Artificial ligaments in anterior cruciate ligament reconstruction: Coating strategies for PET-based materials

Abstract

Artificial ligaments, as an important implant for Anterior Cruciate Ligament reconstruction (ACLR), offer notable advantages in early return to sport. However, most artificial ligaments currently used in clinical ACLR are made of polyethylene terephthalate (PET), a polymer characterized by a smooth and hydrophobic surface that limits cell adhesion and tissue growth, leading to the formation of fibrous scar tissue at the tendon-bone interface. To address these limitations, various surface coating strategies have been developed, including biocompatible, tissue inductive, osteoconductive, drug delivery, and immunomodulatory coatings. These approaches improve biological performance, promote ligamentization, and enhance integration with host tissues. Additionally, the application of composite functional coatings and smart responsive coatings offers new directions for future research. Despite promising preclinical results, most studies remain at the animal experiment stage, and the underlying mechanisms need further investigation. This review summarizes recent advances in coating strategies for artificial ligaments, highlighting their functional classification, technical development, and potential for clinical translation.

Translational potential statement

To address key challenges in the application of artificial ligaments, such as limited biocompatibility and poor tissue integration, it is essential to understand the current research progress. This review provides a comprehensive overview of the coatings used for ligaments, highlighting the promising role of surface modification in enhancing implant performance. It offers valuable insights for improving the clinical success rate of artificial ligaments and their long term effectiveness in ACLR, thus holding significant clinical translational potential.

Graphical abstract

To address key challenges in the application of artificial ligaments, such as limited biocompatibility and poor tissue integration, it is essential to understand the current research progress. This review provides a comprehensive overview of the coatings used for ligaments, highlighting the promising role of surface modification in enhancing implant performance. It offers valuable insights for improving the clinical success rate of artificial ligaments and their long term effectiveness in ACLR, thus holding significant clinical translational potential.

1. Introduction

The anterior cruciate ligament (ACL) comprises the anteromedial and posterolateral bundles, both of which are crucial structures for preventing tibial anterior translation and maintaining knee stability. ACL injures can lead to anterior-posterior and rotational instability, significantly impairing athletic performance. ACL injuries are among the most common knee pathologies encountered in sports [1], especially in activities that require frequent changes in direction and landing. According to statistics, the incidence of ACL injuries in adolescent athletes is 6.9 %, with a rate of 6.0 % in males and 8.4 % in females [2]. In the general population across all age groups, the incidence is 686 cases per million person-years, with higher rates observed in males [3]. These injuries are characterized by long recovery periods and significant economic burden. Non-surgical treatment may expose patients to the risk of persistent or recurrent instability, as well as secondary meniscal and/or cartilage damage [4]. Therefore, for patients with complete ligament rupture and knee instability, the primary treatment is arthroscopic anterior cruciate ligament reconstruction (ACLR), especially for younger individuals [5,6]. Over 400000 ACLR surgeries are performed annually worldwide [7], with approximately 80 % of patients successfully returning to sports [8]. Current guidelines recommend ACLR for active young individuals (ages 18–35) with ACL tears [9].

The prognosis of ACLR largely depends on graft healing, which includes bone integration at the tunnel interface and intra-articular (IA) graft remodeling, known as “ligamentization” [10]. Graft selection plays a crucial role in the healing process. Currently, the graft options for ACLR include autografts, allografts, and artificial ligaments [11], each with distinct advantages and limitations. Autografts, particularly bone-patellar tendon-bone (BPTB) and hamstring tendons, are considered the gold standard by many clinicians and researchers due to their reliable clinical outcomes [[12], [13], [14]]. BPTB grafts, in particular, offer superior bone-to-bone healing at both the femoral and tibial tunnel interfaces. However, both autografts and allografts undergo the “ligamentization” following implantation. During this process, the biomechanical properties significantly decrease, and the risk of re-injury under external forces is higher [15]. In contrast, artificial ligaments can quickly achieve optimal mechanical strength and restore knee stability in the early postoperative period, providing an advantage for early return to sports [16]. However, artificial ligament materials generally exhibit poor bioactivity and are unable to support biological regeneration and ligamentization after implantation [17]. Although a variety of artificial ligaments have been developed for ACLR in recent decades, most were ultimately abandoned due to early mechanical failure, which leads to synovitis and osteoarthritis [18,19]. However, significant progress has been made in the research and application of artificial ligaments in recent years. The introduction of Ligament Augmentation and Reconstruction System (LARS) provides a high-performance alternative to autografts and allografts for ligament reconstruction. As a new generation of synthetic ligament, LARS mimics the natural ACL, consisting of both a bone tunnel segment and an intra-articular segment, and is pre-treated with external rotation [20]. However, LARS is made from PET fibers, which are smooth and hydrophobic. Therefore, despite its good mechanical properties, LARS faces challenges in terms of biological activity and integration with host tissue. Growing evidence suggests that the poor biological activity of artificial ligaments may be a cause of implant failure [21,22]. Overall, no single type of graft can be universally regarded as ideal, as each graft has inherent strengths and limitations. Therefore, graft selection should be personalized based on individual patient characteristics, such as age, activity level, anatomical factors, rehabilitation plan, and recovery expectations, rather than relying on a one-size-fits-all standard.

To enhance the performance of artificial ligaments and address challenges related to their biological activity and biocompatibility, researchers have explored various strategies and developed numerous materials and techniques. Carbon nanotubes (CNTs), which are cylindrical structures formed by rolling graphene, possess excellent mechanical properties with high strength and toughness, as well as good conductivity. These attributes make CNTs highly promising for composite materials. Recently, Wang et al. [23] utilized CNTs as ligament materials, demonstrating their capacity to withstand one million cycles of continuous bending and complex deformation, with successful outcomes in both rabbit and sheep ACLR models. Interestingly, Lu et al. [24] reinforced silk fibers by feeding silkworms with mulberry leaves sprayed with CNT dispersions, offering a promising approach to incorporating naturally high-strength fibers into artificial ligaments. However, a study has reported that CNTs can induce inflammatory responses [25], underscoring the need for thorough characterization and application specific evaluation of CNTs in biomedical and materials science research. In addition, several reviews have discussed strategies for improving artificial ligaments [[26], [27], [28], [29]]. For example, Li et al. [26] summarized strategies to enhance the biological activity of artificial ligaments, dividing them into scaffold based and cell based categories. Yao et al. [29] explored the application of biological modulation in ACLR, highlighting how the timing, location, and method of biological modulation influence mechanical performance. However, there is currently no systematic review dedicated to the surface modification of artificial ligaments via coating strategies. Therefore, this review focuses on the latest research progress in coating modifications of artificial ligaments in ACLR, discussing animal models, coating techniques, and functional outcomes, providing theoretical support for future clinical translation.

