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Review

Protein Adsorption and Cell Adhesion on Metallic Biomaterial Surfaces

by
Satoshi Migita
* and
Masaki Sato
Graduate School of Science and Engineering, Yamagata University, 4-3-16 Jonan, Yonezawa 992-8510, Japan
*
Author to whom correspondence should be addressed.
Adhesives 2025, 1(4), 15; https://doi.org/10.3390/adhesives1040015
Submission received: 29 September 2025 / Revised: 14 November 2025 / Accepted: 19 November 2025 / Published: 18 December 2025
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Abstract

Metallic biomaterials play essential roles in modern medical devices, but their long-term performance depends critically on protein adsorption and subsequent cellular responses at material interfaces. This review examines the molecular mechanisms governing these interactions and discusses surface modification strategies for controlling biocompatibility. The physicochemical properties of oxide layers formed on metal surfaces—including Lewis acid-base chemistry, surface charge, surface free energy, and permittivity—collectively determine protein adsorption behavior. Titanium surfaces promote stable protein adsorption through strong coordination bonds with carboxylate groups, while stainless steel surfaces show complex formation with proteins that can lead to metal ion release. Surface modification strategies can be systematically categorized based on two key parameters: effective ligand density (σ_eff) and effective mechanical response (E_eff). Direct control approaches include protein immobilization, self-assembled monolayers, and ionic modifications. The most promising strategies involve coupled control of both parameters through hierarchical surface architectures and three-dimensional modifications. Despite advances in understanding molecular-level interactions, substantial challenges remain in bridging the gap between surface chemistry and tissue-level biological performance. Future developments must address three-dimensional interfacial interactions and develop systems-level approaches integrating multiple scales of biological organization to enable rational design of next-generation metallic biomaterials.

1. Introduction

Metals have been used as biomaterials from ancient times to today because of their excellent mechanical properties and corrosion resistance [1]. In ancient Egypt, around 1500 BC, gold and silver were used for dental treatment [2]. Records show that bronze artificial limbs were used in ancient Rome [3]. Modern biomedical metals started with stainless steel in the early 20th century [4,5]. The biocompatibility of titanium was discovered in the 1940s, and clinical applications began in the 1960s [6,7,8]. Titanium alloys, especially Ti-6Al-4V, started to be used in the 1960s. These alloys provided better mechanical strength while maintaining good biocompatibility. Today, titanium alloys and many other metals are still essential materials for medical devices.
The development of metallic biomaterials is important for treating musculoskeletal diseases that are common in modern society. Changes in work styles and population aging have increased musculoskeletal diseases [9,10]. The chronic pain from these diseases can contribute to mental health problems, leading to reduced quality of life and work productivity [11]. As a result, treatments for musculoskeletal diseases like osteoarthritis are increasing. The demand for metal devices is also increasing for disc and spinal fixation [12].
The development of high-performance metal materials is required to improve healing effects and shorten treatment time. Titanium alloys and stainless steel are widely used now because they have excellent mechanical strength and corrosion resistance [1,5]. However, these materials need several months for osseointegration [13]. They also have risks of complications such as infections and chronic inflammation [14]. Although rare, metal hypersensitivity reactions (allergies) have been clinically reported, and they have also been suggested to potentially affect the long-term stability of implants [15]. For these reasons, improving biocompatibility and adding biofunctionality are now required, for example, tissue compatibility and antibacterial properties [16]. To design such advanced materials, we must understand how metal surfaces interact with biological tissues at the molecular level [16].
After metal implantation, water and ions are immediately adsorbed on the surface, followed by protein adsorption [17]. The protein layer changes over time, and this phenomenon is called the Vroman effect [18]. This initial biological reaction process controls later cell adhesion and tissue integration. It ultimately affects the long-term stability of implants. However, it is difficult to predict protein adsorption and cell adhesion patterns [19]. The effects of oxide layer properties on protein adsorption and cellular responses are not well understood [20,21].
This review presents the molecular mechanisms of interactions between metal surfaces and biological systems. It also discusses surface modification strategies to control these interactions. First, we present the basic physicochemical properties of oxide layers formed on metal surfaces. Then, we analyze how these properties influence protein adsorption behavior, using titanium, stainless steel, and cobalt-chromium alloys as representative examples. Next, we explore cell adhesion mechanisms through protein layers and discuss surface treatment strategies to enhance biocompatibility. These insights will help design better metallic biomaterials for the future.

