1. Introduction
The anterior cruciate ligament (ACL), which connects the femur and tibia, is a major stabilizer of the knee joint
[1]. ACL injury causes excessive anterior tibial translation, destabilizes the knee joint, and accelerates the degeneration of both cartilage and bone in the joint
[2],
[3],
[4]. Annually, over 400 000 anterior cruciate ligament reconstruction (ACLR) operations are performed worldwide
[5]. Among these cases, 54% to 74% of injured people have successfully return to sports activities
[6],
[7]. Rehabilitation is affected by numerous factors: the integration status of the ligament and bone within the bone tunnel and ligamentization within the intra-articular portion determine the progression of postoperative rehabilitation, function, and biomechanical strength of the reconstructed ligament
[8],
[9]. In addition, although previous studies have reported ACLR success rates exceeding 90%
[10], with 63% of subjects reporting good to excellent outcomes
[11], the rate of post-ACLR revision surgery was reported to range from 10% to 15% in a recent study
[12]. Radiographic results from a 20-year follow-up study of athletes who underwent non-operative treatment for ACL injury also demonstrated a similar frequency of osteoarthritis signs
[13]. These observations highlight the demand for biological enhancement in ACLR.
Great progress has been made, and previous reviews have drawn conclusions about the results of clinical application of artificial ligaments
[14],
[15],
[16] and preclinical studies on the modification of artificial ligaments
[9] and have explored perspectives on applying bioactive coatings and surface modification methods to artificial ligaments
[16]. However, no comprehensive review has outlined the research and development of artificial ligaments from bench to bedside or the promising materials and emerging technologies for improving artificial ligaments, resulting in a research knowledge gap. First, an ideal substitute for the ACL should exhibit mechanical strength comparable to that of the native tissue
[14]. Second, as the healing of a reconstructed ACL can be generally divided into two parts, namely graft–bone integration inside bone tunnels and ligamentization of the intra-articular portion, the biological and histological performance of the chosen ACL substitute should closely mimic those of the native ligament after tissue regeneration and remodeling. In other words, any biomaterials that can promote these healing processes are theoretically able to accelerate the healing of a reconstructed ACL. Moreover, biocompatibility, which is characterized by inertia and a minimal inflammatory response, was previously considered to be a basic requirement of substitute materials
[17]; however, the bioactivity of biomaterials has been increasingly recognized as an essential feature that accelerates tissue healing
[18]. Therefore, the purpose of this review is to introduce graft materials and associated healing processes in the context of ACLR (2 Current graft options for ACLR, 3 Healing of the reconstructed ACL), evaluate the challenges associated with current artificial graft strategies applied in ACLR (Section 4), and highlight the surge in bioactive artificial ligaments (Section 5).
2. Current graft options for ACLR
2.1. Autografts
The main types of autologous grafts are the bone–patellar tendon–bone (BPTB) graft, hamstring tendon (HT), and quadriceps tendon (QT)
[19],
[20]. Previously, surgeons considered the BPTB graft to be the standard for ACLR because of the natural property of its ligament, which connects bone to bone
[21]. However, common postoperative complications associated with the BPTB graft include anterior knee pain, difficulty kneeling, patellar fracture, and patellar tendon rupture
[22]. The HT graft has been reported to induce less anterior knee pain and a lower incidence of long-term osteoarthritis than the BPTB graft
[23],
[24]. However, complications associated with the HT graft, such as weakness in hip extension and terminal knee flexion, graft laxity, an increased infection rate, increased objective laxity in female subjects over time, tendon truncation during harvesting, and variable graft sizes and lengths, remain problematic
[25],
[26]. The QT graft is a more recent alternative for ACLR; this graft is thicker and has more favorable tensile properties than the BPTB and HT grafts
[27]. However, the QT graft is associated with postoperative complications such as donor site pain and a higher graft failure rate than that associated with the BPTB or HT graft
[28],
[29].