2. ACL animal joint models

The selection of an appropriate animal joint model in artificial ligament research significantly influences experimental outcomes and should be closely aligned with the specific objectives of the study. Commonly used models include the intra-articular (IA) ACLR model (Fig. 1A), and the extra-articular (EA) tendon-bone healing (TBH) model (Fig. 1C) [30,31]. Among these, the EA TBH model involves creating an EA bone tunnel and filling it with a graft. This model involves a simpler surgical procedure with lower technical demands, minimal trauma, and more stable fixation. Because it avoids mechanical stress and places fewer demands on tunnel positioning and graft fixation, it often yields more consistent outcomes. It also avoids the influence of synovial fluid, which has been shown to negatively affect graft-bone integration post-implantation [30,32,33]. Additionally, bone formation typically occurs more rapidly in EA tunnels compared to IA ones [34]. However, this model does not replicate the biomechanical and biological environment of clinical ACLR. In contrast, the IA TBH model (Fig. 1B) more closely simulates the biological conditions of ACLR due to the presence of synovial fluid, though it is less commonly used [31]. The IA ACLR model simulates the clinical ACLR surgery by severing the natural ACL and then reconstructing it, involving tibial and femoral tunnels. However, it is technically more demanding, causes greater surgical trauma, and is associated with a higher risk of complications. Moreover, the graft is subjected to substantial mechanical loading, which may lead to variability in outcomes [35].

Fig. 1.

Diagram of ACL animal joint models. (A–C) The animal models mentioned in the study include: extra-articular TBH model, intra-articular TBH model, and intra-articular ACLR model. (D) Distribution of Joint Models in the Studies. (E) Distribution of Animal Species in the Studies.

As shown in Fig. 1 and Table 2, a total of 33 studies conducted in vivo animal experiments. Among these, 64 % (21/33) employed IA ACLR models, while 33 % (11/33) utilized EA TBH models. Only one study (1/33) used an IA TBH model. Regarding the animal species used, rabbits (21/33) were the most commonly used, followed by rats (6/33), pigs (2/33), beagle dogs (2/33), sheep (1/33), goats (1/33), and mice (1/33). This distribution highlights the frequent use of rabbits due to their appropriate size, cost-effectiveness, and well established experimental protocols. Overall, each animal model presents distinct advantages and limitations. Careful selection based on the experimental objective and desired clinical relevance is essential to ensure meaningful and reproducible outcomes.

Table 2.

Animal experimental studies on coated ligaments.

