental implant integration, the interaction between PRP and fibroblasts may vary depending on membrane properties such as resorbability and surface characteristics, which could influence its effectiveness in promoting PSTI [397].

Fig. 18.

PRP-coated surfaces for PSTI; a) Light microscopic images (imaged area of 0.25 mm²), b) SEM image at 500x magnification and c) SEM image at 2000x magnification of human periodontal ligament fibroblasts (HPLFs) on uncoated and RPR-coated membranes, which show improved HPLFs adhesion and proliferation with interconnected clusters and multiple extensions on the PRP-coated membranes [396].

In another study, PRP was immobilized on Ti substrates using mussel adhesive proteins (MAPs), which served as intermediates anchoring the platelets within PRP to the surface [398]. The study aimed to enhance the biological seal at transcutaneous implant sites by improving the response of dermal fibroblasts. The results demonstrated significantly higher fibroblast adhesion and proliferation on MAPs/PRP-coated Ti substrates compared to the uncoated counterparts. This enhancement was attributed to the GFs in PRP, which activate cell integrins and promote the expression of adhesion proteins, thereby improving fibroblast attachment [398]. The GFs in PRP can also stimulate intracellular signaling pathways and cell cycle progression, leading to an increased proliferation rate of fibroblasts. In that study, dermal fibroblasts adhered to the MAPs/PRP-coated surface spread effectively and exhibited enhanced intercellular connections, characterized by numerous filopodia and fibbers at their periphery. The activation of signaling pathways by GFs released from PRP can induce cytoskeletal reorganization in fibroblasts, facilitating filopodia extension and enhanced cell spreading [398]. Although not directly related, the effectiveness of the composite MAPs/PRP-coating in enhancing the adhesion and proliferation of dermal fibroblasts can be extended into other soft tissue cells such as gingival and epithelial cells in the context of dental implants.

8.1.2. PRF-functionalized surfaces

Although PRP is widely used in fields such as dentistry, orthopedics, and aesthetics for regenerative purposes, its potential is hindered by the inclusion of anticoagulants, which are known to suppress tissue regeneration [399]. PRF, a second-generation platelet concentrate, has gained particular attention in tissue engineering and regenerative medicine, addressing the limitations of PRP [400].

Unlike PRP, PRF is obtained through a simplified centrifugation process without anticoagulants, making it fully autologous and highly biocompatible. Blood is spun at a lower speed and longer duration than PRP, yielding a fibrin clot enriched with platelets, GFs, and leukocytes. The fibrin matrix acts as a natural scaffold, supporting cell migration, proliferation, and angiogenesis [401].

PRF, free from anticoagulants, naturally forms a fibrin clot within minutes of blood collection that can act as a three-dimensional scaffold for tissue regeneration [402,403]. Its complete immune biocompatibility and ability to promote angiogenesis can significantly enhance wound healing [404]. PRF can also enhance PSTI via sustained release of GFs [401] that increase cellular adhesion, proliferation, and differentiation [401,405]. These advantages have made PRF a widely utilized biological agent in oral surgery, with applications continuing to expand rapidly.

PRF serves a multifunctional role by also stimulating gene expression pathways associated with tissue regeneration [401]. In a study by Summa et al., machined Ti discs and collagen membranes were treated with PRF lysates to evaluate the adsorption of TGF-β activity and its subsequent impact on fibroblast gene expression. Gingival fibroblasts exposed to the PRF lysates exhibited significant upregulation of TGF-β target genes, including IL11, NOX4, and PRG4, specifically at higher concentrations of PRF lysates (Fig. 19a and b), which are important for ECM remodeling, angiogenesis, and cell differentiation [401]. These findings suggest the potential of PRF in enhancing PSTI by promoting tissue regeneration.

Fig. 19.

Impact of PRF and i-PRF biofunctionalization on HGF gene expression, spreading, and migration.a) Reverse transcription PCR analysis for IL11, NOX4 and PRG4 in HGFs stimulated with soluble PRF lysates with or without TGF- β receptor I kinase antagonist SB431542, showing a significant increase in the expression of TGF-β target genes (IL11, NOX4, and PRG4) in HGFs when exposed to the PRF lysates [401]. b) Reverse transcription PCR analysis for IL11, NOX4 and PRG4 in HGFs incubated with various concentrations of soluble extracts of PRF, which indicates that the effect of PRF on the activation of TGF-β signaling in HGFs is more pronounced at its higher concentrations [401]. c) Fluorescence images, d) average surface area and e) migration of HGFs cultured with either PRP or i-PRF on tissue culture plastic (TCP), smooth (PT) and acid-etched (SLA) Ti substrates, exhibiting an improvement in the spreading of HGFs when cultured on either TCP or PT, and their migration on all three substrates in the presence of either PRF or i-PRF (scale bar = 100 μm) [409].

The promising role of PRF in improving peri-implant PSTI was also reported in another study, where the effects of PRF-conditioned medium on gingival fibroblasts cultured on Ti discs were investigated [406]. The findings revealed a significant increase in FN production by the gingival fibroblasts in the PRF-conditioned groups. Furthermore, the PRF-conditioned medium significantly enhanced the expression of paxillin, AKT, and vinculin at 24 h compared to the controls [406]. Paxillin is a focal adhesion protein that can enhance gingival fibroblast adhesion and spreading [407]. AKT, as part of the PI3K/AKT signaling pathway, promotes fibroblast survival, migration, and ECM remodeling [408]. Another focal adhesion protein, vinculin, strengthens fibroblast-ECM interactions, stabilizing cell adhesion [407]. These effects can cumulatively enhance soft tissue regeneration and sealing around the dental implants.