2.2. Allografts
The clinical applications of allografts for ACLR vary from BPTB, HT, and QT to the tibialis anterior tendons
[30],
[31]. The tensile load and stiffness of allografts are similar to those of autografts
[32]. However, sterilization using irradiation has been found to adversely affect the biomechanical properties of allografts, delay healing, and increase the risk of graft failure
[33],
[34],
[35],
[36]. Allografts have some advantages, including the elimination of donor site morbidity, fewer postoperative limitations on activity, and less knee pain than that associated with autografts
[37],
[38]. In addition, allografts are more controllable than autografts in terms of the length and diameter, which are also important for healing
[39],
[40]. Of note, immunologic responses and virus transmission are concerns associated with allografts. One report described an association of fresh-frozen allograft use with postoperative inflammatory synovitis, and the transmission incidence of bacteria such as staphylococcus and peptostreptococcus was estimated to be 0.15%
[41]. In addition, the allograft failure rate is 2.3 times higher in young subjects (< 30 years) than in older ones
[33].
2.3. Artificial grafts
Artificial ligaments have been a popular means of overcoming the drawbacks of autografts and allografts since the 1950s
[1]. Various materials have been explored and used in the construction of artificial ligaments, including but not limited to polypropylene (PP) in the 1970s, polytetrafluoroethylene (PTFE) and carbon fibers in the 1980s, and polyethylene terephthalate (PET) in the 1990s
[42]. Artificial ligaments have advantages with respect to donor site morbidity and disease transmission. Nevertheless, complications such as chronic effusions, synovitis, and graft failure are not rare. To date, artificial ligaments available in the market continue to undergo optimization
[14],
[15],
[43] (
Fig. 1).
3. Healing of the reconstructed ACL
In general, the healing of a reconstructed ACL involves two parts: ① graft–bone integration inside bone tunnels and ② intra-articular ligamentization. Orthopedic surgeons must have a thorough understanding of both stages to choose an appropriate graft and be able to initiate the optimal rehabilitation protocols after ACLR. Graft healing is increasingly accepted as occurring in three main stages: the early stage, which is accompanied by necrosis; the middle stage, which involves cell proliferation and tissue remodeling; and the late stage, which involves tissue maturation. These three stages occur in both the intra-articular portion and inside the bone tunnel
[44],
[45]. Nonetheless, the two healing sites undergo specific and detailed processes at different stages, which can last from several weeks to months (
Fig. 2).
Due to the lack of bioactivity in existing artificial ligament materials, the above-described healing process occurs only when autografts and/or allografts are used for ACLR. Therefore, how to improve the bioactivity of artificial ligaments through tissue engineering while ensuring reliable mechanical performance and biocompatibility is a research question worth exploring. Such exploration aims to provide a postoperative healing process more similar to that of autografts and allografts.
3.1. Graft–bone integration
At the early healing stage, the graft mid-substance undergoes necrosis due to avascularity and insufficient nutrition and oxygen supply
[46],
[47]. Next, the recruitment of neutrophils and macrophages to the graft occurs as early as one week postoperatively
[48]. Relatively small amounts of cells and vessels can be observed during this stage
[49],
[50]. Various cytokines and factors, including matrix metalloproteinase-1 and -13 (MMP-1 and MMP-13), tumor necrosis factor-α (TNF-α), and interleukin-6 (IL-6), are released during necrosis
[50],
[51],
[52]. This stage lasts about 4 weeks
[53],
[54].
The middle stage is distinguished among the three stages by maximum cell proliferation, and these cells may be derived from various tissues, including bone marrow, synovium, synovial fluid, and the graft
[55],
[56],
[57],
[58]. In addition, angiogenesis from the periphery toward the central part of the graft and progressive mineralization of the graft–bone interface tissue occur gradually, with the subsequent ingrowth of newly formed bone into the outer layer of the graft and integration of the graft into the surrounding bone
[59],
[60],
[61],
[62]. This stage occurs approximately 4 to 12 weeks after reconstruction
[53],
[54]. The final stage can be defined as continuous remodeling and incorporation of the graft and bone. Collagen fibers between the graft and bone undergo a gradual reconstruction of continuity, leading to the re-establishment of a graft–bone junction, or enthesis
[63],
[64],
[65].