2. Physicochemical Properties of Surface Affecting Protein Adsorption

Physicochemical properties of surfaces control the initial adsorption of proteins. Therefore, understanding these properties is important for controlling protein adsorption and potentially improving biocompatibility and biofunctionality of metals.
Surface states on solid materials can act as Lewis acids or Lewis bases, significantly affecting protein adsorption [22]. Coordinatively unsaturated sites on oxide surface act as Lewis acids and are enriched at steps or in the vicinity of oxygen vacancies [23]. At physiological pH, most carboxy groups are deprotonated carboxylates (-COO−), which donate electron pairs to Lewis-acidic, low-coordinated surface metal centers to form inner-sphere coordination complexes [24]. In parallel, surface oxygens and hydroxy groups act as Lewis-basic sites. NH3+ of amino acids forms strong hydrogen bonds with these sites rather than coordination bonds [25]. These acid-base interactions can cause quasi-irreversible protein adsorption and structural changes. Desorption is possible under specific conditions such as pH changes, competing ligands, or altered ionic strength [26,27]. The density and distribution of these sites depend on material composition, crystal structure, and surface treatment conditions, which directly influence protein selectivity and stability. Therefore, controlling surface acid-base properties through rational design represents an effective strategy for tuning protein-surface interactions [28]. Water adsorbs either molecularly (H-bond) or dissociatively to yield terminal Ti-OH and H+ on neighboring oxides (Figure 1) These modes are governed by Lewis/Brønsted acid-base chemistry at coordinatively unsaturated sites.
Hydroxyl groups on solid surfaces typically become charged in aqueous solutions through protonation/deprotonation equilibria (Figure 2). The density of surface charge depends on pH and electrolyte conditions. The pH at which the apparent charge becomes zero is called the point of zero charge (PZC) [29]. Each material has a characteristic PZC value that reflects its surface acid-base properties. Negatively charged surfaces may inhibit the adsorption of negatively charged proteins through electrostatic repulsion, while positively charged surfaces may enhance it [30,31]. However, protein adsorption is controlled by multiple factors including hydrophobic interactions and van der Waals forces, not only electrostatic interactions [32]. Surface charge affects not only the adsorption amounts but also the kinetics of the adsorption process and protein conformation [33,34]. Strong electrostatic interactions may cause protein structural changes in the initial stage of adsorption [26].
Surface charge can be modulated through various approaches including pH control, ionic strength adjustment, and surface functionalization with charged groups [35]. These strategies provide methods for tuning protein adsorption behavior across different material systems.
Surface free energy (SFE) is commonly estimated from static contact angles because it cannot be measured directly. Representative models for calculating SFE are Owens-Wendt-Rabel-Kaelble (OWRK) and van Oss-Chaudhury-Good (vOCG). OWRK divides SFE into dispersive and polar components, while vOCG splits the total SFE into a Lifshitz-van der Waals (γLW) term and acid-base (γAB) term. The latter is further decomposed into an electron-acceptor (γ+) and an electron-donor (γ−) parameter. Choice of model and probe liquids (e.g., the routine use of diiodomethane as “apolar”) as well as measurement protocol can bias SFE estimates, so reporting conditions matters [36,37]. The balance of SFE components modulates protein adsorption by tuning specific Lewis acid–base contacts and noncovalent interactions at the interface, thereby affecting adsorption amounts and conformations [17].
Wettability often correlates with the polar component of SFE in OWRK analysis. Protein adsorption on hydrophilic surfaces is influenced by the polar component of SFE [38]. Hydrophobic surfaces generally accelerate protein adsorption via hydrophobic interactions. Globular proteins such as albumin show maximum adsorption amounts on moderately hydrophobic surfaces. Hydrophobic interactions can induce conformational changes of proteins, exposing internal hydrophobic amino acid residues that enhance surface interactions. Interestingly, the relationship between wettability and protein adsorption is non-monotonic [30]. Protein adsorption is suppressed on both superhydrophobic and superhydrophilic surfaces. Superhydrophilic surfaces suppress adsorption because strong hydration layers cannot be displaced by proteins [39], while superhydrophobic surfaces create air-water interfaces that prevent protein-surface contact [40]. Machine learning-based analyses further show that increasing the density of hydrophilic polymer brushes, which strengthen interfacial hydration, effectively prevents protein adsorption [41].
The permittivity (dielectric constant) of solids also modulates initial protein adsorption via image charge interactions. The permittivity discontinuity between water (εr = 78–80, 25 °C) and solid surfaces creates electrostatic image forces. When the surface εr is lower than water, repulsion occurs, while higher values cause attraction. Most solid materials have εr values lower than water (Table 1), suggesting that image charge repulsion often dominates protein-surface interactions. In contrast, TiO2 can exceed water εr values depending on crystal phase: anatase (40–55) and rutile (80–170) [42]. Therefore, rutile-rich films can create attractive image forces and have been correlated with enhanced FXII activation, fibrin formation, and platelet adhesion at otherwise comparable roughness and surface energy [43].

