Tissue engineering of cartilage, tendon and bone

Abstract

Tissue engineering aims to produce a functional tissue replacement to repair defects. Tissue reconstruction is an essential step toward the clinical application of engineered tissues. Significant progress has recently been achieved in this field. In our laboratory, we focus on construction of cartilage, tendon and bone. The purpose of this review was to summarize the advances in the engineering of these three tissues, particularly focusing on tissue regeneration and defect repair in our laboratory. In cartilage engineering, articular cartilage was reconstructed and defects were repaired in animal models. More sophisticated tissues, such as cartilage in the ear and trachea, were reconstructed both in vitro and in vivo with specific shapes and sizes. Engineered tendon was generated in vitro and in vivo in many animal models with tenocytes or dermal fibroblasts in combination with appropriate mechanical loading. Cranial and limb bone defects were also successfully regenerated and repaired in large animals. Based on sophisticated animal studies, several clinical trials of engineered bone have been launched with promising preliminary results, displaying the high potential for clinical application.

Keywords

Tissue engineering / cartilage / bone / tendon / recent advances

Introduction

Limited availability of donor tissue is still a major obstacle for the repair of tissue defects. Tissue engineering offers a promising perspective in tissue regeneration by employing cells and biodegradable biomaterials to reconstruct or restore living tissues in vitro or in vivo [1]. Significant progress has been made in this field since the 1980s when it was first defined. Advancements have been made from the use of mature, differentiated cells as a cell source to multipotent stem cells; from primitive biomaterials with simple structures to adjustable biomaterials with specific shapes and sizes; from descriptive demonstration to more mechanistic insights; and from bench to some attempts, if not many, to bedside. This review placed emphasis on the recent progress in engineering of cartilage, tendon and bone in our laboratory.

Cartilage reconstruction

Cartilage reconstruction is one of the most widely pursued feats in tissue engineering. In our laboratory, we focused on the in vivo reconstruction of articular, auricular and tracheal cartilages.

Tissue engineering has been employed to repair articular defects with chondrocytes in mammal animal models [2,3]. As the main cell type in cartilage, autologous chondrocytes are an obvious choice for the cell source in cartilage engineering, as they are already programmed to express cartilage-specific genes and deposit extracellular matrices (ECMs) [4]. In our laboratory, we have successfully reconstructed hyaline cartilage and repaired full-thickness articular cartilage defects with autologous chondrocytes in vivo in a porcine model [5]. An 8-mm full-thickness defect was created by trephining the weight-bearing areas of medial and lateral femoral condyle. The defects in the experimental group were then repaired by pressing the cell-scaffold constructs into the defects manually. Non-cell-seeded scaffolds or untreated defects served as controls. The results showed that typical hyaline cartilage structure was generated, with ideal interface formation between the engineered cartilage and the adjacent normal cartilage at 24 weeks post-implantation. In addition, glycosaminoglycan (GAG) levels of the engineered cartilage (around 110 μg/mg) reached 80% of that of native cartilage (around 130 μg/mg). This study provides insight into the clinical translation of this technology to repair articular cartilage defects. However, some cartilage formation was observed in the subchondral bone region [5]. For this reason, bone marrow stem cells (BMSCs) were utilized to repair osteochondral defects, since they can undergo both chondrogenesis and osteogenesis. Articular osteochondral defects were repaired using autologous BMSCs seeded on a polyglycolic acid/polylactic acid (PGA/PLA) scaffold. Osteochondral defects were created in non-weightbearing areas of the knee joints (two in each knee) and were repaired by a chondrogenic factor-induced BMSC-PGA/PLA construct. Dexamethasone-treated BMSC PGA/PLA constructs, non-cell-seeded PGA/PLA constructs or unrepaired defects served as controls. After 6 months, all of the defects in the experimental group were completely repaired by hyaline cartilage and cancellous bone in their appropriate anatomical locations. In the control group utilizing dexamethasone-treated BMSCs, only 5 of 16 defects were repaired with hyaline cartilage and cancellous bone, while the others were repaired by fibrocartilage and cancellous bone. In contrast, no obvious repair but only fibrotic tissue was observed in the control group without cells or those left untreated. Moreover, the compressive moduli of reconstructed cartilage in the group treated with chondrogenically induced BMSCs (30.43±3.45 MPa) reached 80.27% of the normal amount (37.95±1.94 MPa) at 6 months, with a high level of GAG content (9.40±1.02 mg/g) that was not statistically different from normal cartilage (10.08±0.36 mg/g). These results demonstrated that the BMSCs and PGA/PLA complex could form engineered cartilage and bone to repair osteochondral defects (Fig. 1). In order to elucidate the mechanism by which BMSCs-seeded PGA/PLA scaffolds repair the defects, we retrovirally labeled BMSCs with green fluorescent protein (GFP) before in vivo implantation, which provided reliable evidence that implanted BMSCs could undergo chondrogenic and osteogenic differentiation in vivo in articular osteochondral defects to form hyaline cartilage and subchondral bone [6].