Two types of entheses have been reported in previous studies: direct and indirect entheses. The former type comprises four structurally contiguous but histologically gradient layers: graft, unmineralized fibrocartilage, mineralized cartilage, and bone
[65]. The latter type is composed of obliquely or perpendicularly arranged fibers, known as Sharpey’s fibers, inserted into the surrounding bone
[66]. Both direct and indirect entheses are recognized as good outcomes, and the chondrocytes surrounding the graft–bone interface play a temporary role and eventually undergo endochondral ossification
[55],
[67],
[68],
[69].
3.2. Tissue remodeling of the intra-articular portion
Tissue remodeling of the intra-articular portion is referred to a histological alteration of the graft when implantation is performed as a substitute for the native ACL
[46]. Similar to the early graft-stage process that occurs inside the bone tunnel, the intra-articular portion undergoes tissue necrosis and inflammation. In addition, the mid-substance of the graft may be subject to extended inflammation and necrosis caused by immunological cells and cytokines, accompanied by circulating synovial fluid
[70]. Interestingly, Rougraff and Shelbourne
[71] showed that the intra-articular portion of a BPTB graft underwent early revascularization instead of tissue necrosis in the central region of the graft as early as 3 weeks post-ACLR. Other studies involving animal models also reported that soft-tissue grafts such as the HT and extensor digitorum longus did not undergo early necrosis postoperatively
[72],
[73].
During the middle stage, cell proliferation is accompanied by angiogenesis, and tissue remodeling is followed by cellular metabolism
[74]. From the perspectives of cellularity, vascularity, and extracellular matrix (ECM) alteration, as tissue remodeling continues, the cells transform from an oval to a spindle shape and align along the longitudinal axis of the graft. Hypervascularization that occurs at the early stage slowly decreases to a normal level during the middle stage, and the regenerated ECM, especially collagen type III, increases as the original ECM gradually degrades
[75],
[76].
Adequate cellular repopulation and tissue remodeling rates are crucial determinants of prognosis
[77]. On one hand, deficient cell proliferation and revascularization may cause inferior tissue reorganization, thus postponing the process of rehabilitation
[77]. On the other hand, excessive cellular and histological activities may over-weaken the mechanical properties of the graft
[44]. Excessive type III collagen, which is inferior to type I collagen in terms of mechanical strength, may have negative effects on the graft and lead to an increase in the graft failure rate. An excess of cells throughout the graft undergo necrosis and decrease to the normal level seen in an intact ACL by 6 to 12 months postoperatively, and vessels are distributed evenly in the graft
[44],
[78]. However, the endpoint of remodeling is not clear because certain changes continue to occur even years after reconstruction
[49],
[63].
Undoubtedly, cellularity and osteogenesis are the crucial factors affecting intra-articular tissue remodeling and graft–bone integration, respectively. Autografts are recognized as the gold standard, mainly due to their intrinsic bioactive advantages in cell adhesion, proliferation, and osteogenesis, as these grafts contain cells, ECM, and cytokines to facilitate healing processes. In comparison, artificial ligaments often lack tissue remodeling and graft–bone integration abilities because of their poor biological activities
[79],
[80]. As a result, how to improve the bioactivity of artificial ligaments remains a challenging topic.