However, a recognized limitation of PRF is its limited integration with biomaterials due to the fibrin scaffold structure, which lacks the adaptability of the liquid or gel consistency found in PRP [409]. To address this challenge, injectable platelet-rich fibrin (i-PRF) was developed as a liquid alternative to solid PRF. I-PRF was specifically designed to overcome limitations such as the high viscosity of solid PRF, which can hinder its application in narrow or multi-rooted teeth. In contrast to PRF, i-PRF is prepared at lower centrifugation speeds and shorter durations, resulting in a liquid formulation that remains injectable and easy to apply in targeted areas [409]. While i-PRF provides immediate delivery of GFs like PDGF, VEGF, and TGF-β, PRF offers a more gradual release of these bioactive molecules over time, supporting long-term tissue regeneration [409].

I-PRF has been recognized as more effective than PRP in improving PSTI due to its unique preparation method and biological properties [409]. In a study, i-PRF was prepared and applied to the smooth (PT) and roughened (acid-etched, SLA) Ti implant substrates before seeding gingival fibroblasts on the surface. The study then compared the ability of i-PRF to promote cellular interactions, such as fibroblast migration, adhesion, proliferation, and ECM protein expression, to that of standard PRP [409]. As shown in Fig. 19c and d, both PRP and i-PRF significantly enhanced fibroblast spreading on tissue culture plastic (TCP) and smooth Ti (PT), as evident from their improved morphology with extended filopodia and larger surface areas. i-PRF also increased cell migration by over 350 % compared to the controls (Fig. 19e) and significantly boosted cell proliferation, particularly at three and five days of culture [409]. These benefits may be attributed to its ECM-mimetic fibrin structure that serves as adhesion points for fibroblasts and supports enhanced migration [409]. This fibrin structure also results in a sustained release of GFs such as PDGF, VEGF, and TGF-β [410] that can enhance tissue repair and regeneration [409,410]. In contrast, PRP releases these factors more rapidly, resulting in a shorter duration of biological activity [411]. PRF also contains a higher concentration of leukocytes, which play a crucial role in immune modulation and wound healing by secreting additional cytokines and GFs [412]. Leukocytes help create a bioactive environment that promotes angiogenesis, ECM remodeling, and collagen synthesis, further enhancing its regenerative potential [412]. In a study [409], i-PRF is also shown to create a favorable microenvironment for ECM remodeling by increasing collagen type 1 expression in fibroblasts.

Another form of PRHds is advanced PRF (A-PRF), an intermediate platelet concentrate positioned between PRF and i-PRF in terms of viscosity and biological properties [413,414]. A-PRF is created using a modified centrifugation process at moderate speeds (1300 rpm) for 8 min, optimizing the retention of bioactive components, including platelets, leukocytes, and GFs, within a solid fibrin matrix. This process results in a structure that supports cellular activity and facilitates a sustained release of GFs, enhancing its regenerative potential [413]. A-PRF, with its intermediate viscosity, combines the structural integrity of PRF with the adaptability of i-PRF, offering enhanced biological activity and versatility for regenerative applications [413].

In a recent study, i-PRF and A-PRF were compared for their effectiveness in inducing the proliferation of HGFs [413]. While both i-PRF and A-PRF were effective in enhancing cellular growth, i-PRF demonstrated higher levels of IGF-1 release, which is crucial for promoting cell survival, proliferation, and migration. A-PRF, however, excelled in providing TGF-β1 and PDGF-AA, essential for matrix remodeling, angiogenesis, and tissue regeneration [413]. A-PRF also resulted in a more extended release of the GFs in a sustained manner over several days. This can make A-PRF a more efficient and cost-effective solution in regenerative therapies by creating a stable, growth factor-rich microenvironment favorable for enhancing PSTI while reducing the need for repeated applications [413].

Advanced Platelet-Rich Fibrin Plus (A-PRF+), as a new version of A-PRF, can be also used in regenerative medicine and dentistry [415]. While both share similarities in function, there are key differences that distinguish A-PRF+ from A-PRF. In comparison to A-PRF, the preparation of A-PRF + involves lower centrifugal forces and specific time adjustments [416], resulting in an enrichment of bioactive cells, platelets, and GFs. In addition, A-PRF + has a denser and more organized fibrin scaffold than A-PRF, which can support better cellular attachment, migration, and angiogenesis. A-PRF + also exhibits a higher concentration of monocytes and macrophages, which are essential for modulating healing and regeneration processes. Upon implantation, monocytes entrapped within the fibrin matrix are recruited to the wound site and undergo differentiation into macrophages, guided by a cytokine-rich microenvironment shaped by platelet-derived factors such as IL-10, TGF-β, and VEGF [417]. This milieu suppresses pro-inflammatory M1 polarization and promotes the emergence of an anti-inflammatory M2 phenotype, which facilitates inflammation resolution [220], ECM remodeling [418], and neovascularization [419]. In addition to their direct reparative roles, M2 macrophages engage in paracrine signaling with stromal and endothelial cells, enhancing granulation tissue formation and supporting tissue maturation. By modulating the local immune landscape and providing a cell-instructive fibrin scaffold, A-PRF + supports a coordinated regenerative response conducive to high-quality soft tissue repair [220].

Comparison of the regenerative potential of A-PRF+ and i-PRF [414] showed that both formulations promoted stem cells from the apical papilla (SCAP) proliferation and osteo-/odontogenic differentiation; however, A-PRF + demonstrated superior effects, enhancing cell proliferation and migration more effectively than i-PRF. Such enhancements in the proliferation and migration of SCAPs result in a greater population of functional cells at the implant site that can form gingiva or periodontal-like tissues for a tight and protective seal around the dental implants [420]. SCAPs can also differentiate into fibroblasts, endothelial cells, and other connective tissue cell types, further contributing to the formation of soft tissue structures [421].