3. Protein Adsorption on to Metallic Biomaterials

The chemical composition of oxide layers on metallic biomaterials is the most fundamental factor determining protein adsorption [17,18]. In particular, the elements and their oxidation states dominate the coordination chemistry of surfaces and determine the interactions with proteins [48].
Amino acid residues in proteins form metal-protein complexes. For example, imidazole rings of histidine (His) in proteins coordinate to metal ions at physiological pH [49]. Also, sulfur atoms of cysteine (Cys) act as soft bases and preferentially bind to soft metal centers (e.g., Cu+, Ag+, Au+) [50]. In contrast, carboxy groups of aspartic acid (Asp) and glutamic acid (Glu) act as hard bases that bind strongly to hard, high-valent oxophilic centers such as Ti4+ on TiO2 surfaces, typically in inner-sphere bidentate on bridging carboxylate motifs [24,51]. For trivalent chromium, Cr3+ is a hard acid that forms robust carboxylate complexes, consistent with this trend [52].
These differences in protein adsorption determine the cell attachment behavior on each metal material. The distinct surface characteristics of these materials (Table 2) lead to different protein adsorption patterns and subsequent cellular responses. Figure 3 shows cell attachment on Ti, SUS316L, and cobalt-chromium-molybdenum (CCM) alloy after 6 h cultivation. The cell shape and spreading area strongly depend on material surface. This observation means protein layers formed on various material surfaces totally differ in cell attachment.

3.1. Titanium

The high biocompatibility of titanium is widely recognized, and it is extensively used for medical implants [18]. The oxide layer with a thickness of 3–7 nm on titanium is highly hydroxylated [53]. This layer is mainly TiO2 and forms stable hydration layers in water [54]. The TiO2 oxide layer easily adsorbs calcium ions, phosphate ions, and hydrogen ions in the surrounding environment [55,56]. In particular, divalent cations such as calcium ions have bridging effects between proteins and titanium surfaces [57]. Barberi et al. evaluated the adsorption of albumin and fibronectin on titanium [58]. They revealed that hydroxy group density and their acidic properties play important roles in controlling the secondary structure of proteins. Also, Duderija et al. evaluated protein adsorption on titanium with electrochemically controlled surface potential, and they concluded that electrostatic interactions are the dominant factor in protein adsorption amounts [65]. According to first-principles molecular dynamics (MD), the crystal phase of TiO2 determines the surface acid-base equilibrium, charge density, and PZC [59]. BSA adsorption on TiO2 strongly depends on pH conditions, which influence the coverage and orientation of proteins [60]. Huang et al. reported that TiO2 films with high rutile content increased FXII activation, fibrin formation, and platelet adhesion, suggesting the contribution of permittivity [43]. Tanaka et al. concluded that the relative permittivity of oxide layers affects thrombosis formation through protein adsorption, with different behaviors observed between Cr2O3-based and TiO2-based systems [66].

3.2. Stainless Steel

Stainless steel is used for stents and orthopedic components due to its excellent mechanical strength, chemical stability, and relatively low cost. Its corrosion resistance is mainly maintained by passive films composed of Fe2O3 and Cr2O3 with a thickness of 1–3 nm [61,62]. Iron and chromium are transition elements, and these elements easily form complexes with protein molecules [67]. This complex formation induces pH changes and leads to local corrosion [68]. Recent studies have revealed mechanisms by which chromium-rich films and protein complexes promote metal ion release. Atapour et al. reported that chromium concentrates in the oxide layers in the presence of albumin, and chromium release is promoted in whey protein solution [69]. Similar phenomena have been confirmed by Varmaziar et al., who pointed out that proteins with complex-forming ability can reconstruct oxide layers and induce local corrosion [70]. Furthermore, Du et al. revealed that albumin adsorption layers take chromium atoms from the crystal lattice, resulting in increased point defect density [71]. These results highlight the critical importance of controlling protein adsorption on stainless steel. Surface design strategies are essential to prevent undesirable protein-metal interactions while maintaining appropriate cellular responses.

3.3. Co-Cr Alloy

Cobalt–chromium (Co-Cr) alloys are widely used materials in orthopedic and dental fields due to their excellent mechanical properties. The oxide layer of Co-Cr-Mo alloys preferentially releases cobalt in biological environments and forms hydrated oxide layers containing chromium and molybdenum [63]. In the initial oxidation stage, Co and Cr oxidize selectively, and the thickness and composition of the oxide layer change over time [64]. This oxide layer is considered to contribute to both biocompatibility and corrosion resistance. For albumin adsorption on Co-Cr-Mo alloy surfaces, hydrophobic interactions are the dominant mechanism, and side-on orientation with monolayer formation tends to be preferred [72]. Zhou et al. found that heat treatment of Co-Cr-Mo alloys causes changes in albumin adsorption on the surface [73]. This is thought to be due to changes in the polar component of surface free energy by heat treatment and competition between electrostatic and hydrophobic interactions. The heat-treated surfaces of Co-Cr-Mo alloys show changes polar component of SFE, which affects albumin adsorption.