4. Artificial ligaments from biocompatibility to bioactivity: experience and inspiration
In past decades, attempts have been made to construct ideal artificial ligaments via tissue engineering methods, using materials ranging from synthetic polymers to natural silks, seed cells, and various types of growth factors. Biomaterials for tissue engineering have been selected with an emphasis on biocompatibility
[81]. Advances in regenerative medicine and engineering technology have substantially broadened the concept of biomaterials. For example, in the field of artificial ligament construction, clinical practice has proven that biocompatibility is not the only requirement for biomaterials. Biological effects such as angiogenesis and osteogenesis also play critical roles in the healing process after ACLR
[82]. Here, we sort information about the research and development of artificial ligaments, including ligament scaffolds and fixation devices, using different materials. Although some types of artificial ligaments have fallen into disfavor because of their unsatisfactory medium- and long-term follow-up results, the demand for artificial ligaments continues to increase around the world because of the popularity of all types of sports, which has led to increases in sports-related injuries. As a result, research and development efforts focused on new forms of artificial ligaments are important for meeting these demands.
4.1. Ligament scaffolds
To date, several types of materials, mostly polymers, have been applied as ligament frameworks
[83]. As the full history of the use of synthetic devices for ACLR is outside the scope of this review, the representative materials are listed instead. A structured literature search was performed in the Cochrane Library (from 1996 to June 30, 2023), Embase (from 1910 to June 30, 2023), and PubMed (from 1979 to June 30, 2023) using the following key words: “anterior cruciate ligament reconstruction,” “Active Biosynthetic Composite,” “ABC ligament,” “Goretex,” “Gore-Tex,” “Kennedy ligament augmentation device,” “Kennedy-LAD,” “ligament augmentation and reconstruction systems,” “LARS,” “Leeds–Keio,” and “LK ligament.” Research articles, including randomized controlled trials (RCTs), observational studies, comparative studies, cohort studies, and case series with at least ten subjects, on primary ACLR surgery using the artificial ligaments listed above and having minimum postoperative follow-up times of 2 years were included in this review. Animal studies, cadaveric studies, reviews, commentaries, retracted studies, studies with no available full text, studies without postoperative quantitative scores, studies on ACL revision surgery, and case series with fewer than ten subjects were excluded. For cohort studies that included the same patient cohort but reported different follow-up time points, only the latest publication was included in this review. The process of selecting eligible studies is presented in
Fig. 3.
Twenty-five research articles were included in the qualitative synthesis; however, no studies were subjected to meta-analysis due to the heterogeneity of the studies and inconsistent reporting of outcomes. Clinical outcomes, including the total number of subjects at the final follow-up, follow-up time, postoperative subjective clinical assessments (Lysholm score, Tegner score, and Internal Knee Documentation Committee (IKDC) score), and complications, were recorded. Graft failure was defined as partial or complete rupture of the reconstructed ligament or when “graft failure” was recorded by an included study. The scores of different postoperative subjective clinical assessments and the occurrence of complications were listed only if the studies reported the outcomes. Otherwise, the outcome was given as “not recorded.” The clinical outcomes of different artificial ligaments are presented in
Table 1,
Table 2,
Table 3,
Table 4,
Table 5 [83],
[84],
[85],
[86],
[87],
[88],
[89],
[90],
[91],
[92],
[93],
[94],
[95],
[96],
[97],
[98],
[99],
[100],
[101],
[102],
[103],
[104],
[105],
[106],
[107].
The Gore-Tex ligament (W. L. Gore; Flagstaff, USA) is a permanent prosthesis composed of PTFE; it was applied for clinical ACLR in the 1980s
[84],
[85],
[86].
Table 1 [84],
[85],
[86] shows the clinical outcomes of studies involving application of the Gore-Tex ligament. The follow-up times ranged from 2 to 15 years. The Lysholm score shows a declining trend over time. The most frequently reported complications were graft failure, effusion, and infection.
The Active Biosynthetic Composite (ABC; Surgicraft Ltd., UK) ligament is composed of braided carbon and polyester fibers and was widely applied in the late 1980s and early 1990s
[83],
[87],
[88],
[89]. Research using the ABC ligament for ACLR has mainly focused on the middle-term (3–7 years) follow-up results (
Table 2 [83],
[87],
[88],
[89]). Bashaireh et al.
[83] reported an inferior IKDC score when comparing the ABC ligament with an autograft. The complications associated with the ABC ligament mainly include graft failure, joint stiffness, and knee synovitis.