8.1.3. PL-functionalized surfaces

Platelet lysate (PL) is a cell-free concentrate of GFs obtained by disrupting platelets, typically via freeze-thaw cycles or chemical treatments [394]. Unlike PRP and PRF, which retain cellular and structural components, PL provides a purified growth factor cocktail devoid of fibrin or plasma.

PL has not been extensively used for PSTI in dental implants. However, in a study by Babo et al., it was used to promote periodontal tissue regeneration in a bilayered scaffold made from two distinct layers of a PL scaffold (with and without rat periodontal ligament cells) and a calcium phosphate cement with PL incorporated poly (lactic-co-glycolic acid) (PLGA) microspheres [422]. The results showed that the bilayered scaffold promotes periodontal ligament regeneration and connective tissue attachment by enhancing cell adhesion, proliferation, and metabolic activity of periodontal ligament cells. The bilayered system promoted the regrowth of alveolar bone and functional periodontal tissue, which includes the periodontal ligament, gingiva, and connective tissues, supporting the teeth. The promotion of periodontal tissue regeneration was attributed to the sustained release of GFs from PL that can enhance the activity of periodontal ligament cells [422]. It was also reported that the periodontal ligament cells seeded within the construct were not clearly detected after six weeks and did not directly contribute to new tissue formation. Instead, their effects were likely due to the release of bioactive molecules influencing the host cells [422]. Although not directly focused on dental implants, this study shows the potential of PL for improving PSTI. In this regard, PL can be applied as a bioactive coating to stimulate the activity of periodontal ligament cells, which are essential for soft tissue attachment around implants [423]. PL can be also incorporated into hydrogels or nanostructures that are applied around the dental implants, harnessing its discussed bioactive properties [422] to enhance soft tissue healing and integration.

PL has also shown promise as an alternative to fetal bovine serum (FBS) for expanding dental pulp stromal cells (DPSCs) [424,425]. It facilitated DPSC proliferation, while preserving their functionality and stability [424,425]. DPSCs, as a type of mesenchymal stem cell, can enhance ECM production and reduce bacterial adhesion, thereby improving wound healing and soft-tissue sealing around implants [426]. PL was also reported to maintain the multipotency of DPSCs [424,425], which can enhance the integration of dental implants with surrounding tissues by promoting osteogenesis and angiogenesis [427].

The effect of PL on HGFs was also investigated by enriching serum-free (SF) medium with PL (SF + PL). The findings demonstrated a 1.5-fold increase in HGF proliferation in the PL-enriched medium compared to a FBS-supplemented control, while maintaining cytoskeletal stability and chromosomal integrity [428]. Moreover, SF + PL medium did not induce excessive secretion of ECM remodeling molecules, such as matrix metalloproteinases (MMPs), tissue inhibitors of metalloproteinases (TIMPs), or cytokines. These molecules, when overproduced, can disrupt tissue homeostasis and impair PSTI [429]. Based on these findings, PL can promote the rapid proliferation of HGFs and regulate the secretion of ECM remodeling molecules that are favorable for PSTI.

Another benefit of PL lies in its ability to neutralize hydrogen peroxide (H₂O₂)-induced toxicity and alleviate oxidative stress in gingival fibroblasts [430]. Peroxide-induced toxicity can arise following dental implant placement due to local inflammation caused by surgical trauma, bacterial contamination, or the properties of the implant surface. During this process, immune cells release reactive oxygen species (ROS), including H₂O₂, which induces oxidative stress. Oxidative damage compromises gingival fibroblast viability, leading to necrosis marked by a loss of membrane integrity [428]. It has been shown that PRF lysates, derived from PRF, effectively neutralized H₂O₂-induced toxicity in gingival fibroblasts. The protective mechanism of PRF lysates was attributed to heat-sensitive catalase, which is a temperature-sensitive enzyme that decomposes H₂O₂ into water and oxygen, thereby mitigating its harmful effects [430].

8.2. Customizable multi-agent functionalized surfaces for enhanced PSTI

Biofunctionalizing implant surfaces with multiple agents presents a promising strategy to enhance PSTI by harnessing the synergistic and complementary effects of various bioactive molecules. The success of platelet-rich derivatives (PRDs), such as PRP- or PRF-coated surfaces, is largely attributed to their diverse biological composition, which includes GFs like TGF-β and PDGF, plasma proteins such as Fg and FN, leukocytes, cytokines, and other regenerative molecules. However, optimizing biofunctionalization requires a more targeted approach by precisely selecting and loading specific biological agents to achieve tailored therapeutic outcomes and maximize implant performance.

For instance, combining various peptides, proteins, and GFs can have a synergetic effect on cell adhesion. A foundational study published in 1987 by Lowenberg et al. demonstrated this synergistic effect by integrating PDGF into collagen coatings on porous Ti alloy (Ti6Al4V) discs [23]. The results demonstrated that collagen-coated Ti significantly improved fibroblast attachment and orientation compared to uncoated Ti, which exhibited the lowest attachment index (AI) and orientation index (OI) due to its bio-inert properties. However, the addition of PDGF to the collagen coating further enhanced its performance, showing significantly higher OI values on days 1 and 2 compared to collagen alone [23].