Ear reconstruction is one of the earliest models of tissue engineering. As early as the 1990s, Cao et al. engineered human ear-shaped cartilage 12 weeks after subcutaneous delivery of bovine chondrocytes seed on PGA scaffolds into nude mice [7]. In contrast to articular cartilage, auricular cartilage requires more precise fabrication technologies because of its more sophisticated and complicated structure. It is of great significance to construct a prefabricated bioscaffold in the shape of an ear with mechanical integrity so that it can retain its original shape and size during culture and chondrogenesis. In our laboratory, an ear-shaped PGA/PLA scaffold was fabricated by computer-aided design and manufacturing (CAD/CAM) technique according to a mirror-symmetric model built from a computed tomography (CT) scan of a patient’s normal ear. The PLA content was optimized by balancing the scaffold's biocompatibility and mechanical strength. The ear-shaped scaffold reached a level of similarity of 97.8% prior to culture and 86.2% after 12 weeks of culture, separately, largely retaining its original shape and size (Fig. 2). Furthermore, the constructs formed ear-shaped cartilage-like tissues at 12 weeks with abundant cartilage extracellular matrices, mature lacunae, and with good elasticity and mechanical strength. This study demonstrates a useful strategy for reconstructing cartilage with complex shapes in an in vitro approach [8].

Due to the lack of a sufficient cell source, BMSCs have been introduced as a novel cell source in ear reconstruction. As mentioned above, BMSCs can be induced toward chondrogenic lineage in vitro by soluble factors. However, application of high levels of soluble factors led to hypertrophy and calcification in vivo [9,10]. In our laboratory, we took advantage of a coculture system with chondrocytes and BMSCs to address these problems. Interestingly, we discovered that whereas BMSCs alone failed to form engineered cartilage in a subcutaneous area, they could successfully differentiate into chondrocytes when cocultured with mature chondrocytes, both subcutaneously and even in vitro [11,12]. These results indicate that the chondrogenic niche provided by chondrocytes is sufficient to induce BMSC chondrogenesis. Notably, soluble factors secreted by chondrocytes, such as transforming growth factor β1 (TGF-β1), insulin-like growth factor-1 (IGF-1) and bone morphogenetic protein-2 (BMP-2), are currently used as inducers for chondrogenic differentiation of mesenchymal stem cells [13]. They play an important role in directing BMSC chondrogenesis in vivo in a coordinated way, even though their contents in coculture media were greatly lower than those used in conventional induction protocols [11]. It sheds a light to further application of coculturing in ear reconstruction.

