Articular cartilage degeneration can occur due to injury, aging, or both. Repetitive microtrauma, instability, or undetected intra- or extra-articular injuries may lead to an accelerated degeneration of the joint surface. This process can occur rapidly or over the course of several decades. Because of the growing number of patients afflicted with osteoarthritis at increasingly younger ages, more attention has been directed toward tissue engineering techniques to disrupt or reverse the osteoarthritis development process [83]. A recent biomechanical study reported on the suture retention strength of in vitro prepared neocartilage. The authors found that neocartilage had 33% of the tensile strength of native cartilage. This study also reported survival of neocartilage grafts sutured into osteochondral defect rabbits. The authors concluded that neocartilage can be reliably secured with sutures [84]. Additional biomechanical studies have demonstrated early arthritic changes in adjacent tissues in patients with suture-secured cartilage implants [85]. However, a recent in vitro study of nanoindentation in repair and native cartilage revealed that repair cartilage had a ten times lower contact stiffness than that of native cartilage [86]. Other biomechanical studies have also reported decreased contact stiffness and modulus in repair tissues compared to native articular cartilage [87]. Although techniques to produce neocartilage have advanced, there is still further research needed to develop tissue more similar to that of native articular cartilage.

Many surgeons currently hold microfracture as the gold standard procedure in cartilage restoration for small lesions (less than 1.5 cm²) [88]. However, recent studies have indicated that although microfracture delays cartilage degeneration at short-term follow-up, the beneficial effects may not last beyond 5 years after surgery [89, 90]. Moreover, studies have reported that treatment failure after microfracture is possible regardless of osteochondral defect size [91, 92]. As the shortcomings of microfracture have been elucidated, there has been growing interest in developing treatments with the capacity to fully restore articular cartilage at extended follow-up intervals. In this regard, a recent study reported that repair of chondral injury using a hyaluronic acid-based scaffold with activated bone marrow aspirate concentrate provides better clinical outcomes and more durable cartilage repair at medium-term follow-up compared with microfracture [93].

Gobbi et al. investigated the clinical outcome in a group of active patients with large full-thickness chondral defects of the knee treated with one-step surgery using bone marrow-derived MSCs and a second-generation matrix. The authors reported that MRI scans showed good stability of the implant and complete filling of the defect in 80% of patients, and hyaline-like cartilage was found in the histological analysis of the biopsied tissue at a minimum of 3 years. No adverse reactions or postoperative complications were noted [94].

Autologous chondrocyte implantation (ACI) has developed into a promising new frontier for cartilage restoration, with potentially more consistent results than microfracture. Case series with >10 years follow-up have shown ACI to be effective and biomechanically durable for large (>4 cm2) chondral defects [95, 96]. Peterson et al. evaluated indentation measurements at the central part of each ACI graft and reported a range of 0.3–3.7, which was 90% or more of the value observed in the control group [95]. Gooding et al. [97] recently compared ACI to osteochondral transplantation at 10 years follow-up and reported superior outcomes in the ACI group compared to the osteochondral transplant group. Similar findings have been reported in other studies reporting on patients with >3 cm2 osteochondral lesions [98, 99]. These studies have increased the confidence in ACI as a reliable cartilage regeneration tool; however, further studies are needed to identify the specific indications for ACI and the long-term intra-articular biomechanical benefits of ACI. One of the major drawbacks of ACI is the need for a biological matrix to ensure correct placement of ACI in the joint being treated. Some studies have reported using porcine membrane with mixtures of type I and II collagen and/or hyaluronic acid scaffolds [97, 100, 101]. However, using such grafts may expose the patient to developing an immune reaction and have a subsequent failure of the ACI procedure. The use of matrices in ACI is necessary; however, future studies should focus on developing non-immunogenic and biomechanically stable ACI matrices.

Scaffold-based chondrocyte implantation has also emerged as a promising treatment for osteochondral defects. Hyaluronic acid-based scaffolds have been shown to facilitate development of hyaline-like cartilage 1 year after implantation in early case series [101, 102]. In a recent prospective cohort study, Kon et al. reported comparable patient-reported outcomes between a group that received a hyaluronic acid-based matrix-induced autologous chondrocyte implantation (MACI) or microfracture [102]. However, at 7.5 years follow-up, the hyaluronic acid-based MACI patients had significantly better IKDC scores than the microfracture group. The authors who also performed a sub-analysis of return to sport in the athletes in their cohort found that microfracture required 8 months of recovery, whereas the hyaluronic acid scaffold plus MACI patients required 12.5 months of recovery [102]. Gobbi et al. compared the clinical outcomes of two similar groups of patients with patellofemoral full-thickness cartilage lesions, treated with MACI or BMAC, employing the same scaffold [103]. Both groups showed significant improvement in all scores, from preoperative to final follow-up (P = 0.001), but there was no significant difference in improvement between the two groups, except for the IKDC subjective score (P = 0.015), which favored the BMAC group. Deterioration in MACI and improvement in BMAC group scores were noticed, from 2 years to final follow-up, but were nonsignificant [103]. Given the longer rehabilitation period of the scaffold plus MACI group, future studies must focus on decreasing rehabilitation times while maintaining biomechanical viability of MACI treatment.