4. Surface Modification Strategies for Controlling Protein Adsorption and Cellular Responses

The physicochemical properties and protein adsorption behaviors discussed in previous chapters determine cellular responses on metal biomaterials. To design materials with optimized biocompatibility and biofunctionality, surface modification strategies can be systematically categorized based on two key control parameters: effective ligand density (σ_eff) and effective mechanical response (E_eff) [74]. As illustrated in Figure 4, σ_eff represents the surface density of bioactive ligands that are accessible and properly oriented for cellular recognition. This encompasses both the quantity of adsorbed proteins and the quality of their presentation—specifically, whether key binding motifs such as RGD sequences are exposed and accessible to cell surface receptors. The spatial distribution of these ligands is crucial, as integrin clustering and focal adhesion formation require ligand spacing within approximately 60–70 nm [75,76]. E_eff represents the mechanical properties that cells sense at the material interface. This is not simply the bulk mechanical properties of the substrate, but rather the effective stiffness and viscoelasticity of the integrated system comprising the oxide layer and adsorbed protein layers. These properties influence mechanotransduction pathways that control cell proliferation, differentiation, and function [77,78,79]. Based on this framework, surface modification strategies can be classified into three approaches: (1) direct control of σ_eff, (2) direct control of E_eff, and (3) coupled control of both parameters. This classification provides a rational basis for selecting and designing surface treatments to achieve specific cellular responses.

4.1. Design Principles for σ_eff and E_eff

4.1.1. Geometirical Design of σ_eff

Recent nano-biopatterning studies show that immobilizing adhesive ligands on ~100 nm-diameter nanodisc array accelerates cell spreading and focal adhesion formation [76]. In Nour et al., integrin-binding peptides were grafted as side chains on a polymer and the local nanocluster valency was tuned [80]. At comparable global surface coverage, increasing the local ligand clustering (e.g., ≥3–4 ligands/cluster) enhanced early cell adhesion and promoted myoblast activation and myotube formation—indicating that cellular responses are governed by local ligand density (cluster geometry) rather than total amount. Moreover, cells confined to defined ECM areas (~2–9 µm2), cells switch to a high-adhesion state per unit area: adhesion stress (force per ECM area) increases by roughly 4–15 times high compared with larger patterns (≥~35–60 µm2) [81]. These are consistent with evidence that local, rather than global, ligand density governs adhesion strength and focal adhesion formation, because clustering within ≤60–70 nm nearest-neighbor spacing enables multivalent engagement and rapid rebinding [82]. Thus, σ_eff is not fixed by total ligand amount but can be rationally controlled via local ligand density, i.e., cluster geometry such as island size and inter-island spacing.

4.1.2. Physicochemical Design of E_eff

Protein adlayers formed on metallic oxides exhibit thickness and viscoelasticity that depend on surface chemistry and hydration; these interfacial properties define E_eff. Under otherwise identical ionic conditions, adlayers that are thicker and more dissipative tend to delay early spreading and focal adhesion formation, whereas excessively rigid, low-dissipation layers can hinder subsequent rearrangement and maturation [83]. The amount of protein adsorption—approximated by the effective thickness—increases with the surface density of free amines; amine-rich coatings favor serum protein binding and thereby strengthen early cell adhesion [84]. Collectively, these observations indicate that the protein adlayer behaves as a dissipative interphase whose properties shifts the interfacial mechanical response (E_eff) and modulates early cell adhesion behavior. E_eff depends not only on adlayer thickness but also on the molecular architecture of the interphase and the structuring. Interfaces that co-tune hydration and dissipation—e.g., by increasing chain stretching, segmental friction, and interfacial water retention—balancing early receptor access with subsequent focal adhesion maturation [85]. E_eff is conceptually aligned with hard/soft protein-corona framework established for nanoparticles, in which an inner tightly bound “hard” layer and an outer, exchangeable “soft” layer jointly determine the apparent interfacial mechanics sensed by cells [86].