Another study on multi-agent biofunctionalization investigated the combined effects of Ti-binding peptides (TiBP) and FN-derived RGD peptides on silk fibroin coatings applied to Ti surfaces [431]. TiBP are short amino acid sequences engineered or selected to bind specifically and strongly to Ti or TiO₂ surfaces [432]. The dual-functionalized silk–TiBP/RGD coating demonstrated robust results, with a 60 % increase in cell density compared to bare Ti (Fig. 20a and b), better cell adhesion and spreading (Fig. 20c), higher vinculin expression, and numerous well-defined focal adhesion points (Fig. 20d and e). Endothelialization testing, performed using chick aorta explants, demonstrated that the TiBP/RGD coating exhibited the strongest endothelial adhesion, with only 17 % of cells detaching under trypsin treatment, compared to 40 % for bare Ti.

Fig. 20.

Multi-agent biofunctionalized surfaces for PSTI.a) Fluorescence micrographs showing DAPI-stained nuclei of fibroblasts cultured for 48 h on uncoated T (control) and Ti coated with silk–TiBP/RGD 5.3 %, demonstrating increased cell attachment on functionalized surface. Scale bar = 500 μm. b) Quantitative analysis of cell density 48 h post-seeding showing a 60 % increase in fibroblast density on the TiBP/RGD-coated surface compared to bare Ti. c) SEM images of fibroblast adhesion and spreading on unmodified Ti disks and Ti functionalized with silk–TiBP/RGD 5.3 % after 48 and 96 h of culture, highlighting improved cell morphology and coverage on silk–TiBP/RGD-coated Ti compared to uncoated Ti. d) Immunofluorescence images showing vinculin (red), actin (green), and nuclei (blue) in Swiss 3T3 cells cultured for 4 h on uncoated Ti and silk–TiBP/RGD 5.3 % functionalized Ti, revealing higher vinculin expression and more focal adhesion points on the functionalized surface, e) Quantification of focal contact points indicating enhanced cell attachment. Reproduced fromRef. [431]with permission from Elsevier. © 2012 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved. f) Adhesion of fibroblasts on smooth and coated Ti surfaces. Control: polystyrene culture plate; Ti: smooth Ti surface; Ti/MAP: MAP-coated Ti surface; Ti/MAP/PRP: MAP/PRP composite-coated Ti surface. No significant difference was observed between the two coatings, suggesting MAP alone enhances adhesion. g) Proliferation of fibroblasts on smooth and coated Ti surfaces, indicating a synergistic effect of PRP and MAP in promoting cell growth. h) Fluorescence images of fibroblast spreading at 1 h, 2 h, and 1 day, showing greater spreading on MAP and MAP/PRP coatings. Scale bar = 100 μm. Reproduced fromRef. [398]with permission from Springer Nature. © 2015 The Korean Society for Biotechnology and Bioengineering and Springer-Verlag Berlin Heidelber. i) Proliferation rate of HGF-1 cells after 48 h of stimulation with EMD, prAMEL, and TRAP showing the highest proliferation rate across all tested dosages, while prAMEL and prTRAP had moderate effects. j) Real-time migration analysis of HGF-1 cells, with prAMEL exhibiting the strongest migration enhancement, followed by prTRAP, while EMD had minimal impact [435].

The synergistic effect of multi-agent biofunctionalization was further explored in another study, where researchers investigated the complementary roles of MAPs and PRP in biofunctionalizing Ti implants to promote PSTI [398]. Both the single-agent MAP coating and the dual-agent MAP + PRP coating demonstrated significantly enhanced early cell adhesion (Fig. 20f) and spreading (Fig. 20h) compared to bare Ti surfaces, with no significant difference observed between the two coatings. However, fibroblast proliferation results (Fig. 20g) showed significantly higher cell numbers on MAP/PRP-coated Ti surfaces compared to MAP-coated surfaces, which the authors suggested could be due to the synergistic effect of PRP and MAP on cell proliferation [398]. Nevertheless, in this study, PRP was used in combination with MAPs rather than as an independent coating on Ti surfaces. MAPs acted as mediators, anchoring PRP to the Ti surface through their strong adhesive properties, forming a stable and functional coating. Independent evaluation of PRP-coated surfaces could reveal its baseline effects without MAP and allow for direct comparisons with MAP/PRP composite coatings.

Aside from having synergistic effects, multi-agent surface loading can also provide complementary benefits. For example, AMPs can be combined with GFs to simultaneously provide antibacterial activity and enhance host responses [433]. While this approach has not yet been explored for dental implants, it has shown promising results in wound healing applications [434]. For instance, Vijayan et al. developed a PLGA-based nanoparticulate system designed to co-deliver AMPs and GFs, such as VEGF and bFGF [434]. This system demonstrated dual benefits: the AMPs provided broad-spectrum antibacterial activity to combat infections, while the GFs promoted angiogenesis and stimulated cell proliferation [434].

Another important factor in designing multi-agent biofunctionalized surfaces is the careful selection and optimization of dosages for each agent, as well as consideration of their potential interactions. In one study, Wyganowska-Swiatkowska et al. investigated the effects of enamel matrix derivative (EMD), porcine recombinant amelogenin (prAMEL), and tyrosine-rich amelogenin peptide (prTRAP) at different dosages of 12.5 μg/mL, 25 μg/mL, and 50 μg/mL on HGF behavior [435]. EMD exhibited the most pronounced effect on proliferation at all tested dosages (Fig. 20i), while its impact on migration was minimal, with no significant differences observed across concentrations (Fig. 20 j). prAMEL moderately enhanced proliferation (Fig. 20i) in comparison to the control, although significantly less than EMD. However, it notably promoted migration at all dosages compared to other groups, with the strongest effect observed at 25 μg/mL after 72 h (Fig. 20j). prTRAP, similar to prAMEL, showed limited influence on proliferation, significantly less than EMD, but it strongly enhanced migration, surpassing EMD yet remaining less effective than prAMEL after 72 h.