Scaffold-free constructs also offer an option for cartilage replacement. These options are prepared by culturing bovine chondrocytes in high-density molds over 8 weeks. By week 1, these constructs could be handled [104], and by week 4, they had biomechanical properties similar to that of juvenile cartilage [105]. However, other biomechanical studies have demonstrated that cartilage repair tissue to have more rapid cyclic compression values compared to normal articular cartilage. Although scaffold-free constructs reduce the risk of immunologic reaction at the time of implantation, further studies are needed to ensure the scaffolds can withstand the biomechanics of the joint that is implanted into.

The future of cartilage regeneration lies within the realm of in vitro chondrocyte-seeded scaffold development. This process involves the introduction of chondrocytes into a 3D matrix, which is subsequently cultured in vitro for 4–6 weeks [106]. This process produces neotissue that has produced promising short- and long-term patient-reported outcome results [107–109]. Because walking exposes articular cartilage to force equal to 4–5 times a person’s body weight [110], in vitro development chondrocyte-seeded scaffolds may not produce biomechanically robust constructs. Because of this, researchers have begun to expose cell-laden matrices to hydrostatic pressure [107] or dynamic compression [108] to enhance matrix maturity and biomechanical strength. In a phase I FDA clinical trial, eight patients received autologous chondrocyte-implanted collagen I scaffolds, which were subjected to hydrostatic pressure prior to implantation. At 1 year after surgery, seven of the eight patients had almost completely filled chondral defects, with mature repair tissues [111]. In the subsequent phase II trial, the scaffold technique was found to have a similar safety profile to microfracture, but was found to have better clinical outcome scores at 2 years follow-up.

Tissue engineering has demonstrated a distinct ability for cartilage regeneration that could potentially advance the field of treating both focal chondral lesions and OA by providing a functional biological equivalent tissue to replacing lost or damaged tissue. The standard concept of tissue engineering is to seed cells on three-dimensional (3D) biomaterial scaffolds to help regenerate damaged tissue. The scaffold is designed to create a 3D environment that promotes tissue regeneration through the cells that are seeded within the scaffold [112, 113]. Multiple scaffolds have been proposed for tissue engineering. Among these, Shimomura et al. compared ahydroxyapatite (HA)-based artificial bone coupled with a mesenchymal stem cell (MSC)-based scaffold-free tissue-engineered construct (TEC) with a beta-tricalcium phosphate (bTCP) for repair of osteochondral defects in a rabbit animal model. The authors found that osteochondral defects treated with the TEC/bTCP implants showed more rapid subchondral bone repair at 1 month, but the cartilaginous tissue deteriorated over time out to 6 months after implantation. Osteochondral defects treated with the TEC/HA implants maintained good histological quality out to 6 months after implantation and also exhibited better biomechanical properties at 6 months as compared with the TEC/bTCP implants [114].

Furthermore, biodegradable hydrogels have been suggested as a promising scaffold for articular cartilage tissue engineering. A main advantage of using hydrogels is the ability to inject the hydrogel as a solution in situ and then polymerizing it in vivo. In addition, recent advances in bioprinting have granted tissue engineers the ability to assemble hydrogels into anatomically relevant functional tissues or organ parts. Moreover, hydrogels could also be potentially used as drug delivery systems, allowing a controlled and sustained release of drugs intra-articularly over several weeks or months [115] for the treatment of joint disease such as OA or rheumatoid arthritis (Fig. 2.3).

The body of the literature concerning cartilage tissue engineering in animal models is rapidly expanding; however, it has been reported that 90% of the new approaches that are successful in animal studies subsequently fail clinical trials [116]. Therefore, meticulous analysis of the existing short-term clinical outcomes is advocated to be able to guide the mixed biomimetic and bioactive approach to more successful and reliable clinical outcomes.