Authors	Coating materials	Coating type	Coating methods	Animal models	Joint model	Follow-up min to max	Site of action	Results	References
Yang et al.	HPC	Biocompatibility Oriented Coating	Dip coating	Rabbit	EA TBH	4–8 weeks (wks)	Graft-bone interface	Improved graft hydrophilicity and osseointegration	[76]
Yang et al.	HPC	Biocompatibility Oriented Coating	Spraying	Beagle dog	IA ACLR	4–16 wks	Graft-bone interface	Improved hydrophilicity, enhanced BMSCs survival rate, better growth status, and increased maximum failure load and stiffness. HPC/PET group vs. PET group (Failure load at 4 weeks: 502.3 ± 20.0 N vs. 182.7 ± 14.5 N; Failure load at 16 weeks: 325.0 ± 32.6 N vs. 142.0 ± 5.3 N; Stiffness at 16 weeks: 46.7 ± 7.8 N/mm vs. 27.3 ± 4.5 N/mm)	[77]
Cai et al.	SF	Biocompatibility Oriented Coating	Chemical Grafting Coating	Rabbit	IA ACLR	6–12 wks	Intra-articular Graft	Improved cell compatibility, reduced inflammation, and achieved better ligamentization	[81]
Kawai et al.	Chitin	Biocompatibility Oriented Coating	—	Rabbit	IA ACLR	8 wks	Graft-bone Interface, Intra-articular Graft	Knee joint inflammation was reduced, and TBH was enhanced. The peak resistance increased from 19.2 ± 15.3 N in the control group to 42.2 ± 12.7 N (the peak resistance of the intact ligament was 137.1 ± 18.1 N)	[84]
Li et al.	HA and CG	Biocompatibility Oriented Coating	LBL	Rabbit, Pig	IA ACLR	3–12 wks	Intra-articular Graft	The graft's hydrophilicity and cell compatibility were improved, enhancing vascular remodeling and collagen reorganization	[85]
Cho et al.	HA and CG	Biocompatibility Oriented Coating	LBL	Pig	IA ACLR	12 wks	Graft-bone interface	Promoted the healing of the grafted bone at the distal end of the bone tunnel	[42]
Balasubramaniyan et al.	HA and CG	Biocompatibility Oriented Coating	LBL	Rabbit	EA TBH	4 wks	Graft-bone interface	Cell compatibility was improved, enhancing tissue regeneration	[86]
Li et al.	Fibronectin	Biocompatibility Oriented Coating	Ultrasonic Spray Coating	rat	IA ACLR	4 wks	Intra-articular Graft	Improved cell compatibility and promoted tissue ingrowth into graft fibers	[52]
Shi et al.	Polyallylamine hydrochloride, chondroitin sulfate, magnesium silicate	Biocompatibility Oriented Coating	LBL	Rabbit	IA ACLR	12 wks	Graft-bone interface	Promoted BMSC adhesion and proliferation, induced osteogenic differentiation, and enhanced graft-bone integration. The maximum failure load (104.3 ± 26.9 MPa) and stiffness (35.7 ± 10.3 N/mm) were superior to the PET group	[90]
Wang et al.	Graphene	Tissue Inductive Coating	Chemical VaporDeposition	Rabbit	EA TBH	4–12 wks	Graft-bone interface	Cell compatibility was improved, promoting osteogenic differentiation, stimulating new bone formation, and enhancing graft-bone integration	[91]
Wang et al.	CNTs	Tissue Inductive Coating	Wrap	Sprague–Dawley rat	IA ACLR	1–2 wks	Graft-bone interface	The failure load and stiffness were similar to those of natural ligaments, with bone tunnel narrowing. The pull-out forces at the 1st and 2 nd weeks were 14.29 ± 1.06 N and 20.85 ± 3.91 N, respectively, approximately 1.4 times that of bare helical PET fibers	[92]
Jiang et al.	SF and HAp	Tissue Inductive Coating	Dip coating	Rabbit	EA TBH	4–8 wks	Graft-bone interface	Cell compatibility was improved, osteogenic activity was enhanced, and new bone formation at the graft-bone interface was promoted. At 8 weeks post-surgery, the failure load (91.00 ± 5.39 N vs. 56.60 ± 5.37 N) and stiffness (8.10 ± 1.78 N/mm vs. 5.20 ± 0.84 N/mm) were both higher than those in the PET group	[50]
Li et al.	Cu-BG	Tissue Inductive Coating	Pulsed laser deposition	Goat	IA ACLR	12 wks	Graft-bone interface	The graft's hydrophilicity was improved, cell compatibility was enhanced, and osteogenic and angiogenic capabilities were strengthened	[96]
Vaquette et al.	PolyNaSS	Tissue Inductive Coating	Grafting	Sheep	IA ACLR	3–12 months	Graft-bone Interface, Intra-articular Graft	Promoted osteogenesis, with graft failure loads at 3 months (188 ± 52 N vs. 144 ± 69 N) and 12 months (322 ± 170 N vs. 260 ± 126 N) both higher than those in the PET group	[44]
Wang et al.	HAp	Osteoconductive Coating	Plasma spraying	Rabbit	IA ACLR	4–12 wks	Graft-bone interface	The graft's hydrophilicity and cell compatibility were improved, enhancing graft-bone healing. At 12 weeks, the failure load (72.30 ± 3.80 N vs. 52.60 ± 5.80 N) and stiffness (29.72 ± 2.14 N/mm vs. 20.18 ± 2.04 N/mm) were significantly higher than those in the PET group	[38]
Li et al.	HAp	Osteoconductive Coating	Dip coating	Rabbit	EA TBH	4–8 wks	Graft-bone interface	Enhanced graft-bone healing. At 8 weeks, the average failure load (90.7 ± 13.5 N vs. 60.5 ± 9.1 N) and stiffness (9.3 ± 2.0 N/mm vs. 5.7 ± 1.2 N/mm) were higher than those in the control group	[39]
Cai et al.	HAp	Osteoconductive Coating	Biomineralization	Rabbit	IA ACLR	12 wks	Graft-bone interface	Improved cell compatibility, promoted osteogenesis, and enhanced graft-bone healing. At 12 weeks, the failure load (76.4 ± 6.3 N vs. 52.1 ± 6.3 N) and stiffness (14.9 ± 3.1 N/mm vs. 10.1 ± 1.6 N/mm) were higher than those in the PET group	[97]
Cai et al.	CaP	Osteoconductive Coating	Biomineralization and electrodeposition	Rabbit	IA ACLR	6–12 wks	Graft-bone interface	Cell compatibility was improved, promoting osteogenesis. At 12 weeks post-surgery, the ultimate failure load and stiffness showed an increasing trend in the control group, biomineralization group, and electro-deposition group	[37]
Ma et al.	Strontium-substituted hydroxyapatite	Osteoconductive Coating	Dip coating	Rabbit	IA ACLR	4–12 wks	Graft-bone interface	The graft's hydrophilicity was improved, promoting osteogenesis. At 12 weeks, the maximum failure load (121.5 ± 25.2 N vs. 48.3 ± 16.1 N) was significantly higher than that in the PET group	[40]
Egawa et al.	SrSiP	Osteoconductive Coating	Dip coating	Rabbit	EA TBH	8 wks	Graft-bone interface	Promoted osteogenesis and enhanced new bone formation	[100]
Li et al.	BG	Osteoconductive Coating	Dip coating	Rabbit	EA TBH	3–12 wks	Graft-bone interface	Promoted osteogenesis and angiogenesis, enhancing graft-bone integration. At 6 weeks post-surgery (92 ± 12.0 N vs. 58.2 ± 9.4 N) and 12 weeks post-surgery (114 ± 16.0 N vs. 68 ± 9.1 N), the average failure load in the experimental group was higher than that in the control group	[101]
Wu et al.	BG	Osteoconductive Coating	Dip coating	Rabbit	IA TBH	6–12 wks	Graft-bone interface	Increased osteoblast activity and enhanced graft-bone integration. At 6 weeks and 12 weeks post-surgery, the average failure load of the grafts was higher than that of the control group	[31]
Zhang et al.	Collagen, simvastatin	Drug Delivery Coating	Dip coating	Rabbit	IA ACLR	4–8 wks	Graft-bone Interface, Intra-articular Graft	Improved cell compatibility, promoted osteogenesis, and enhanced graft-bone healing. At 8 weeks post-surgery, the average failure load (86.4 ± 4.1 N) and stiffness (26.1 ± 2.0 N/mm) were higher than those in the PET group	[104]
Han et al.	HA, CHI	Drug Delivery Coating	LBL	Rabbit	IA ACLR	4–8 wks	Graft-bone interface	Hydrophilicity was improved, enhancing BMSCs osteogenic activity and graft-bone healing. At 8 weeks post-surgery, the maximum failure load and stiffness of the grafts were increased by 87.1 % and 80.5 %, respectively, compared to the control group	[105]
Gao et al.	Multilayer RSF loaded with curcumin and Zn2+	Drug Delivery Coating	Dip coating	Sprague–Dawley rat	IA ACLR	1–8 wks	Graft-bone interface	The graft's hydrophilicity and cell compatibility were improved. Sequential release of curcumin and Zn2+ inhibited acute inflammation in the early stage and enhanced osteogenesis in the later stage	[106]
Wu et al.	Strontium calcium phosphate and magnesium ions	Drug Delivery Coating	Magnetron sputtering and hydrogel coating	Sprague–Dawley rat	IA ACLR	2–12 wks	Graft-bone interface	Improved cell compatibility and osteogenic activity. Sequential release of Mg2+ and Sr2+ promoted vascular and neural generation at the graft-bone interface, synergistically enhancing graft-bone integration	[107]
Ding et al.	Simvastatin, CS, HAp	Drug Delivery Coating	Plasma spraying and Dip coating	Sprague–Dawley rat	EA TBH	3–6 wks	Graft-bone interface	Sustained drug release for 25 days, promoting osteogenesis	[119]
Kang et al.	Polydopamine, bone morphogenetic protein-2	Drug Delivery Coating	Dip coating	Rabbit	EA TBH	8 wks	Graft-bone interface	Enhanced hydrophilicity and cytocompatibility; promoted osteogenesis. Failure load at 8 weeks (79.93 ± 6.49 N vs. 44.25 ± 4.01 N) significantly higher than PET group	[120]
Wang et al.	GelMA, PEGDA, and SCS	Immune Modulation Coating	Hydrogel coating	Mouse	IA ACLR	2–8 wks	Graft-bone interface	Hydrophilicity was improved, regulating macrophage polarization and enhancing graft-bone integration. It promoted osteogenesis and angiogenesis while inhibiting scar formation. At 8 weeks post-surgery, the pull-out strength was approximately 12 N, three times that of the control group	[111]
Cai et al.	SF, HAp	Composite Functional Coating	Chemical Grafting Coating	Beagle dog	IA ACLR	6–12 wks	Graft-bone Interface, Intra-articular Graft	Surface roughness and hydrophilicity were improved, promoting osteogenesis and enhancing graft-bone integration. It facilitated ligamentization of the graft. At 6 weeks post-surgery (290.9 ± 27.3 N vs. 169.8 ± 19.4 N) and 12 weeks post-surgery (493.2 ± 27.3 N vs. 198.3 ± 17.8 N), the failure load was higher than that in the control group, with the 12-week failure load approaching that of the natural ACL (516.4 ± 12.6 N)	[36]
Li et al.	Polydopamine, Chondroitin Sulfate, BMP-2	Composite Functional Coating	Chemical Grafting and physical fixation	Sprague–Dawley rat	EA TBH	3 days-12 wks	Graft-bone Interface	Regulated macrophage polarization, optimized BMSC status, promoted osteogenesis, and enhanced graft-bone healing	[112]
Li et al.	SF, Chondroitin Sulfate, Sodium Hyaluronate, HAp	Composite Functional Coating	Dip coating	Rabbit	IA ACLR	8–24 wks	Graft-bone interface	Enhanced graft-bone healing. At 16 weeks post-surgery (28.2 ± 3.7 N vs. 16.2 ± 1.5 N) and 24 weeks post-surgery (42.7 ± 4.3 N vs. 21.7 ± 2.5 N), the pull-out force in the experimental group was significantly higher than that in the control group	[118]
Gerasimenko et al.	CNTs, collagen type II	Composite Functional Coating	Spin coating	Rabbit	EA TBH	4–12 wks	Graft-bone interface	Improved hydrophilicity and hemocompatibility, promoting osteogenesis	[121]