The Kennedy Ligament Augmentation Device (Kennedy-LAD; 3M, USA) is composed of braided PP and has been applied mainly to augment grafts for over three decades
[90],
[91],
[92],
[93],
[94],
[95],
[96]. Most relevant research has compared the Kennedy-LAD with autografts for ACLR in terms of the short-term (2–4 years) and long-term (25–30 years) follow-up results; the clinical outcomes with respect to the mean Lysholm and Tegner scores were similar between graft types, although the Kennedy-LAD was associated with a higher rate of graft failure (
Table 3 [90],
[91],
[92],
[93],
[94],
[95],
[96]).
The Leeds–Keio (LK; Neoligaments Ltd, UK) artificial ligament, composed of PET, was jointly designed by Seedhom at Leeds University and Fujikawa at Keio University in 1982
[97],
[98],
[99],
[108]. A 5-year follow-up analysis showed no significant differences between the LK ligament and autografts in terms of the mean Lysholm and Tegner scores, although other studies have reported inferior scores for the LK ligament
[97]. Graft failure and pain during exercise were the most frequently reported complications associated with the LK ligament (
Table 4 [97],
[98],
[99]).
The ligament advanced reinforcement system (LARS; Surgical Implants and Devices, France) is another artificial ligament composed of PET and has been applied clinically since the 2000s
[100],
[101],
[102],
[103],
[104],
[105],
[106],
[107]. LARS has been compared with various grafts, including autografts, allografts, and other artificial grafts, in a series of clinical studies. The follow-up results, ranging from 4 to 16.5 years postoperatively, have demonstrated satisfying outcomes according to the Lysholm, Tegner, and IKDC scores. However, attention has been given to important complications such as graft failure, screw loosening, and knee synovitis (
Table 5 [100],
[101],
[102],
[103],
[104],
[105],
[106],
[107]). As shown in
Table 1,
Table 2,
Table 3,
Table 4,
Table 5, the LARS ligament has demonstrated the most satisfactory clinical outcomes for ACLR to date, based on the short- to long-term follow-up results, when compared with other types of artificial ligaments.
4.2. Fixation devices
One of the most widely applied fixation devices in the bone tunnel is the titanium (Ti) interference screw, which is popular because of its good mechanical strength
[109],
[110],
[111]. Apart from the representative metallic interference screw, which remains permanently in the implant site, some bioabsorbable screws have been developed using polymers such as poly-L-lactic acid (PLLA), polylactic acid (PLA), and polyglycolic acid (PGA)
[2],
[112],
[113]. The focus has shifted recently from the idea of providing stable fixation to the concept of accelerating new bone formation in the bone tunnels for the reconstructed ACL. In addition, with the advancement of surgical techniques, reports have increasingly described cortical suspension with a Ti button and the hybrid fixation method, which uses both an interference screw and cortical suspension
[114],
[115],
[116]. In the past decade, some clinical studies have also investigated the efficacy of using degradable interference screws, and some researchers have additionally studied the use of hybrid screws containing osteoinductive materials, such as β-tricalcium phosphate (β-TCP) and hydroxyapatite (HA), in ACLR
[117],
[118],
[119],
[120],
[121],
[122],
[123],
[124],
[125]. The Lysholm, Tegner, and IKDC scores at follow-up were not significantly different from the scores associated with Ti screws. However, studies on hybrid screws have reported complications such as bone edema, cysts, bone tunnel enlargement, and screw breakage; accordingly, which osteoinductive materials to add to fixation devices, and how to do so, remain controversial topics.
4.3. Modification of artificial ligaments
As mentioned in the previous section, some artificial ligaments have been abandoned, mainly because of a high incidence of graft failure; additionally, some of these ligaments may cause recurrent effusion or foreign body reactive synovitis. Currently, the LARS ligament appears to be the most commonly used, especially in mainland China, due to its relatively low incidence of complications
[42]. Of note, the LARS ligament is fabricated from PET, and the accompanying Ti interference screws are both composed of biocompatible but not bioactive materials.