4.2. Implementation Strategies to Raise σ_eff

Direct immobilization of extracellular matrix (ECM) proteins (such as fibronectin and type I collagen) or cell adhesion peptides, such as RGD (Arg-Gly-Asp) and PHSRN (which is fibronectin derived sequences Ac-Pro-His-Ser-Arg-Asn-NH2) represents a fundamental strategy for directly controlling σ_eff [87]. Various implementation strategies for controlling effective ligand density are summarized in Table 3. For achieving both material specificity and stable immobilization, chimeric molecules that use solid-binding peptides as anchors and incorporate sequences like RGD are effective [88]. For example, chimeric peptides fusing titanium-binding peptides (e.g., RKLPDA, written in the one-letter amino acid code [89]) with RGD have achieved both specific binding to titanium surfaces and cell adhesion capability [90]. Similar approaches have been reported for stainless steel-binding peptides [91] and CCM-binding peptides [92,93], enabling selective modification according to material type.
Self-assembled monolayers (SAMs) provide another versatile platform for σ_eff control [94]. Alkanethiol SAMs on gold surfaces and silane-based SAMs on oxide surfaces allow precise control of ligand density and spacing through mixed monolayer approaches [95,96]. Click chemistry has also enabled efficient and selective peptide conjugation to functionalized surfaces [97,98,99].
Ionic surface modification—encompassing both ion implantation and ion exchange/doping—offers a composition-level route to tune the native oxide and thereby control the σ_eff. For instance, oxygen-ion implantation into TiO2 increases the attraction of the fibronectin PHSRN domain, improving cell adhesion and differentiation [100]. This is direct evidence of motif-specific adsorption and hence higher σ_eff. In parallel, silver ion exchange routes have been reported to increase the total protein adsorption capacity on metal surfaces by shifting wettability, polar component, and charge state [101]. Calcium-enriched surfaces further reinforce provisional matrix formation by accelerating adsorption–desorption dynamics of plasma proteins at early times [102,103]. Therefore, ionic modification rewrites coordination and hydration chemistry at the outer oxide, which modulates adsorption amount, orientation, and motif accessibility without changing the bulk substrate.
Table 3. Implementation strategies for controlling effective ligand density (σ_eff) on metallic biomaterial surfaces. ↑: increase relative to control.
Table 3. Implementation strategies for controlling effective ligand density (σ_eff) on metallic biomaterial surfaces. ↑: increase relative to control.
StrategyMaterialAnchorPresented MotifWithin-Study OutcomesRefs
Solid binding peptideTiminTBP-1
(RKLPDA)		minTBP-1-RGDMC3T3-E1: adhesion ↑, spreading ↑; some studies report osteogenic markers ↑[89,90]
SUS316LSUS-binding peptide (SBP)SBP-RGDHUVEC: adhesion/retention ↑[91]
Co-Cr-MoCCM-binding peptide (CBP)CBP-RGDHUVEC: adhesion ↑; MC3T3-E1: proliferation ↑, osteogenesis markers ↑[92,93]
Self-assembled monolayer (SAM)AuAlkanetiol SAMRGD
GRGDS
GRGDS + PHSRN		Baby hamster kidney: adhesion ↑
3T3 Swiss fibroblasts: cytoskeleton, focal adhesion ↑
NIH3T3: spreading, focal adhesion ↑		[96,97]
Au
Fe-MPN		Click-SAMRGDSP
cRGD		hMSC: adhesion, spreading, focal adhesion ↑
NIH3T3: adhesion, proliferation, migration ↑		[98,99]
Ionic
modification		TiO2O ion beam [100]
bioactive glassAg ion exchangeAlbumin
Fibronectin		Human gingval fibroblasts: adhesion, differentiation ↑[101]
TiCa ion coatingSerum proteinPrimary human alveolar osteoblasts: adhesion (ND), osteocalcin, procollagen type I ↑[102]