Although EMD contains migration stimulators like amelogenin degradation products (TRAP) [436], it did not demonstrate the same migration-enhancing effects observed with prAMEL and prTRAP, likely due to insufficient concentrations of these stimulatory components. The levels of these components within EMD may fall below the threshold required to achieve the significant migration observed when prAMEL and prTRAP are used individually. Additionally, EMD's complex composition, containing various protein fractions, could result in interactions that modulate or suppress the migration-enhancing effects of prTRAP present in the EMD [435]. Another reason for this lower migration rate observed in EMD groups which is not mentioned in this study could be the presence of PGA in EMD [437]. While PGA is beneficial for wound healing, it may alter the extracellular environment in ways that restrict migration.

However, EMD significantly enhanced cell proliferation compared to prAMEL and prTRAP, probably due to the complex composition of EMD, primarily comprising amelogenin and other proteins, which synergistically promote proliferation but may not effectively activate migration-specific pathways. Proliferation and migration are regulated by distinct signaling pathways, and EMD seems more potent in activating ERK and TGF-β signaling, which drive cell proliferation, rather than pathways involved in adhesion and migration [435]. The addition of prAMEL to the EMD composition could potentially provide complementary effects for future studies, with EMD promoting proliferation and prAMEL enhancing migration.

Therefore, in addition to selecting the appropriate agents, it is important to consider their optimal doses and potential interactions with other components to maximize their effectiveness for cell proliferation, migration, and adhesion.

9. Future directions

Biofunctionalized dental implant surfaces have undergone significant advancements, yet key challenges remain in ensuring stability, scalability, and long-term efficacy under physiological conditions. The harsh in vivo environment accelerates biomolecule degradation and poses risks such as coating delamination, which can compromise integration and implant longevity. Refining biofunctionalization techniques is essential to enhance coating durability, strengthen adhesion, and develop multifunctional strategies that withstand biological and mechanical stress. Understanding how these modifications respond to bacterial exposure, mechanical loading, and oral fluid interactions will be key to predicting their long-term clinical success. Fig. 21 provides an overview of the current challenges, key biofunctionalization strategies, emerging fabrication technologies, and prospective directions, which are elaborated in detail in the following sections.

Fig. 21.

Main challenges, key biofunctionalization strategies, emerging fabrication technologies, and future directions for enhancing PSTI of dental implants.

9.1. Enhancing biointerface stability through scalable covalent strategies: from wet chemistry to plasma-based technologies

Covalent immobilization of biomolecules via wet chemistry is a well-established strategy for soft tissue sealing, offering strong and stable attachment. Covalent linkages can be selectively positioned outside enzymatic recognition domains, thereby enhancing resistance to proteolytic cleavage. In addition, densely packed or crosslinked architectures introduce steric hindrance that limits enzymatic infiltration, while the constrained conformational dynamics of immobilized proteins further suppress enzyme-substrate interactions. Together, these features confer enhanced biochemical resilience, enabling covalently functionalized surfaces to maintain bioactivity and structural integrity under physiologically demanding conditions. This is particularly advantageous in the transmucosal region of dental implants, where interfaces are continuously exposed to fluid shear, enzymatic degradation, and immune surveillance. However, translating these methods into scalable, reproducible, and environmentally sustainable remains a key challenge. Since covalent bonding relies on functional groups, automation with precise reaction control presents a viable path toward industrial scalability. Plasma-based surface modification has emerged as a promising approach, introducing reactive functional groups [[438], [439], [440]] and long-lived radicals [441,442], facilitating covalent attachment of biomolecules. Unlike conventional wet chemistry methods, plasma techniques are inherently scalable, allowing precise control over surface chemistry while eliminating the need for solvents and hazardous waste. Recent developments in this domain to generate uniform plasma-activated biointerfaces on complex 3D structures further enhance their clinical applicability in creating next-generation biofunctionalized implants [[443], [444], [445], [446], [447]].

9.2. Optimizing biomolecule dosage and spatial presentation to prevent adverse tissue responses

An important area of future study is optimizing the dosage of biological agents to achieve effective PSTI without adverse effects. For instance, TGF-β as a tissue regeneration agent can promote ECM production and cellular differentiation. However, excessive TGF-β levels can provoke an overactive immune response, increasing macrophage density in the injured area. This heightened macrophage activity may inadvertently lead to fibrosis, characterized by excessive scar tissue formation, which can compromise the functional integration of soft tissues [448]. Similarly, achieving an optimal thickness is essential when functionalizing surfaces with proteins. For example, excessively thick collagen layers can result in steric hindrance [170], creating physical barriers that block cells from adhering to the surface by preventing their adhesion molecules from reaching the binding sites [170]. Thus, precise control of biomolecule concentration and spatial distribution remains critical in designing biofunctionalized surfaces that support stable and predictable soft tissue attachment.

9.3. Integrating hierarchical topographies with biochemical cues to mimic the native soft tissue interface

Another promising direction lies in integrating physical and biochemical cues to enhance PSTI. By optimizing microgroove dimensions [449] in combination with biochemical coatings, such as FN or laminin, implants can be tailored to mimic the native ECM environment [159]. Before biofunctionalization, hierarchical surface topographies can be created using techniques, such as plasma [450], electrochemical anodization [451,452], or femtosecond-laser restructuring [453], to produce multi-scale surface features, such as micro- and nano-scale patterns. These structures mimic natural biological designs, enhancing properties like adhesion, cell growth, and electrochemical activity.