Studies that have focused on various methods of modifying PET ligaments and fixation devices are listed in
Table 6 [126],
[127],
[128],
[129],
[130],
[131],
[132],
[133],
[135],
[136],
[137],
[138],
[139],
[140],
[141],
[142],
[143]. ECM components, including hyaluronic acid (HA) and collagen, have been added to ligament scaffolds to enhance cell adhesion and proliferation and thus facilitate tissue regeneration
[132],
[134]. The use of ECM analogs such as silk, also called silk fibroin (SF), to modify PET ligaments has been proven to facilitate the adhesion and proliferation of fibroblasts
in vitro and to promote the growth of autologous tissue into intra-articular portion of the ligament
[51],
[84],
[138],
[146]. In addition, Wang et al.
[144] attempted to replace PET with a different material. In that study, the authors developed a hierarchical helical carbon nanotube ligament for ACLR in rabbits and goats. The postoperative results showed that compared with the control group, which was treated using PET ligaments, the experimental group exhibited enhanced osseointegration in bone tunnels and a higher pull-out force.
Other studies on the modification of fixation devices were carried out in animal models and used various promising materials to replace the unabsorbable metal (mainly Ti) screws. The modification attempts have largely focused on improving osteogenesis during graft–bone integration. One common method involved adding HA, a component of the bone tissue frame, to the intraosseous parts of PET ligaments or the interference screw; subsequently, the
in vivo results showed enhanced new bone formation following this approach, compared with a PET ligament or PLLA screw alone
[80],
[136]. Another method involved the addition of bioactive metal or metal compounds. Coating of the intraosseous parts of PET ligaments with sodium hydroxide (NaOH) and GRGDSPC peptide, or with copper bioactive glass, was observed to improve bone regeneration at the graft–bone interface in animal models, and mechanistically, this effect may have been driven by the enhanced expression of osteogenic genes such as
S100A10,
BMP2, and
OCN [82],
[127].
In recent years, magnesium (Mg) has attracted much attention from medical researchers, especially those in the fields of orthopedics, traumatology, and sports medicine
[147]. Mg has been proven to efficiently promote osteogenesis via the increased release of calcitonin gene-related polypeptide (CGRP) by dorsal ganglion root cells in the spinal cord
[59],
[148]. Downstream cells and signaling molecules, such ascyclic adenosine monophosphate (cAMP)-responsive element binding protein 1 (CREB1) and osterix (SP7) in periosteum-derived stem cells (PDSCs), were upregulated and thus enhanced the osteogenic effect of Mg
[148]. Additionally, increased vascular endothelial growth factor (VEGF) levels were found to support the neo-formation of type H (CD31
hiEndomucin
hi) vessels in the metaphysis, periosteum, and endosteum
[149],
[150], which in turn promoted bone regeneration
[146]. Further studies on the use of Mg to modify interference screws have reported considerable progress. In animal experiments, the use of Mg interference screws in ACLR led to accelerated fibrous tissue mineralization and the formation of distinct fibrocartilage transition zones at the graft–bone interface, whereas these effects were not seen when Ti screws were used
[137],
[138],
[139],
[144]. Mechanistically, Mg promoted the recruitment of BMSCs toward the peri-implant bone tissue and the local upregulation of bone morphogenetic protein (BMP-2), VEGF, TGF-β1, and platelet derived growth factor (PDGF)-BB, as well as the downregulation of MMP-13
[137],
[138],
[139],
[140],
[144]. Researchers also have investigated the use of Mg-based alloy screws to balance mechanical properties, degradation speeds, and bioactivity. In animal experiments, Mg-based alloy screws were shown to increase the peri-tunnel bone mass, thus creating a sufficient connection between the graft and bone
[140],
[141] (
Fig. 4).
The canonical configuration for ACLR was constructed based on ligament scaffolds and fixation devices. However, advances in materials and technology have led to increasingly diverse configuration patterns and surgical techniques. Li et al.