Functional maintenance and structural completeness of tracheal cartilage is vital to human health. Tracheal defects, especially those of long segments, usually require substitutes for tissue reconstruction. In our laboratory, satisfactory tubular cartilage has been generated with autologous chondrocytes-PGA/PLA constructs in a rabbit model for tracheal reconstruction in vivo following 2 weeks of in vitro precultivation. This time in in vitro culture also alleviated the post-implantation inflammatory reaction when compared with constructs implanted after just two days of in vitro cultivation [14]. Autologous tubular chondrocyte-PLA/PGA constructs were subcutaneously implanted into rabbits, after in vitro culture for either 2 weeks or 2 days. In the group that was precultivated for 2 weeks, most of the fibers were found to be completely embedded in ECM produced by the chondrocytes. Importantly, no obvious inflammatory reaction was observed after in vivo implantation and homogeneous cartilage-like tissue was formed with biomechanical properties close to native tracheal cartilage at 4 weeks post-implantation. In the group with just 2 days of precultivation, however, an obvious inflammatory reaction was observed within and around the cell-scaffold constructs at 1 week implantation and only sporadic cartilage islands separated by fibrous tissue were observed at 4 weeks. Other than PGA/PLA complex, it has also been reported that a functional trachea was constructed by seeding mesenchymal stem cell-derived chondrocytes into decellularized trachea, and used to replace the diseased bronchus in a patient with a success [15]. In order to promote cell migration while maintaining the original structure of the matrix, we proposed a sandwich model with acellular cartilage matrix for cartilage engineering as follows: cartilage was cut into pieces and decellularized to achieve acellular sheets. Chondrocytes from recipient were seeded on each acellular sheet and the cell-sheet constructs were stacked on top of one another for the total of 20 sheets to engineer new cartilage. In this model, native chondrocytes were easily removed from the sheets, while the seeded cells were evenly distributed in the constructs, so that homogenous neo-cartilage was formed in vitro or in vivo (unpublished data). This strategy may be useful for engineering cartilage with sophisticated structure.

Tendon reconstruction

Tenocytes, spindle-like and elongated in shape, are naturally considered an optimal cell source in tendon engineering and have been utilized to reconstruct tendon in some studies. The chicken craw is a good animal model due to its high structural similarity to the tendons in human hands. We used this model for flexor tendon engineering inside the tendon sheath by using autologous tenocytes and PGA fiber scaffolds. The results demonstrated that engineered tendon after 1 week of in vitro culture and 14 weeks of in vivo repair was similar to normal tendon in its gross view, histology and biomechanical strength (around 90 N), providing a possibility that functional tendon can be generated in vivo [16].

In contrast to the flexor tendon, the extensor tendon is a complex structure that performs important hand functions. We isolated human fetal extensor tenocytes and seeded them onto PGA fibers formed in a scaffold that mimicked the human extensor tendon complex. After in vitro culture for 6 weeks, some constructs were further cultured in vitro with dynamic mechanical loading for an additional 6 weeks in a bioreactor. Constructs were implanted subcutaneously into nude mice for another 14 weeks with or without mechanical loading. The results demonstrated that conditioning with dynamic mechanical loading in a bioreactor leads to a more mature and compact tissue in vitro, but the generated tissue tends to become atrophic. Interestingly, the maturation of the engineered tendon was significantly enhanced by suturing the constructs to the animals' fascia, by which animal movement provided natural mechanical loading to the constructs. Even without mechanical loading in vivo, the implanted tendon revealed relatively more mature collagen superstructure than in vitro samples, as evidenced by the formation of more fibrils possessing an apparent D-band. However, these fibrils appeared in a random and disorganized pattern due to lack of proper mechanical loading. After mechanical loading in vivo, engineered tendons increased more in tissue volume and became more mature in their gross view, tissue structure, and collagen superstructure with stronger mechanical properties, when compared to in vitro loaded tendon and in vivo non-loaded tendon (Fig. 3) [17]. This study indicates that in vivo mechanical loading might be an optimal niche for engineering functional tendon.