4.3. Implementation Strategies to Tune E_eff

E_eff represents the effective mechanical response that cells experience at the interface. This encompasses not only bulk elasticity but also the viscoelasticity of ECM and coating layers, which is a critical parameter for cellular mechanotransduction. Cells integrate E_eff through integrins, focal adhesions, and the cytoskeleton to regulate adhesion, migration, and fate. In our framework, E_eff complements σ_eff: identical ligand densities can yield distinct outcomes under different E_eff, while viscoelastic changes can mask or amplify ligand presentation. Table 4 presents various strategies for tuning E_eff at biomaterial interfaces.
On titanium, a quartz crystal microbalance with dissipation (QCM-D) resolved dissipation (ΔD) correlated positively with fibroblast spreading and negatively with motility on tannic-acid nanocoatings, indicating that the interfacial viscoelasticity (E_eff) directly modulates early adhesion dynamics [83]. Hydroxyapatite (HAp) coating is also a representative method for E_eff control. ECM layers formed on HAp-coated surfaces are rich in viscoelasticity and contribute to enhanced osteoblast differentiation. Tagaya et al. used QCM-D analysis of FBS adsorption and revealed that carbonate ions inhibit FBS adsorption on HAp-modified electrodes and control the viscoelasticity of adsorbed layers [104], demonstrating that solution ionic composition provides a practical handle to adjust interfacial viscoelasticity.
Surface polyethylene glycol (PEG) modification is also effective for E_eff control [105,106]. On titanium surfaces modified with RGD-conjugated PEG, osteoblast differentiation is enhanced, with longer-chain PEG showing greater effects [107,108]. Similar results have also been reported using zwitterionic polymers which enhanced osteogenesis [109]. These effects are consistent with the dynamic viscoelasticity of PEG or zwitterionic polymer adlayers influencing cellular responses. Additionally, reduction in non-specific adsorption through zwitterionic polymers or PEG modification enables selective protein adsorption and E_eff control [110,111]. Functional polymer modifications can thus function as dynamic interfaces to precisely control protein adsorption and cell adhesion through their involvement in E_eff.
Table 4. Implementation strategies for controlling effective mechanical response (E_eff) at metallic biomaterial interfaces. ↑: increase relative to control; ↓: decrease relative to control.
Table 4. Implementation strategies for controlling effective mechanical response (E_eff) at metallic biomaterial interfaces. ↑: increase relative to control; ↓: decrease relative to control.
StrategyMaterialCoatingAdlayer MetricValue
(vs. Bare)		Within-Study OutcomesRefs
Polyphenol filmTiTannic acidΔD/thickness of FBS~2×/1.3–1.7×Fibroblasts: adhesion ↑[83]
Inorganic layerAuHydroxy apatiteΔD/thickness of FBS~2×/2–4×Viscoelasticity of FBS ↑
Hepatocyte: morphological change		[104,112]
Tether rigidityAucRGD + spacerSpacer typePolyproline vs. aminohexanoic acidPolyproline spacer: adhesion, focal adhesion ↑[113]
Hydrated brushTiPEGthickness of PEG-RGD1.5–2× in airMC3T3-E1: differentiation↑
Osteoblast: osteogenesis in vivo ↑		[107,108]
GlassPEGPEG lengthLong/short = 318 nm/9.5 nmLonger PEG: focal adhesion, spreading ↓[114]
TiZwitter ionic polymerprotein adsorption0.1–0.3×MC3T3-E1: differentiation ↑[109]

4.4. Coupled Control of σ_eff and E_eff

Micro- and nanometer-scale surface roughness and topography can control both σ_eff [58,115,116] and E_eff [117]. Appropriate roughening of titanium surfaces promotes preferential adhesion of osteoblasts [118,119,120], whereas certain submicron regimes have been reported to impair cellular function [121,122,123,124]. Together, these observations indicate that surface topography co-regulates σ_eff and E_eff simultaneously, via changes in protein coverage/orientation (σ_eff) and in the viscoelasticity of the adsorbed layer (E_eff). Hierarchical surface architectures further enable multiscale cellular response control and have emerged as a practical paradigm for coordinated regulation of cell proliferation, differentiation, and functional expression [125,126,127].
TiO2 nanotubes provide compact geometric knobs—diameter and length—that rewrite σ_eff while concurrently shifting E_eff through changes in interfacial hydration, oxide phase, and thin-wall mechanics [128]. Sub-100 nm tubes favor early attachment and focal adhesion maturation; larger diameters bias differentiation and mineralization [129,130]. The nanotube topography can tune σ_eff and E_eff by modulating protein adsorption and orientation, and the viscoelasticity of the interfacial protein corona. Consistently, XB130, an actin filament-associated adaptor, has been shown to be involved in TiO2 nanotubes-induced osteogenic differentiation via mechano-biochemical signaling [131]. Lateral spacing of nanotubes could also influenced osteogenesis: in an in vivo model, nanotube arrays with lateral spacing of 92 nm exhibits greater osteogenic activity in vivo model than those with 25 nm, plausibly via the FAK/RhoA/YAP cascade [132]. Moreover, the pore spaces and gaps between neighboring nanotubes can induce protein aggregation at the top rim of TiO2 nanotubes [133], indicating that lateral spacing is a key geometric parameter that modulate mechanotransduction through σ_eff and E_eff. The same array thus co-tunes ligand presentation and the interfacial viscoelastic milieu without altering the bulk substrate [117].
Micro-arc oxidation (MAO) builds a ceramic, microporous TiO2 layer whose chemistry (Ca/P, Nb, Ag, etc.) and porosity/architecture can be scripted [134,135]. The polar/ionic dopants reshape σ_eff [136,137], while the porous, water-rich ceramic elevates E_eff, offering coupled control on osteogenic interfaces.

5. Perspective

This review has comprehensively discussed the mechanisms of protein adsorption and cell adhesion on metallic biomaterial surfaces, as well as surface modification technologies for controlling these interactions. The molecular mechanisms by which oxide layer properties formed on metal surfaces initiate protein adsorption and subsequently trigger cell adhesion have become increasingly clear. However, realizing next-generation metallic biomaterials requires addressing new challenges beyond conventional understanding focused on planar surfaces.