3. Overview of ligament related coating methods

Coating technology is a simple and intuitive surface modification strategy, which can be classified into chemical and physical approaches. Some of these techniques have already been industrialized. Among them, chemical synthesis is a common and classic coating approach, which involves synthesizing the target coating by controlling the reaction conditions and material concentrations. Common techniques in this category include dip coating, sol–gel, and electrochemical deposition. Physical deposition methods are those that directly deposit materials onto the substrate using physical techniques. These techniques include spraying (vacuum spraying, plasma spraying), sputtering deposition (DC sputtering, RF sputtering, magnetron sputtering), and pulsed laser deposition (see Fig. 2).

Fig. 2.

Common coating methods for ligaments. In current research, the most commonly applied coating techniques for ligaments include dip coating, sol–gel, electrochemical deposition, plasma spraying, sputter deposition, PLD and biomineralization.

Specifically, coating methods such as dip coating, chemical deposition, and spraying have been widely reported [[36], [37], [38], [39], [40]]. It is important to note that most studies involve pre-treating the surface of the artificial ligament to improve coating adhesion. Firstly, dip coating, as a convenient and cost-effective coating technique, has been widely applied in the preparation of artificial ligament coatings [41]. Common organic coating materials, such as collagen, hyaluronic acid, Chitosan (CS), and silk fibroin (SF), are typically applied using the dip coating method [36,[42], [43], [44], [45]]. Additionally, layer-by-layer (LBL) assembly can be integrated into the dip coating process to enable differential degradation and release kinetics, allowing for the sequential delivery of functional agents from distinct inner and outer layers [46]. However, dip coating lacks stability and uniformity, and various dipping parameters, such as solvent type, deposition temperature, and polymer concentration, can significantly influence the coating performance [47].

Electrochemical deposition is a versatile technique that is independent of the substrate shape. By altering deposition parameters such as electrolyte concentration, pH, temperature, salt concentration, current, and deposition time, researchers can tailor surface morphology, film thickness, and chemical composition [48]. The most common electrochemical technique is electrochemical deposition, which can achieve uniformly thick coatings on substrates [41]. Inorganic coating materials, mainly silicon, calcium and phosphorus, are typically obtained by chemical deposition [37,49,50]. However, chemical deposition generally involves dispersing inorganic minerals or nanoparticles in a biopolymer solution, making it difficult to avoid using crosslinkers or organic solvents. The residual solvents may introduce toxicity concerns [51].

Spraying is a commonly used technique for artificial ligament coatings, which can help reduce the toxicity associated with chemical residues from biochemical reagents [38]. Techniques such as vacuum spraying and plasma spraying are widely applied in this context [38,52]. Among these, vacuum spraying is a promising method, as it can produce uniform nanocoatings without the need for heat treatment [53,54]. On the other hand, plasma spraying involves high working temperatures, which may damage the substrate and potentially alter the crystallinity and phase composition of hydroxyapatite (HAp) [55]. However, some researchers have pointed out that despite the plasma jet reaching temperatures as high as 10,000 K or more, the temperature drops sharply once the jet leaves the nozzle [56]. Nevertheless, ion beam-assisted deposition has stringent production conditions that are difficult to control [57], and thick coatings are prone to cracking during spraying [58].