Dermal fibroblasts are more accessible for future clinical applications due to less donor site morbidity and more abundant cell sources than tenocytes. Considering the facts that both tenocytes and dermal fibroblasts are derived from the mesoderm and have similar characteristics in cell morphology and extracellular matrix components, skin fibroblasts were also examined for tendon engineering in vivo and in vitro. We utilized a porcine model to explore the possibility of tendon engineering using dermal fibroblasts. Constructs were formed with autologous dermal fibroblasts seeded on PGA unwoven fibers and cultured in vitro for 7 days before in vivo implantation to repair a defect of flexor digital superficial tendon. Constructs seeded with tenocytes were used as a control. After 26 weeks following implantation, both the fibroblast- and tenocyte-engineered tendons exhibited histology similar to that of the natural tendon. Parallel collagen fibers were observed throughout the tendon structure. Furthermore, both fibroblast- and tenocyte-engineered tendons reached similar tensile strengths of 31.0±2.2 MPa and 32.4±3.6 MPa, respectively, or about 75% of natural tendon strength (42.4±3.9 MPa) [18]. Based on this, we further investigated the possibility of engineering human neo-tendon tissue in vitro using dermal fibroblasts. Since tendon maturation is dependent on local mechanical stimulation, a bioreactor is necessary to provide a mechanical loading to engineered tendons. In our laboratory, a stainless-steel U-shape spring was employed to introduce tensile force to a scaffold of PGA unwoven fibers fixed between hooks at its two ends. The PGA scaffold could be kept tension-free or constantly strained by adjusting the distance between two ends of the spring [19]. These scaffolds were seeded with human dermal fibroblasts and compared to those seeded with human tenocytes. The fibroblast/PGA constructs formed neo-tendon tissue in vitro under static strain and the tissue structure became more mature with increase in culture time, reaching a peak at 14 weeks. Interestingly, there was no significant difference between the neo-tendon tissues generated with human dermal fibroblasts and with human tenocytes [20]. These results suggest that human dermal fibroblasts can replace tenocytes for future tendon graft development in vitro.

Bone reconstruction

BMSCs are considered a viable cell source for bone engineering. Successful repair of bone defects has been demonstrated using BMSCs in small immuno-deficient animals, such as athymic mice [21,22], as well as in large immunocompetent animals [23,24]. BMSCs have recently been used in allotransplantation for bone regeneration without harmful immune responses in a canine model [25], in several other animal models of allogeneic rejection [26-28], and in patients [29], showing a promising prospect in bone regeneration. We have also reported that both primary and passaged adipose stem cells (ASCs) have immunosuppressive capacity and retain low immunogenicity when treated with IFN-γ [30], providing a perspective into clinical application using allogeneic ASCs.

In our laboratory, we created bilateral full-thickness parietal defects (20 mm in diameter) in sheep to examine the potential of BMSCs to repair cranial defects. The defects were repaired with implants constructed of osteogenically induced BMSCs encapsulated in calcium alginate. Defects treated with calcium alginate alone or left unrepaired were used as controls. Reconstructed bones were observed in the defects as early as 6 weeks after implantation, with more mature bone formation at 18 weeks. Almost complete repair of the defects was shown at 18 weeks with three dimensional CT scanning. In contrast, the defects in control groups were filled with fibrous tissues and remained unrepaired at 18 weeks post-surgery [31].

After cranial bone was successfully engineered, the same technique was utilized to repair alveolar defects in a canine model. A horizontal alveolar bone defect was created by removing alveolar bone (5 mm high at buccal side) at the location of the third and fourth mandibular premolars and the first molar to expose tooth roots. The defects were then repaired by in vitro induced BMSCs in calcium alginate. Defects repaired by calcium alginate alone or left untreated were used as controls. Bone nodule structure was observed at 4 weeks post-surgery in engineered bone, which became more mature after 12 weeks. In contrast, limited bone regeneration was found in control groups with the alveolar crests adjoining to the notch at that time point [32].