5.1. Three-Dimensional Modification of Porous Metal Internal Walls: Development of Material-Inspired Biomaterials

Most current surface modification technologies target planar or external surfaces, but three-dimensional (3D) interfacial interactions play crucial roles in actual biological environments. The development of internal wall modification technologies for porous metal materials is extremely important for advancing metamaterial-inspired biomaterials [138,139]. Within porous structures, inadequate fluid exchange often leads to air bubble retention, making uniform modification difficult with conventional liquid-phase methods [140]. Solving this requires developing new technologies such as selective modification using molecular self-assembly [141]. Precise control of the chemical and physical properties of porous internal walls is expected to enable advanced functions difficult to achieve with conventional surface modifications, including directional cell migration, guided angiogenesis, and spatiotemporal control of drug release [142,143,144].

5.2. Bridging Hierarchical Understanding Gaps: From Molecular to Tissue Level

A fundamental challenge in metallic biomaterial research is the understanding gap between different hierarchical levels. While physicochemical principles can explain surface-protein interactions well [145,146], this understanding breaks down when transitioning to cellular responses [147,148]. The relationship becomes even more complex at the tissue level, where individual cell behavior does not necessarily predict tissue compatibility [149,150]. The transition from protein adsorption to cellular response involves multiple uncertainties [151]. Protein layers undergo dynamic reorganization after cell contact, with cellular ECM production potentially overwriting initial adsorption patterns [152]. Cell membrane receptors integrate multiple simultaneous interactions rather than responding to single ligand-receptor pairs, making it difficult to predict cellular responses from surface-adsorbed protein characteristics alone [153,154].
Addressing these hierarchical gaps requires new research strategies beyond traditional reductionist approaches [155]. Systems biology approaches integrating omics analyses, real-time in vivo imaging with spatiotemporal resolution, mathematical models linking different scales, and reverse design approaches starting from desired tissue responses are essential for achieving true rational design of biomaterials [156,157,158].

5.3. Sustainability-Oriented Design of Metallic Biomaterials

Sustainability has emerged as an essential design consideration for metallic biomaterials, complementing traditional requirements of functionality and biocompatibility [159]. Despite the increasing importance of resource efficiency and waste reduction in medical devices, the recycling and lifecycle management of metallic implants have received limited attention [160]. Design for disassembly principles enable efficient material recovery and reuse at end-of-life, while minimizing waste throughout the implant lifecycle [161]. Advanced manufacturing approaches, particularly 3D printing technology, allow precise customization with minimal material waste [162,163]. Additionally, biodegradable magnesium alloys offer inherent sustainability advantages by eliminating revision surgery requirements and long-term foreign body concerns [164]. Surface modification strategies for Mg-based materials further contribute to sustainable biomaterial development by enabling controlled degradation profiles [165]. These sustainability-oriented approaches will become increasingly critical as healthcare systems seek to balance clinical performance with environmental responsibility.

6. Conclusions

This review has examined the molecular mechanisms governing protein adsorption and cell adhesion on metallic biomaterial surfaces. The physicochemical properties of oxide layers—including Lewis acid-base chemistry, surface charge, surface free energy, and permittivity—collectively determine protein adsorption behavior and subsequent cellular responses. Surface modification strategies can be systematically categorized based on control of effective ligand density (σ_eff) and effective mechanical response (E_eff). The most promising approaches involve coupled control of both parameters through hierarchical surface architectures and three-dimensional modifications. Despite advances in understanding molecular-level interactions, substantial challenges remain in bridging the gap between surface chemistry and biological performance. Future developments must address three-dimensional interfacial interactions and develop systems-level approaches that integrate multiple scales of biological organization to enable rational design of next-generation metallic biomaterials.

Author Contributions

Conceptualization, writing—original draft preparation, and editing, S.M.; writing—review and editing, M.S. All authors have read and agreed to the published version of the manuscript.

Funding

A part of this work was supported by JSPS KAKENHI Grant Number JP25K15886.

Data Availability Statement

The data that support the findings of this study are available from the corresponding author upon reasonable request.

Conflicts of Interest

The authors declare no conflicts of interest.