When performing coating processing, researchers typically aim to achieve stable, well-adhered thin coatings. Although successful integration has been achieved through methods like dip coating, plasma spraying, and other techniques involving bioactive calcium phosphate materials, certain drawbacks remain [53]. In contrast, radio frequency magnetron sputtering (MS) is effective in producing thin coatings with high adhesion strength between the coating and the substrate. MS technology offers fast deposition rates and low substrate temperatures, making it suitable for heat-sensitive materials. The HAp coatings produced by MS show excellent density, uniformity, and high bonding strength [59]. The advantages and limitations of commonly employed coating techniques for ligament applications are summarized in Table 1. In general, to achieve reliable coatings, researchers should choose the appropriate coating technique based on the characteristics of the substrate and coating material, as well as the required coating thickness and layer number. Thin coatings are generally preferred to minimize interfacial delamination and promote better mechanical integration. When choosing an appropriate coating technique, researchers should consider the chemical compatibility between the coating and the substrate, the mechanical requirements (such as flexibility and modulus), as well as the degradation or release kinetics of the coating.

Table 1.

Comparison of commonly used coating methods for ligaments.

Coating Method	Working Principle	Advantages	Disadvantages
Dip coating	Relying on molecular interactions between the coating material and the substrate surface, the substrate is immersed in the coating solution, where a liquid film forms and subsequently solidifies into a stable coating upon curing	The reaction steps are minimal [60]	Insufficient coating rate and lack of precise control over the coating results [61]
Sol–gel	A sol is formed in the solution, which then undergoes gelation by adjusting the conditions, ultimately forming a solid gel material	The process is simple, operates at low temperatures, and has minimal impact on the internal structure, making it suitable for polymer materials with complex geometries and large surface areas [62,63]	The coating is prone to cracking [64]
Electrochemical deposition	Using an electric field to deposit particles from the suspension onto the substrate surface	The equipment is simple, the film formation time is short, and the coating exhibits good uniformity and integrity [41,65]	Residual solvents may pose toxicity risks [51]
Plasma spraying	High-temperature plasma is generated by arc heating of the working gas, and the material is then sprayed at high speed onto the substrate surface to form a coating	The deposition rate is high, and the resulting coating is uniform and stable [38]	The working temperature is high, and process control is difficult [57]
Sputter deposition	In a vacuum environment, high-energy ion beams (usually argon ions) accelerated by an electric field bombard the target surface, causing atoms or molecules to be “sputtered” from the material surface, which are then deposited onto the substrate surface	The coating on flat substrates has a uniform thickness, is dense, and exhibits high adhesion [66]	Long processing time and high cost [67]
PLD	A high-energy pulsed laser is used to bombard the target material, causing the surface to rapidly evaporate and generate plasma. This plasma then condenses onto the substrate, forming a thin film	Ultra-thin coatings (on the order of a few atoms in thickness) can be achieved [68,69]	Splat particles affect the surface quality [66,70]
Biomineralization	The substrate is immersed in simulated body fluid	The material properties are similar to bone minerals, and mild processing conditions allow for the binding of biological molecules [71]	Long soaking time [72]

4. Surface modification of ACL artificial ligaments through coatings

Artificial ligaments offer significant advantages for early return to sports, but the grafts are typically composed of synthetic polymers, which have smooth, hydrophobic surfaces that limit cell adhesion and tissue ingrowth. After implantation, the tendon-bone interface often forms fibrous scar tissue rather than the fibrocartilage transition zone found in natural ACL [23]. Consequently, research on ligament coatings primarily focus on improving biocompatibility and tissue inductivity. The goal is to enhance integration with the host tissue while maintaining the structural and functional integrity of the artificial ligament, promoting the formation of transitional tissues, such as newly formed bone and Sharpey's fibers, that similar to the natural tendon-bone interface (tendon, unmineralized fibrocartilage (UFC), mineralized fibrocartilage (MFC), and bone) [17,73,74]. In recent years, significant efforts have been made to enhance the performance of artificial ligaments, including: improving biocompatibility, promoting cell adhesion and proliferation, and facilitating ligamentization; imparting tissue inductivity to stimulate bone tissue formation and angiogenesis; conferring osteoconductivity to promote bone integration; Enabling controlled and sequential drug release; and Modulating the immune microenvironment to reduce inflammation and minimize fibrotic scar tissue formation (Fig. 3).

Fig. 3.

Overview of surface modification strategies for artificial ligament coatings. (A) Firstly, based on the design of coating location and quantity, segmented or multilayer coatings can be applied to achieve hierarchical modification of the graft. (B–F) Furthermore, due to the diverse functions of coating materials, the coatings can endow artificial ligaments with various properties, including biocompatibility, tissue inductivity, osteoconductivity, drug delivery, and immune modulation.

These strategies provide both theoretical foundations and technical support for more efficient artificial ligament repair. Since most studies use multiple materials to coat ligaments, many coated ligaments have been reported to exhibit multiple effects in animal models. We categorize the coating strategies based on the main focus of the researchers’ efforts.

4.1. Biocompatibility oriented coatings

Biocompatible coatings primarily aim to improve surface hydrophilicity and cell compatibility, promoting cell adhesion and proliferation. Organic materials have shown significant advantages in this regard. For example, hydroxypropyl cellulose (HPC), a non-ionic cellulose ether polymer, is soluble in water and organic solvents, making it an important pharmaceutical excipient. When used as a surface coating for PET grafts, HPC improves hydrophilicity and adhesiveness [75]. Yang et al. [76] coated PET grafts with HPC and implanted them into rabbits, observing improved bone integration within the bone tunnel. Further in vivo studies confirmed significantly enhanced ligament tissue regeneration [77]. Additionally, Lessim et al. [78] grafted sodium poly(sodium styrene sulfonate) (polyNaSS) onto LARS artificial ligaments, improving their biocompatibility and promoting fibroblast colonization, collagen fiber deposition, and integrin adhesion.

Among natural materials, SF is a natural polymer extracted from silkworms, known for its tensile strength and toughness. It has been used for sutures and researched as a coating for PET grafts [79,80]. Cai et al. [81] cross-linked SF onto PET grafts and demonstrated that the organic coating improved hydrophilicity, facilitated fibroblast adhesion and proliferation, and enhanced extracellular matrix secretion. CS, a natural polysaccharide known for promoting wound healing, has been shown to stimulate cell proliferation and collagen synthesis [82,83]. Kawai et al. [84] coated CS onto nonwoven polyester scaffolds and implanted them into rabbit knee joints. They found that compared to uncoated scaffolds, CS coatings enhanced bone tissue formation in the bone tunnel and soft tissue formation in the joint cavity, increasing graft-bone attachment strength.