9.4. Functional application of hydrogel coatings with optimized thickness

A key advancement in biomaterials for both in-vitro and in-vivo applications is the transition from 2D to 3D structures to better replicate the physiological environment. Hydrogel coatings can provide excellent 3D microenvironments for cells, allowing biomolecule encapsulation, protection from harsh conditions, and controlled release [[454], [455], [456]]. However, the thickness of hydrogel coatings is also a key factor, as excessive thickness can hinder cell adhesion [32] while potentially reducing mechanical stress on dental implants [457]. This balance becomes particularly important for small, precise implants like dental implants, where achieving the optimal coating thickness is essential for preserving functionality. Future research should aim to optimize hydrogel thickness to strike a balance between promoting cell adhesion, reducing mechanical stress, protecting biomolecules, and enabling their controlled release [457,458].

9.5. Harnessing vascularization for PSTI

Vascularization represents an important yet underexplored factor in the maturation and long-term stability of peri-implant soft tissues. While its role in bone regeneration is studied [459], limited studies have addressed how biofunctionalization can enhance vascularization for better PSTI [34]. Local vascular networks not only facilitate nutrient exchange and immune modulation but also support ECM organization and soft tissue maturation [460]. Future investigations should focus on tailoring approaches to selectively promote vascularization within soft tissue environments, while minimizing pro-inflammatory cues, to advance the development of biointerfaces optimized for durable mucosal integration.

9.6. Simultaneous soft and hard tissue integration via dual immunomodulation strategies

Building on the central role of immune regulation in tissue healing, future implant designs may benefit from the strategic incorporation of anti-inflammatory biomolecules, such as cytokines and immunoregulatory peptides, that orchestrate a pro-regenerative microenvironment across both soft and hard tissue interfaces [217,255,461]. Although current strategies often focus on either mucosal or osseous integration in isolation, their underlying biological processes are governed by overlapping, inflammation-mediated pathways. A unified immunomodulatory approach holds considerable translational potential by concurrently supporting PSTI and bone regeneration. Key mediators such as IL-4, IL-10, and IL-13 have demonstrated the ability to reprogram macrophages toward an anti-inflammatory M2 phenotype, thereby promoting ECM synthesis, epithelial adhesion, and osteogenic activity [217]. The localized presentation of these cytokines at the implant surface—via covalent immobilization or controlled-release systems—may enable spatially defined immune modulation. Such strategies align with the emerging paradigm of immune-instructive biomaterials [217,270,462] that actively engage endogenous repair pathways to enhance long-term clinical outcomes.

Nonetheless, a key challenge in immune modulation lies in its potential to attenuate the host's antimicrobial response. Suppressing pro-inflammatory signaling may impair the recruitment and activation of immune cells responsible for early microbial defense [463,464]. To address this, a combined strategy integrating immunoregulatory agents with antimicrobial elements—such as bactericidal peptides or antibacterial nanocoatings—may prove essential to preserve host protection while fostering tissue regeneration [465,466].

9.7. Multi-agent biofunctionalization for regenerative, antimicrobial, and immunomodulatory performance

Future advancements in dental implantology can focus on the biofunctionalization of implant surfaces using a combination of different types of proteins, peptides, and GFs to enhance their performance and ensure effective tissue integration. By strategically combining these agents, future implants can simultaneously promote cell adhesion, proliferation, and migration while actively minimizing the risk of infection [398,431,467]. For example, integrating GFs into implant surfaces is expected to improve cellular responses, while incorporating antibacterial peptides can provide robust antimicrobial activity to prevent infections. Furthermore, immune modulation will be an indispensable area of research, with an emphasis on designing surfaces that reduce chronic inflammation and foster a pro-regenerative environment to support successful tissue integration [229,468]. Developing multi-agent biofunctionalized surfaces tailored to enhance host cell adhesion, mitigate inflammation, and combat infection represents a promising direction for next-generation implants. Incorporation of antibacterial biomolecules, such as AMPs, imparts an additional layer of functional defense, facilitating the development of multifunctional surfaces capable of concurrently promoting tissue regeneration, modulating inflammation, and resisting bacterial colonization.

9.8. Controlled and time-dependent delivery strategies: temporal release design for multifactorial systems

To further enhance the therapeutic potential of biofunctionalized implant surfaces, future research should focus on integrating multi-agent systems with controlled and time-dependent delivery strategies. By combining multiple bioactive agents with sustained and programmable sequential release mechanisms, it will be possible to more closely replicate the temporally regulated biological cascades that govern soft tissue regeneration [469]. The next generation of surface engineering should prioritize the development of LBL assemblies, biodegradable and stimuli responsive hydrogels, and nanocarriers with tunable degradation profiles to enable precisely timed delivery of therapeutic cues [[469], [470], [471]].

Integrating smart materials that respond to environmental stimuli represents a compelling avenue for further advancements in soft tissue bioengineering. While no research has explored the use of stimuli-responsive materials to biofunctionalized the surface of dental implants for PSTI, these materials hold significant potential in this domain. Stimuli-responsive coatings can dynamically adapt to the local oral environment, releasing bioactive agents such as GFs, AMPs, or ECM-mimicking molecules in response to triggers like pH changes, mechanical forces, electrical stimulation [472] or enzymatic activity [473]. For example, a pH-sensitive coating can be designed to release antibacterial agents to combat peri-implantitis [474]. At the same time, enzyme-responsive materials offer the opportunity for GF delivery by enabling on-demand release, stabilization, and localized activation [475].