[132] developed an all-in-one artificial ligament composed of a silk scaffold with a TCP/PEEK anchor and demonstrated its use in a porcine model of ACLR. The results of that study showed that TCP substantially promoted graft–bone attachment, and the transition zone from the silk graft to bone, which was characterized as regenerated fibrocartilage, was similar to that seen in native bone attachments. Our team has developed an Mg-based alloy (MgZnGd) wire that can be sutured into the intraosseous portions of ligament grafts to promote bone formation during the early healing stage and enhance fibrocartilage-like tissue formation during the late healing stage
[145]. Bioactivity such as the fibroblastic effect exerted by silk and the osteogenic effect exerted by Mg, along with the emerging all-inside surgical technique for ACLR, will continue to inspire the research and development (R&D) of novel bioactive artificial ligaments.
Nevertheless, these attempts at R&D should first aim to balance feasibility, safety, and healing, while bridging the gap between animal studies and clinical demands
[151],
[152]. Feasibility can be enhanced through more advanced technology and better materials, and healing can be improved by ensuring safety. In other words, by building on the results of previous animal experiments and clinical trials, and through meticulous verification of
in vitro and
in vivo experiments, we aim to address the limitations of prior studies and reach a new, elevated equilibrium in terms of feasibility, safety, and the healing effect.
5. Conclusion and future perspectives
To date, the design approaches for artificial ligaments and fixation devices have reflected the prevalence of ACL injury and attempts to continuously improve surgical outcomes. Previous designs of ACL-related implants have mainly focused on biocompatibility based on material mechanical properties, emphasizing the biosafety or inertia of implanted materials. Recent attempts are increasingly shifting toward accelerating the postoperative healing process by endowing materials with bioactivity, aiming to enhance or eventually replace the ACL with neo-regenerated tissues. The studies discussed in this review demonstrate the desirability of considering the demands of the postoperative regeneration process involving angiogenesis, osteogenesis, and fiber synthesis, in addition to biocompatibility.
Some aspects of design are of significance for refining artificial ligaments. First, emergent manufacturing processes, such as electrospinning and three dimensional (3D) printing, may help to improve the physical and biological properties of artificial ligaments and thus yield better postoperative results. In their review, Legnani and Ventura
[16] included studies on the application of coating and surface modification methods, especially surface soaking, in the manufacturing of artificial ligaments. The studies included in that review reported enhanced angiogenesis and osteogenesis; however, they have also reported the nonuniform coating of bioactive components and low cell affinity. Currently, surface soaking is not an ideal method for modifying polymer fibers such as PET as the components may be subject to local agglomeration, which would restrict biological function to a limited region. In addition, the morphology, thickness, and degradation speed of the coated materials are difficult to control via surface soaking
[153]. Advances in tissue engineering techniques such as electrospinning have shown promise for the fabrication of homogeneous porous scaffolds with a uniform distribution of coated components. Shalumon et al.
[154] used the conjugated electrospinning method to fabricate electrospun fiber-coated PCL yarn bundles with evenly distributed surface coating materials and demonstrated that these could enhance tissue regeneration in an extensor digitorum tendon defect model. Lui et al.
[155] developed a 3D-printed, porous, multiphasic bone–ligament–bone scaffold, which demonstrated both favorable cell affinity and homogeneous distribution in an
in vitro model.
Second, direct modification of the materials is also a pivotal factor in artificial ligament optimization. Natural silk materials that possess both reliable mechanical properties and cell affinity are increasingly the focus of investigation. Silkworm silk is one such material. Li et al.
[132] developed an artificial scaffold from silk fiber and demonstrated a robust transition zone from a silk graft to bone in a porcine model of ACLR. However, the application of raw silkworm silk may induce an immunogenic response in the host if the outer layer of sericin, an antigenic protein, is not removed thoroughly
[132]. Additionally, this risk can be avoided by using spider silk, another natural silk with stronger mechanical properties and lower immunogenicity than silkworm silk
[156]. Kornfeld et al.