LBL self-assembly technology, as an effective surface modification method, has been widely applied in ligament coatings. Li et al. [85] employed LBL to load multilayer coatings of hyaluronic acid (HA) and cationic gelatin (CG) onto PET grafts. In vivo studies in rabbits and pigs showed suppressed inflammation, promoted vascular reconstruction, and collagen remodeling. Similar results were observed for ultra-high molecular weight polyethylene grafts [86].

The combination of organic and inorganic materials has also demonstrated great potential in improving ligament biocompatibility. Due to its simplicity, room temperature reactions, high adhesion, and versatility, polydopamine (PDA) coatings have attracted attention in the textile field. It serves as a multifunctional platform for secondary reactions to achieve desired functions [87,88]. Wu et al. [89] utilized the mussel adhesion principle, using PDA as an adhesive interlayer, and loaded nanohydroxyapatite and silver atoms onto PET grafts. This approach enhanced ligament biocompatibility while adding antibacterial properties. In another experiment, Wang et al. [88] developed composite grafts modified with copper/silver/polydopamine, where the addition of copper endowed the grafts with antimicrobial properties and good electromagnetic interference shielding ability. Furthermore, Shi et al. [90] incorporated magnesium silicate into coatings composed of polyallylamine hydrochloride and chondroitin sulfate. In both in vivo and in vitro experiments, the composite-coated grafts promoted bonemarrow derived mesenchymal stem cells (BMSCs) adhesion and proliferation, inhibited fibrous scar tissue formation, and enhanced TBH.

4.2. Tissue inductive coatings

Tissue inductive coatings promote biological processes such as osteogenesis and angiogenesis, making them an important strategy for enhancing graft function and facilitating repair. Currently, most research on the tissue-inductive properties of artificial ligaments focuses on optimizing their osteogenic capacity. Wang et al. [91] loaded graphene onto the surface of PET grafts, ensuring the porosity and fatigue resistance of the graphene coating. Using a rabbit joint graft-bone healing model, they demonstrated that the graphene coating increased bone formation and mineral deposition rates, significantly improving the structural parameters and biomechanical properties of PET ligaments after grafting, thereby promoting early recovery in ACLR. Wang et al. [92] wound oriented CNTs onto PET fibers. Compared to the bone tunnel expansion caused by commercial artificial ligaments made from bare polyester fibers, the CNT composite fibers promoted new bone regeneration and effective bone tunnel repair. In another study, their CNT-based grafts successfully withstood one million continuous bends and complex deformations in ACLR models of rabbits and sheep [23]. Interestingly, Lu et al. [24] fed silkworms mulberry leaves sprayed with a CNT dispersion solution to produce tough “reinforced silk,” offering potential for the application of natural high-strength fibers in artificial ligaments. In another study, Jiang et al. [50] combined inorganic and organic materials by loading a composite coating of SF and HAp on PET grafts. This composite coating promoted new bone formation and reduced scar tissue formation at the graft-bone interface.

Inducing vascular ingrowth is another direction to promote TBH. During bone repair, the ingrowth of new blood vessels helps establish a more structured cellular layout and extracellular matrix, reducing the formation of disorganized connective tissue [[93], [94], [95]]. Li et al. [96] applied pulsed laser deposition (PLD) technology to load copper-containing bioactive glass (Cu-BG) nanocoatings onto PET grafts. This significantly improved the hydrophilicity of the graft surface, promoted cell adhesion and proliferation, and, in a goat model, facilitated bone regeneration and angiogenesis at the graft-bone interface. Additionally, the mechanical stability of the grafts was enhanced, with the maximum pull-out force of the coated grafts being significantly higher than that of the uncoated group.

4.3. Osteoconductive coatings

In addition to organic materials, many inorganic materials have also been evaluated for improving ligament coatings. Bioactive ceramics, such as HAp, are among the most representative calcium phosphate salts. HAp is the most stable form of calcium phosphate (CaP) and is also the most abundant form of calcium phosphate in bones and teeth. It possesses osteoconductive and osteoinductive properties, promoting bone growth on the surface of biomaterials and is widely used as a mature osteoinductive agent in hard tissue engineering. To enhance TBH after ACL reconstruction using artificial ligaments, several animal studies have used HAp coatings to improve the osteoconductivity of grafts and promote TBH [38,39,97]. Furthermore, Cai et al. [37] compared the effects of biomineralization and electropolymerization methods for loading CaP onto PET grafts. They found that the electropolymerization method was more efficient and enabled uniform distribution of CaP particles on the scaffold surface, effectively avoiding the potential issues of inflammation and slower bone formation caused by random CaP particle distribution [98,99]. Interestingly, Saito et al. [62] used the sol–gel method to prepare a titanium dioxide coating, which effectively induced HAp deposition on the PET surface, exhibiting excellent bone bonding ability, osteoconductivity, and bone biocompatibility. In another study, Ma et al. [40] prepared a strontium-substituted HAp coating on PET grafts, which improved surface hydrophilicity, promoted TBH, and significantly increased biomechanical strength. Similarly, Egawa et al. [100] applied a silicate-substituted strontium coating to PET grafts, finding that the coating promoted bone formation.

Bioactive glass (BG) is a biomaterial with osteoconductivity that can release calcium and phosphorus in vivo, combining with ions in body fluids to form a HAp layer. Li et al. [101] demonstrated the positive effect of BG-coated PET grafts on bone integration in a rabbit extra-articular TBH model. In another study, Wu et al. [31] obtained similar positive results in a rabbit IA model. However, despite the promising effects of the aforementioned bioactive coatings in promoting bone integration, studies on their regulation of cellular signaling pathways, cell interactions, and the local microenvironment remain limited. Therefore, future research should further investigate the molecular mechanisms underlying the effects of ligament coatings to optimize their performance and facilitate clinical translation.