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Figure 1. Schematic illustration of water interaction with a titanium dioxide. (left) A water molecule with two lone pairs of electrons on the oxygen atom (black dots). (center) Formation of a hydrogen-bond bridge between the surface Ti–O groups and the water molecule; δ+ and δ− indicate the resulting partial positive and negative charges. (right) Dissociative adsorption of water, yielding a surface hydroxyl group and a proton bound to neighboring oxygen.
Figure 1. Schematic illustration of water interaction with a titanium dioxide. (left) A water molecule with two lone pairs of electrons on the oxygen atom (black dots). (center) Formation of a hydrogen-bond bridge between the surface Ti–O groups and the water molecule; δ+ and δ− indicate the resulting partial positive and negative charges. (right) Dissociative adsorption of water, yielding a surface hydroxyl group and a proton bound to neighboring oxygen.
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Figure 2. Schematic relationship between pH and the apparent surface charge of a metal-oxide surface based on the site-binding model. The curve represents the apparent surface charge as a function of pH, becoming positive at low pH, zero at the point of zero charge (PZC), and negative at high pH. The cartoons above the curve illustrate representative surface species in each pH region: protonated hydroxyl groups (≡MOH2+) under acidic conditions, neutral hydroxyl groups (≡MOH) around the PZC, and deprotonated groups (≡MO−) under basic conditions, generated by the acid–base equilibria. Illustration of site-binding view of the point of zero charge (PZC).
Figure 2. Schematic relationship between pH and the apparent surface charge of a metal-oxide surface based on the site-binding model. The curve represents the apparent surface charge as a function of pH, becoming positive at low pH, zero at the point of zero charge (PZC), and negative at high pH. The cartoons above the curve illustrate representative surface species in each pH region: protonated hydroxyl groups (≡MOH2+) under acidic conditions, neutral hydroxyl groups (≡MOH) around the PZC, and deprotonated groups (≡MO−) under basic conditions, generated by the acid–base equilibria. Illustration of site-binding view of the point of zero charge (PZC).
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Figure 3. Osteoblast attachment on Ti, SUS316L, and Co-Cr-Mo alloy surfaces after 6 h cultivation. Green fluorescence shows cell morphology and spreading behavior on different metal surfaces. Scale bar = 50 µm.
Figure 3. Osteoblast attachment on Ti, SUS316L, and Co-Cr-Mo alloy surfaces after 6 h cultivation. Green fluorescence shows cell morphology and spreading behavior on different metal surfaces. Scale bar = 50 µm.
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Figure 4. Schematic representation of the two key control parameters for surface modification: effective ligand density (σ_eff) and effective mechanical response (E_eff). The diagram shows how oxide layer properties govern protein adsorption, which mediates cellular responses through coupled control of both parameters.
Figure 4. Schematic representation of the two key control parameters for surface modification: effective ligand density (σ_eff) and effective mechanical response (E_eff). The diagram shows how oxide layer properties govern protein adsorption, which mediates cellular responses through coupled control of both parameters.
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Table 1. Representative relative permittivity (εr) and points of zero charge (PZC) of biomaterial-relevant metal surfaces.
Table 1. Representative relative permittivity (εr) and points of zero charge (PZC) of biomaterial-relevant metal surfaces.
MaterialsεrPZCReferences
SiO23.9–4.62.0–4.6[29,44,45,46]
Al2O39.0–11.56.8–9.9[29,44,46,47]
Cr2O311.9–13.36.0[29,45,46]
Fe2O312–167.1–9.2[29,45,46]
TiO2 (anatase)40–555.5–6.5[29,42]
TiO2 (rutile)80–1703.6–4.4[29,42,45,46]
H2O79.6–[46]
Table 2. Surface characteristics and interfacial features of titanium, SUS316L stainless steel, and CCM.
Table 2. Surface characteristics and interfacial features of titanium, SUS316L stainless steel, and CCM.
MaterialsThickness (nm)Oxide/Chemical CompositionInterfacial FeaturesReferences
Ti3–7Mainly TiO2 (anatase/rutile)Stable hydration layer in water; readily adsorbs Ca2+, PO43−, H+; hydroxyl density/acidity tunes protein secondary structure and motif exposure; crystal phase controls acid–base equilibrium/PZC; pH controls BSA coverage/orientation[43,53,54,55,56,57,58,59,60]
SUS316L1–3Cr2O3-rich passive film with Fe(III) oxidesOutermost monolayers are water- and hydroxyl-rich[61,62]
Co-Cr-Mo2.5–3Cr(III) oxide (Cr2O3) with a smaller amount of Cr(III) hydroxide (Cr(OH)3)Surface contains a large amount of OH− (hydrated/oxyhydroxide character)[63,64]
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Migita, S.; Sato, M. Protein Adsorption and Cell Adhesion on Metallic Biomaterial Surfaces. Adhesives 2025, 1, 15. https://doi.org/10.3390/adhesives1040015

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Migita S, Sato M. Protein Adsorption and Cell Adhesion on Metallic Biomaterial Surfaces. Adhesives. 2025; 1(4):15. https://doi.org/10.3390/adhesives1040015

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Migita, Satoshi, and Masaki Sato. 2025. "Protein Adsorption and Cell Adhesion on Metallic Biomaterial Surfaces" Adhesives 1, no. 4: 15. https://doi.org/10.3390/adhesives1040015

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Migita, S., & Sato, M. (2025). Protein Adsorption and Cell Adhesion on Metallic Biomaterial Surfaces. Adhesives, 1(4), 15. https://doi.org/10.3390/adhesives1040015

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