[157] collected raw spider silk and implanted it into a nerve defect model as an artificial nerve scaffold. The results indicated that spider silk had a capacity for supporting axonal regeneration comparable to that of an autologous nerve graft. However, natural spider silk yields are limited because the cannibalistic nature of spiders makes it very difficult to feed them at a scalable level
[158]. As a result, transgenic technologies have been investigated and used to enable the large-scale production of recombinant spider silk protein (spidroin). Dinjaski et al.
[159], Martín-Moldes
[160], and Li et al.
[161] have developed a series of spidroin-based materials such as recombinant chimeric spidroin films, which were shown to improve the growth, proliferation, and osteogenic differentiation of human mesenchymal stem cells. These recombinant spidroins may be useful in a wide range of orthopedics and sports medicine applications, including the fabrication of artificial ligaments.
Last but not least, the biological features of components also provide additional value. Various types of bioactive materials have been included and investigated in the modification of artificial ligaments, including but not limited to proteins, cells, and metallic elements
[127],
[129],
[140]. However, adding more bioactive materials will not necessarily translate to better clinical outcomes. More research is needed on the key upstream biological effects and their stimulation by as few types of materials as possible. For example, during the regeneration process of ACLR, it is essential to comprehend both the graft–bone integration and intra-articular tissue remodeling processes, as these two distinct portions of the graft require different regenerative and integrative approaches in ACLR. Although multiple factors play a role in ACL rehabilitation, the leading factor remains largely unclear. Further studies are required to clarify the signaling pathways involved in the healing process, such as the molecules downstream of CGRP, and to further investigate the cytokine regulatory networks linking neural perception, osteogenesis, and angiogenesis.
The significance of these enhancements of artificial ligament bioactivity lies in the fact that while the applications of biomaterials in real clinical scenarios are very limited, innovations in this area can prompt research to optimize artificial ligaments. However, the need to balance and optimize the mechanical and biological properties of artificial ligaments remains challenging. Currently, biological properties have been sacrificed to some extent to ensure mechanical strength, and vice versa. Future research should focus on the innovation of biomaterials, novel structural designs such as integrated all-in-one devices to increase overall mechanical strength, and improvements in surgical techniques, such as all-inside reconstruction surgery to reduce bone loss in tunnels. All of these attempts may lead to better clinical outcomes when using artificial ligaments in ACLR.
In conclusion, the healing status of an implant after ACLR is an instrumental determinant of clinical outcomes and, therefore, a key criterion for evaluating the postoperative prognosis. To this end, the use of traditional biocompatible materials with few advantages in terms of healing is shifting to the use of bioactive materials, emphasizing a focus on accelerating tissue regeneration and integration in the field of ACLR; however, the outcomes of currently available products are still incomparable to those of native ligaments. This review contributes to our current knowledge of ACLR by providing details about progress in the field. On one hand, the surge in modification of materials and technologies has largely boosted investigations of tissue regeneration and bioengineering and a shift from biocompatibility to bioactivity. On the other hand, more efforts is needed to conquer the remaining barriers separating bench from bedside.
Acknowledgments
This work was supported by the Areas of Excellence Scheme (AoE/M402/20), the Knowledge Transfer Project Fund (KPF24GWP05), the IdeaBooster Fund (IDBF24MED12), the HKSTP Ideation Fund (Ideation 23-0698), and the ITC Technology Start-up Support Scheme for Universities (TSU24MED05). Professional English language editing support was provided by AsiaEdit.
Compliance with ethics guidelines
Haozhi Zhang, Xin Chen, Michael Tim-Yun Ong, Lei Lei, Lizhen Zheng, Bingyang Dai, Wenxue Tong, Bruma Sai-Chuen Fu, Jiankun Xu, Patrick Shu-Hang Yung, and Ling Qin declare that they have no conflict of interest or financial conflicts to disclose.
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