4.4. Drug delivery coatings

In addition to directly inducing osteogenesis and vascularization to enhance TBH, coatings can also serve as drug delivery systems. Collagen, one of the earliest natural materials used in ACLR, is also commonly employed as a hydrogel matrix, thus making collagen an ideal drug carrier for controlled and sustained drug release. Simvastatin, which has positive effects on osteogenesis [102,103], was studied by Zhang et al. [104], who applied a biocompatible collagen coating loaded with simvastatin-loaded microspheres onto plasma-modified PET grafts. This approach effectively improved graft hydrophilicity and surface roughness, and upon implantation in rabbit knee joints, it not only promoted osteogenesis but also promoted angiogenesis, enhancing graft bone healing while suppressing osteoarthritis progression. LBL self-assembly technique involves alternating deposition of oppositely charged or affinity molecules to not only functionalize the surface but also serve as a drug delivery system. In another study, a LBL self-assembled coating of CS and HA encapsulated electrospun polycaprolactone (PCL) scaffolds to form a multilayer membrane for the localized delivery of stromal cell-derived factor-1α (SDF-1α) and bone morphogenetic protein 2 (BMP-2). This system successfully induced the recruitment and osteogenic differentiation of BMSCs [105].

Furthermore, multilayer coatings allow for differential degradation and controlled sequential release of functional agents from distinct layers. Building on this concept, Gao et al. [106] designed a multilayer coating capable of time-programmed release of curcumin and Zn²⁺, aiming for anti-inflammatory effects during the inflammatory phase and tissue induction during the regenerative phase. In vivo experiments demonstrated that the coated grafts inhibited inflammatory responses and enhanced TBH in a relatively short period. Similarly, Wu et al. [107] proposed a sequential ion release system utilizing magnetron sputtering to deposit a uniform strontium calcium phosphate (Sr-CaP) coating on the inner layer to support osteogenesis. The outer layer featured an HA-based hydrogel containing Mg²⁺, designed to release magnesium ions in the early phase to facilitate vascular and neural regeneration. Animal studies demonstrated that the synergistic effects of this sequential release strategy significantly improved TBH.

4.5. Immune modulation coatings

Current surface modification strategies, such as SF, graphene, and hydroxyapatite, primarily focus on enhancing osteogenesis, but often overlook the regulation of fibrous scar tissue and inflammation. Studies have shown that macrophage accumulation at the graft-bone interface may contribute to the formation of fibrous tissue, potentially leading to graft loosening or even failure [108,109]. In the early postoperative period of ACLR in sprague dawley (SD) rats, significant inflammatory cell infiltration is observed at the tendon bone interface. During this stage, CD68 positive macrophages, recruited from circulating monocytes, exhibit phagocytic activity and peak at postoperative day 7, followed by a gradual decline. At a later stage, CD163 positive macrophages derived from the local tissue microenvironment peak at day 28, playing anabolic and reparative roles [108]. Song et al. [109] analyzed tissue from the femoral tunnel aperture in ACLR patients and found that postoperative M1 macrophage activation and sterile inflammation are key biological factors contributing to graft loosening. Therefore, promoting the early phenotypic transition of local macrophages from a pro-inflammatory M1 state to an anti-inflammatory M2 state is essential for improved tissue regeneration [110].

To address this issue, Wang et al. [111] developed a multilayer hydrogel coating with immunomodulatory properties and a low friction coefficient. The coating, composed of gelatin methacrylate (GelMA), polyethyleneglycol diacrylate (PEGDA), and sulfated polysaccharide (SCS), for surface modification of PET-based implants. In vivo and in vitro experiments demonstrated that this double network hydrogel enabled controlled release of SCS, effectively inhibited M1 macrophage polarization, promoted M2 polarization, enhanced osteogenesis and angiogenesis, and reduced scar tissue and fibrosis formation. This ultimately facilitated effective tissue regeneration and improved graft stability, showing promising potential for scar-free graft–bone integration.

4.6. Composite functional coatings (overall research on promoting ligamentization and TBH)

In order to reproduce the multilayered tissue gradient characteristics present in the natural tendon-bone interface after implantation, researchers have explored multilayer and multiphase scaffolds, enabling the simultaneous observation of ligamentization and TBH in vivo. Cai et al. [36] proposed a segmented coating for PET artificial ligaments, where bone integration within the bone tunnel and ligamentization within the joint cavity were observed. The joint section was coated with SF, and the bone section was coated with HAp. Both in vivo and in vitro evaluations showed that the coating improved the hydrophilicity and cell compatibility of PET fibers, enhancing the ligamentization and bone integration of the implants. Similarly, Li et al. [112] developed a triple nanocoating for PET implants, consisting of polydopamine, chondroitin sulfate, and BMP-2. Their results showed that this multifunctional coating effectively modulated the local immune microenvironment by promoting macrophage polarization from pro-inflammatory M1 to anti-inflammatory M2 phenotypes, thereby reducing fibrotic proliferation [113]. Additionally, it regulated immune cytokine secretion, promoted BMSC osteogenic differentiation, and significantly improved in vivo bone integration, benefiting TBH.

Previously, to enhance the effectiveness of grafts, many studies attempted to seed scaffolds with stem cells [[114], [115], [116], [117]], achieving some success in promoting ligamentization and TBH. Building on this, Li et al. [118] proposed a triphase modification of silk scaffolds seeded with three types of cells, dividing the implant into tendon, fibrocartilage, and bone regions, each loaded with different coatings and cells. Specifically, the tendon region was coated with SF and seeded with BMSCs, the fibrocartilage region was coated with SF, chondroitin sulfate, and hyaluronate and seeded with chondrocytes, while the bone region was coated with SF and HAp and seeded with osteoblasts. This construct effectively promoted TBH, with mechanical properties and a layered structure of the implant–bone interface resembling that of the natural tendon-bone interface post-implantation.

Considering the highly complex biological environment of the human body, segmented or composite coatings, which integrate the functionalities of different coatings, are expected to become a promising approach in coating design. Overall, coated artificial ligaments have shown promising results in animal models, and we summarize the related in vivo studies in Table 2.

5. Future prospects

Despite significant progress in the research of coated ligaments, many challenges remain to be addressed (Fig. 4).

Fig. 4.

Future Prospects and Expected Advancements in Coated Ligaments. To facilitate clinical translation, animal models should transition from small animals to large mammals and non-human primates to more accurately simulate the human knee environment and postoperative healing process, thereby improving the reliability of in vivo data. Moreover, coated scaffolds remain in the early stages of development, and coating parameters must be optimized and precisely controlled to ensure both functional performance and long term stability. Appropriate coating materials should be selected to construct gradient coatings and smart coatings with responsive capabilities. Overall, while the clinical application of coated artificial ligaments remains limited, continued innovation in this area holds significant potential to advance the optimization of artificial ligament design and functionality.