To explore the possibility of repairing critical-sized mandibular defects using engineered bones, we created a 30 mm segmental mandibular defect in a canine model and repaired it using BMSCs on porous, slow-degrading β-tricalcium phosphate (β-TCP) scaffolds. New bone formation was observed from 4 weeks post-operation, and bony-union was detected by radiographic and histological examination after 32 weeks. Control defects treated with β-TCP scaffolds alone were repaired with fibrous tissues, with minimal new bone formation at the cutting end in the defects at 32 weeks. More importantly, the engineered bone with BMSCs/β-TCP achieved satisfactory biomechanical properties at 32 weeks post-operation, which was very close to those of contralateral edentulous mandible and autograft bones. However, 30%~40% of the biomaterial remained after 32 weeks post-operation, even though new bone had already formed [33]. In an ideal tissue engineering strategy, the degradation rate of the implanted biomaterial should be synchronous with the pace of tissue development. New bone formation within the period of natural bone healing would be more applicable for clinical use. It is also likely that the remnants of ceramic material would impede the organization of the newly forming bone. To address these issues, we introduced highly resorbable coral as a scaffold in the same model. Less biomaterial residue of the BMSCs/coral grafts (10%~15%) was observed after 32 weeks, and the engineered bone achieved satisfactory biomechanical properties that were very close to those of the contralateral mandible (Fig. 4) [34].

Furthermore, we challenged the capacity for engineering of weight-bearing bone by repairing femoral bone defects (25 mm long) in a goat model using osteogenically induced BMSCs and coral. In the experimental group, bony union was observed at 4 months, and engineered bone was further remodeled into newly formed cortical bone at 8 months. Most importantly, the tissue-engineered bone segment was similar to the left-side normal femur with no significant differences in terms of bend load strength (1380±183 N vs. 1452±102 N) and bend rigidity (6.13±0.45 N × M/ϕ vs. 6.89±0.69 N × M/ϕ) at 8 months. In contrast, the coral cylinders of the control group without BMSCs showed no bone formation (Fig. 5) [35]. These results showed that the critical-sized segmental bones could be regenerated with BMSCs and coral in the femoral defects within 8 months.

Several studies have reported the successful repair of bone defects in humans with tissue engineering [36-38]. Our laboratory has started a clinical trial of engineered bone with 11 patients since 8 years ago, mainly focusing on the repair of human craniomaxillofacial bone defects using autologous BMSCs and partly demineralized bone matrix (pDBM) [39]. In this study, we have shown that implantation of osteogenically induced BMSCs were essential for stable bone formation in human, as complete resorption was observed with CT scanning and histology when non-cell-seeded pDBM was implanted. More importantly, implanted BMSCs caused new bone formation in humans, which was stable for 18 months, when a biopsy demonstrated the formation of cancellous bone and the expression of osteocalcin and osteopontin by the engineered human bone. However, major challenges still need to be further addressed, such as size limitation, insufficient mechanical strengths at early times, and vascularization of engineered bone.

In recent years, ASCs have attracted considerable attention for their application in tissue engineering because of their high proliferative capacity, multilineage potential, abundant autologous cell availability, and minimal injury to the donor site. Similar to BMSCs, ASCs especially have strong potential of osteogenesis [40] and have the ability to generate bone and repair bone defect in vivo [41], indicating that fat tissue may be a promising cell source in bone regeneration. To understand the long-term stability of the ASC-derived tissues, long-term results still need to be further investigated.

Conclusions and perspective

Significant progress has been achieved in tissue engineering in recent years. In our laboratory, we focus on the engineering of cartilage, tendon and bone. Our previous studies demonstrated that various types of these tissues can be engineered in vivo and in vitro and can be used to repair defects in large animal models and in humans. Notably, a few successful clinical trials utilizing engineered bone point to a promising future for clinical application of engineered tissues. More attention should be paid in further studies to the establishment of a cell source that provides a homogenous and viable cell population in sufficient numbers; the mechanism by which the local niche of tissues promotes differentiation and maturation of target tissues; and the design of biomaterials that are more suitable to tissue reconstruction and closely to human physiological conditions, so as to facilitate the clinical application of tissue engineering.

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