Much of the orthobiologics biomechanical literature has focused on strengthening partially torn rotator cuffs or enhancing rotator cuff repairs. Early studies focused on defining an ideal scaffold to enhance rotator cuff repair strength. Barber et al. [117] proposed a dermal allograft augmentation to strengthen supraspinatus tears. Using ten pairs of human cadaveric samples, they cyclically tested augmented and non-augmented repairs at 10 and 100 N for ten cycles at 20 N/S, followed by destructive testing at 33 mm/S. Bio-mechanical evaluation revealed a significantly greater failure load in the augmented repair shoulders. Furthermore, the two groups had different failure mechanisms: the non-augmented repairs failed at the tendon-suture anchor interface while augmented repairs failed via suture breakage. Moffat et al. [118] engineered a poly(lactide-co-glycolide) [PLGA] nanofiber scaffold to improve rotator cuff repair strength by guiding fibroblast attachment and fiber orientation. Biomechanical analysis revealed that cell distribution and nanofiber organization over the PLGA scaffold were biomechanically superior to those outside of the scaffold area, despite the fact that the PLGA scaffold degraded during in vitro testing. In a later study, Derwin et al. studied eight adult dogs that underwent bilateral shoulder surgery: one shoulder underwent tendon release and repair, while the other shoulder underwent release and augmentation [119]. At time zero, the repair augmentation shoulder had significantly higher failure load. At 12 weeks, the poly-l-lactide scaffold group had significantly greater cross-sectional area, stiffness, and ultimate load than repairs without augmentation. The authors concluded that although using a scaffold augmentation provides a medium for host tissue deposition and ingrowth, it does not fully prevent tendon retraction [119].

Further studies have demonstrated the biomechanical properties of extracellular (ECM) rotator cuff repair augmentation patches to be inferior to that of the native rotator cuff [120]. However, many techniques to reproduce an ECM with properties similar to that of the native rotator cuff have also been described [121–124]. In animal models, rotator augmentation with a porcine intestine patch demonstrated superior stiffness compared to traditional repair but failed to show an increased load to failure [125]. Interestingly, most porcine intestine patch studies report that the graft is weakest at 1–2 weeks after surgery, but increases to strength similar to native rotator cuff tendon by 3 months [126, 127]. Unfortunately, these promising findings produced in the porcine intestine patch models have not scaled into human models. Scalmbert et al. reported MRI confirmed failure in 10 of 11 patients following porcine patch placement for large rotator cuff tears [128].

Cytokine-based therapy has also been proposed and evaluated to enhance rotator cuff strength. Using a sheep model, Rodeo et al. [129] introduced BMP-(2–7) and TGF-(b1–3) in addition to fibroblast growth factor into a type I collagen sponge at the rotator cuff tendinous insertion. At 6 and 12 weeks after surgery, a higher failure load was found in the cytokine plus collagen scaffold group compared to the traditional reconstruction group. However, when these findings were normalized based on tissue volume, no significant differences were found. This study elucidated that although cytokine therapy may accelerate healing, it may not have the capacity to increase rotator cuff reconstruction strength.

Currently, there are two commercially available partial meniscus substitutes: one is made up of type I collagen fibers from bovine Achilles tendon and the other is made up of polyurethane polymers with interdigitating polyester and semi-degradable stiff segments. Early studies on the type I collagen-based implants in humans demonstrated no adverse effects, formation of new tissue, and good clinical outcome scores at 36 months [130]. A later study compared patient outcomes between patients treated with the collagen-based meniscus substitute and patients who underwent partial meniscectomy. This study reported superior patient outcomes in the patients treated with the collagen-based meniscus substitute [131]. In contrast, Rodkey et al. reported that 311 patients with Outerbridge scores less than grade IV were either treated with the collagen-based meniscus substitutive or partial meniscectomy [132]. This study reported that the collagen-based meniscus replacement resulted in improved clinical outcomes in patients with medial meniscus defects, but was not effective for patients with acute chondral injuries or lateral meniscal defects [132]. Despite promising clinical findings, magnetic resonance imaging (MRI) studies have revealed several salient biomechanical changes in the collagen-based meniscus substitutes. The size of the meniscus substitute reduces significantly during follow-up interval, and the MRI signal intensity of meniscus substitutes does not match that of the native meniscus at any follow-up periods [133, 134]. These findings indicate that although clinical results have been encouraging, further studies need to evaluate the biomechanical integrity of meniscus substitutes so that patients can have sustained outcomes beyond 10 years.

Regarding total meniscectomized patients, allograft transplantation remains the gold standard for symptomatic patients [135]. However, meniscal allograft transplants have been observed to undergo collagen remodeling [136] and shrinkage [137], which may lead to alteration of biomechanical properties. Because of these limitations, attention has been directed toward developing a synthetic total meniscus substitute. Early research focused on the biomechanical viability of Teflon and Dacron as meniscus substitutes [138–140]. However, these designs were limited by a lack of biocompatible materials and poor biomechanical performance. Later studies reported on a porous polyurethane scaffold as a total meniscus substitute. Pore size and compressive properties were tuned to stimulate local tissue growth and differentiation into fibrocartilaginous tissue, based on previous studies [141, 142].Compressive properties of the implant increased 24 months after implantation and were not different from native meniscus properties [143]. However, the meniscus implant was not strong enough to resist shear forces in the knee, and collagen type and orientation were not meniscus-like upon further evaluation [143]. Although there has been promising clinical and biomechanical studies reporting on meniscus substitutes, clinicians must remain focused on ensuring that these implants can withstand the stresses of daily life and/or athletic competition so that patients can remain symptom-free over a longer period of time.