First, the selection of animal models for ligament reconstruction. Almost all clinical approaches for humans have been initially developed and validated through animal experiments. Before clinical application of coated artificial ligaments, their safety and efficacy must be tested in vivo. However, current research is predominantly conducted in small animal models, such as rabbits and rats, which pose notable challenges for translation to human clinical settings. These challenges include the difficulty in fabricating ligaments of appropriate size for small joints, limited surgical precision due to anatomical constraints, and significant differences in joint biomechanics. For instance, the knee joints of small animals are typically in a more flexed position at rest and bear mechanical loads in a manner distinct from that of humans. Moreover, their faster metabolism and differing immune responses may not accurately reflect the healing dynamics and tissue integration seen in human patients. These differences limit the ability to accurately replicate human joint kinematics, loading conditions, and healing responses. Therefore, to bridge the gap between preclinical findings and human application, future studies should incorporate large-animal models—such as sheep, goats, or non-human primates, which offer joint sizes, anatomical structures, and biomechanical environments more comparable to humans. As a result, large-animal models serve as a critical translational bridge between preliminary preclinical discoveries and human clinical application, improving the accuracy, reproducibility, and regulatory relevance of in vivo studies. Their use will significantly strengthen the foundation for clinical translation of novel ligament reconstruction strategies, including coated artificial ligaments.

Second, the details of the coating. Different coating methods have varying effects on the coating materials and substrates. Surface treatments of the substrate, working temperature, solvent types, material purity, and other factors can all affect the wear resistance and stability of the coating. Additionally, exploring the interactions and bonding strength between the coating and the substrate is crucial for tissue engineering applications. This can help explain how coatings affect the biological and physicochemical properties of the substrate. As the number of coating interfaces increases, the bonding strength between different layers should also be considered. Any interface with insufficient bonding strength could lead to potential coating detachment. Furthermore, few studies have considered the thickness and roughness of the coating on the substrate. Precisely controlling coating thickness, ensuring it provides sufficient protection without compromising the substrate's flexibility and functionality, is a key optimization direction. Very thin coatings may provide inadequate protection or functionality, while excessively thick coatings may crack, peel, or reduce strength. Currently, reducing the coating thickness while ensuring bonding strength and functionality seems to be more favorable for product translation. Optimizing coating process parameters and selecting appropriate interface treatment methods can enhance the bonding strength between the coating and substrate, preventing detachment or peeling of the coating.

Finally, the selection of coating materials. Currently, beyond basic safety, researchers aim to develop artificial ligaments with the following characteristics: (1) ideal mechanical properties, balancing strength and elasticity; (2) strong tissue induction, promoting bone formation and stimulating ligament tissue regeneration; (3) wear resistance and fatigue resistance, with long-term stability; (4) controllable performance, with functional design based on the biological healing process. In addition to exploring new base materials, ligament coatings are an important solution. Most current research suggests that single materials have limitations and are difficult to use in ideal coated ligaments. Looking to the future, combining multiple materials for coating design has more potential. This approach can optimize mechanical performance and biocompatibility while providing superior tissue induction and durability, thereby better promoting ligament regeneration and improving clinical outcomes. Additionally, the strategies developed by researchers for PET-based ligaments are also expected to be applicable to artificial ligaments made from other materials. Furthermore, the application of smart responsive coatings is a promising direction. By integrating sensor technology, materials that respond to external stimuli (such as pH, temperature, or piezo) can dynamically adjust their properties, enabling controlled drug release, surface property modulation, and more. This would improve the performance of artificial ligaments in various physiological environments, enable controlled internal modification of the ligaments, synchronize tissue regeneration, and provide data support for rehabilitation, ultimately optimizing personalized rehabilitation plans.

6. Conclusion

The development of safe and effective artificial ligaments remains a challenging endeavor. However, significant progress has been made in the field of coated artificial ligaments in recent years. This review focuses on the application of various coating materials and coating technologies in artificial ligaments. Overall, surface modification strategies, such as coating, play a crucial role in artificial ligaments by imparting new mechanical properties and interface characteristics. They can also be used to control drug delivery, guide tissue repair, and promote functional tissue engineering. These strategies help regulate the distribution of different cell types, enable spatiotemporal drug release, and facilitate the application of biomimetic layered scaffolds on multi-tissue surfaces, thereby precisely controlling the composition of complex tissues and their interactions with the underlying matrix. This approach brings artificial ligaments closer to the biomechanical and histological characteristics of ACL after reconstruction, offering significant potential for clinical translation. Nevertheless, most studies are still at the animal experimental stage, and the underlying mechanisms require further exploration. In-depth research and careful evaluation are necessary before clinical application.

Author contributions

LY, YW, ZP, and JC contributed to the conceptualization and design of the study. ZP and XF designed and wrote the entire manuscript. ZP, XF, JT, and LY collected the references for the manuscript and prepared the figures. JT, LS, and YW completed the subsequent revisions of the manuscript. All authors contributed to the manuscript revision and approved the final version of the manuscript.

Funding

This work was funded by the National Natural Science Foundation of China (Grant No. 82372491).

Declaration of competing interest

The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

Abbreviation

ACL, anterior cruciate ligament; ACLR, anterior cruciate ligament reconstruction; PET, polyethylene terephthalate; IA, intra-articular; EA, extra-articular; BPTB, bone-patellar tendon-bone; LARS, ligament augmentation and reconstruction system; CNTs, carbon nanotubes; TBH, tendon-bone healing; CS, Chitosan; SF, silk fibroin; LBL, Layer-by-layer; HAp, hydroxyapatite; MS, magnetron sputtering; PLD, pulsed laser deposition; UFC, unmineralized fibrocartilage; MFC, mineralized fibrocartilage; HPC, hydroxypropyl cellulose; polyNaSS poly(sodium styrene sulfonate), HA; hyaluronic acid, CG; cationic gelatin, PDA; polydopamine, BMSCs; bonemarrow derived mesenchymal stem cells, PLD; pulsed laser deposition, Cu-BG; copper-containing bioactive glass, CaP; calcium phosphate, BG; bioactive glass, PCL; polycaprolactone, GelMA; gelatin methacrylate, PEGDA; polyethyleneglycol diacrylate, SCS sulfated polysaccharide; SrSiP silicon substituted strontium, RSF regenerated silk fibroin; SDF-1α, stromal cell-derived factor-1,BMP-2 bone morphogenetic protein 2; Sr-CaP, strontium calcium phosphate; SD, sprague dawley