9.9. Addressing gaps in the in vitro and in vivo evaluation of biofunctionalized implant surfaces

Future research on biofunctionalized surfaces should involve comprehensive in vitro analyses of their interactions with key immune and gingival cells, including macrophages, fibroblasts, and epithelial cells. A systematic evaluation of how these surfaces modulate immune responses, regulate cell adhesion, and influence proliferation and migration is essential for developing implant coatings that achieve stable and functional PSTI. In addition, although large animal models such as domestic pigs and minipigs have been used to assess PSTI of non-biofunctionalized implant surfaces [476,477], or to analyze hard tissue integration of biofunctionalized surfaces [478,479], their application in evaluating PSTI of biofunctionalized coatings remains limited. Addressing this gap through targeted in vivo studies is crucial to validate the translational potential of these surface modifications.

9.10. Gradient biomolecule delivery mimicking native tissue transitions for region-specific integration

Another important aspect to consider in future designs is gradient biomolecule loading on dental implant surfaces, offering a region-specific approach to enhancing soft and hard tissue integration while preventing infection [480]. For example, in tendon-to-bone or cartilage-to-bone applications, gradients guide the differentiation of MSCs into fibroblasts, chondrocytes, or osteoblasts for seamless tissue transitions [481]. By mimicking the natural transitions in biomolecule concentration across these interfaces, gradients guide cellular behaviors such as migration, adhesion, and differentiation [481]. For example, GFs like BMP-2 and TGF-β direct MSCs to differentiate into fibroblasts, chondrocytes, or osteoblasts, ensuring structured tissue development [481]. For dental implants, this strategy could involve applying higher concentrations of agents such as AMPs, GFs, and adhesion proteins in the transmucosal region to support PSTI and protect against infection. Deeper within the bone, the concentration of these agents could be reduced, with a focus on incorporating biomolecules specifically designed to promote the adhesion and activity of cells associated with hard tissue [482].

9.11. Unlocking the potential of FDA-approved therapies

Building on currently FDA-approved biomaterials for soft tissue regeneration presents an untapped opportunity for implant biofunctionalization. For example, fibrin glue [213,214] has been widely used in wound healing and tissue repair, and PDGF-BB [322,323] has been approved for periodontal therapy. However, their potential for enhancing PSTI in dental implants remains unexplored, necessitating targeted research to evaluate their long-term efficacy and clinical applicability.

In addition to FDA-approved biomolecules, clinically validated cell-based therapies in regenerative medicine offer promising avenues for implant biofunctionalization through direct cellular integration at the material surface. For instance, skin substitute products like Apligraf [483] (collagen matrix cultured with fibroblasts and keratinocytes) and Dermagraft [484] (fibroblasts seeded on a bioabsorbable polyglactin mesh scaffold) are routinely used in clinical practice for burn wounds and chronic ulcers. Building on this concept, pre-seeding implant surfaces with host-derived cells, such as MSCs, presents a compelling strategy to enhance cell adhesion, proliferation, and differentiation, while concurrently mitigating inflammation and infection risks at the implantation site. Translating such cell-based approaches to dental implants may significantly improve both hard and soft tissue integration, thereby increasing long-term implant success rates.

9.12. From lab to clinic: stabilized and immunocompatible biomolecule delivery for scalable and translational bioactive coatings

A persistent obstacle in the development of biofunctionalized implant surfaces lies in the temporal mismatch between the degradation kinetics of bioactive coatings and the protracted course of soft tissue remodeling. The establishment of a stable epithelial seal and functional PSTI unfolds over 6–8 weeks through a cascade of overlapping phases, including inflammation, ECM deposition, cellular proliferation, and tissue maturation [485]. In contrast, coatings incorporating biologically active molecules often degrade rapidly in vivo due to their intrinsically short functional half-lives, limited thermal and structural stability, and heightened susceptibility to enzymatic breakdown under physiological conditions. For example, FGF-1 retains bioactivity for only ∼1 h in serum at 37 °C [486], whereas bFGF [487] and VEGF exhibit even shorter half-lives of approximately 3 and 50 min [488], respectively, following systemic administration. Such pronounced kinetic mismatch between biomolecule stability and the extended timeline of tissue regeneration poses a significant barrier to effective therapeutic delivery. To overcome this challenge, emerging delivery systems—including biodegradable polymer matrices, self-assembling peptide hydrogels, and NP-based platforms—have been designed to stabilize labile biomolecules and modulate their release kinetics in synchrony with the healing cascade. Notably, collagen-derived carriers offer distinct advantages by integrating inherent bioactivity with structural features that preserve molecular conformation and shield encapsulated agents from premature enzymatic degradation. This dual capacity supports sustained bioavailability and prolonged therapeutic efficacy, aligning molecular delivery with the dynamic and temporally orchestrated process of soft tissue remodeling [129].

In parallel, efforts to enable the clinical translation of biofunctionalized coatings must also prioritize strategies that are not only biologically effective but also scalable and compliant with regulatory standards. One promising avenue is the incorporation of clinically validated stabilization techniques, such as PEGylation, into implant surface engineering. PEGylation—the covalent attachment of PEG chains to therapeutic biomolecules—offers a well-established means of enhancing systemic stability, reducing immunogenicity, and prolonging circulation time [489]. By increasing hydrodynamic size and forming a protective hydration shell, PEGylation confers resistance to enzymatic degradation and renal clearance, while improving solubility and biocompatibility [489,490]. Importantly, this strategy is consistent with established pharmaceutical manufacturing practices and benefits from a well-defined regulatory framework, facilitating its clinical translation. The clinical success and FDA approval of several PEGylated therapeutics further underscore its translational potential and regulatory viability [491]. Integrating such scalable modifications into implantable systems could help bridge the gap between laboratory innovation and clinical implementation, facilitating broader adoption of advanced bioactive coatings.

To facilitate the clinical translation of biofunctionalized implant surfaces, future designs must proactively address potential immune responses elicited by heterologous proteins, peptides, or synthetic components. The immunogenicity of non-autologous biomolecules remains a critical obstacle to achieving stable PSTI and sustained implant performance [492]. One promising approach involves pre-coating implant surfaces with autologous formulations, such as platelet-derived concentrates, which not only ensure immunological compatibility but also deliver a rich supply of regenerative cytokines and GFs that promote early healing and immune regulation—while aligning with scalable manufacturing practices. For heterologous systems, additional strategies include encapsulating proteins within immunoizolating carriers such as hydrogels, liposomes, or PLGA microspheres [493,494]. Furthermore, surface modification with PEGylation or zwitterionic polymers can create steric barriers that reduce antigen recognition and prevent immune activation [495].

9.13. Harmonizing surface bioactivity with surgical practice

Recent advances highlight the need to align biomaterial innovation with clinical protocols. The success of biofunctionalized implant surfaces relies on both biological performance and compatibility with surgical handling. Integrating adaptive surface technologies with minimally invasive approaches may enhance mucosal stability, reduce complications, and improve outcomes. Implant placement typically follows either a one-stage or two-stage surgical protocol [496]. In one-stage procedures, the implant and abutment are placed simultaneously, allowing trans-mucosal healing. Two-stage approaches, by contrast, involve initial submersion of the implant and later attachment of the abutment in a second intervention. Each protocol presents distinct implications for soft tissue healing. One-stage protocols offer the advantage of preserving the developing mucosal seal by avoiding additional surgical trauma. In contrast, two-stage procedures can disrupt early healing events that are critical for soft tissue stability [496]. Particularly detrimental is the repeated disconnection and reconnection of abutments, which induces mechanical disturbance of the peri-implant mucosa and facilitates microbial ingress [497]. Such disruptions compromise the biological seal and are associated with increased risk of peri-implant inflammation. The “one-time, one-abutment” protocol—wherein the definitive abutment is placed at the time of surgery and remains undisturbed during healing—has demonstrated promise in maintaining the structural and functional integrity of the soft tissue interface [498]. Future surface design strategies should therefore align with minimally invasive and low-disruption surgical protocols to preserve mucosal integrity, enhance soft tissue resilience, and optimize clinical outcomes.

Looking ahead, advances in surface bio-engineering strategies and scalable fabrication technologies, coupled with the incorporation of newly designed bioactive and smart materials, hold great potential to transform the development of the next generation of dental implants, which rapidly and robustly integrate with soft tissue while ensuring long-term clinical success.

10. Conclusions

Achieving stable PSTI remains a fundamental challenge in dental implantology, directly influencing long-term implant success and infection resistance. Surface bioengineering through attaching bioactive molecules to implant surfaces has emerged as a promising strategy to address these challenges by actively modulating cellular interactions at the implant-soft tissue interface.

We discussed the underlying mechanisms regulating biomolecule-implant interactions in terms of adhesion dynamics and mechanisms of surface attachment. Various biomolecule immobilization techniques were critically evaluated for their effectiveness and limitations in achieving stable and functional integration. While physical adsorption remains a practical approach, its susceptibility to desorption limits long-term stability. Covalent attachment methods, including wet chemistry and plasma-based surface modifications, offer enhanced durability but require further refinement for clinical scalability.

This review also critically evaluated a diverse range of bioactive agents applied for dental implant biofunctionalization, focusing on PSTI, and highlighting their distinct advantages and challenges. ECM-derived proteins such as collagen, FN, and laminin facilitate cell adhesion and tissue repair but may suffer from rapid degradation and inconsistent coating stability. Bioactive peptides offer a targeted approach by enhancing cell proliferation while mitigating bacterial colonization and immune dysregulation; however, their short half-life and susceptibility to enzymatic degradation remain challenges. GFs such as TGF-β, VEGF, and EGF promote angiogenesis, ECM synthesis, and wound healing, accelerating PSTI; however, their uncontrolled release can lead to fibrosis or excessive inflammation. Platelet-derived formulations (PRP, PRF, and PL) provide a multi-faceted biological approach to support tissue healing, but their efficacy can vary depending on preparation methods and patient-specific factors. Multi-agent-loaded surfaces present an emerging strategy to achieve synergistic effects. Yet, their design requires careful control over relative surface-attached concentrations and release kinetics to maximize therapeutic efficacy while avoiding unintended cellular responses.

Moving forward, multi-functional implant surfaces integrating optimized combinations of proteins, peptides, and GFs present a promising avenue for improving clinical outcomes. When coupled with emerging advanced fabrication technologies—including plasma-enabled surface modifications, stimuli-responsive coatings, and scalable biofabrication strategies—these approaches hold great potential for precisely engineering implant surfaces to achieve rapid and durable PSTI.

CRediT authorship contribution statement

Ghazal Shineh: Writing – original draft, Visualization, Methodology, Investigation. Leila Mamizadeh Janghour: Writing – review & editing, Investigation. Yiyun Xia: Visualization, Methodology. Jiayan Shao: Visualization, Investigation. Karan Gulati: Writing – review & editing. Giselle C. Yeo: Writing – review & editing, Supervision, Funding acquisition. Behnam Akhavan: Writing – review & editing, Supervision, Funding acquisition, Conceptualization.

Ethics approval and consent to participate

Ethics approval and consent to participate are not applicable to this review article.

Declaration of competing interest

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.