Abstract
Osteochondral lesions treatment and regeneration demands biomimetic strategies aiming physicochemical and biological properties of both bone and cartilage tissues, with long-term clinical outcomes. Hydrogels and scaffolds appeared as assertive approaches to guide the development and structure of the new osteochondral engineered tissue. Moreover, these structures alone or in combination with cells and bioactive molecules bring the mechanical support after in vitro and in vivo implantation. Moreover, multilayered structures designed with continuous interfaces furnish appropriate features of the cartilage and subchondral regions, namely microstructure, composition, and mechanical properties. Owing the potential as scaffolding materials, natural and synthetic polymers, bioceramics, and composites have been employed. Particularly, significance is attributed to the natural-based biopolymer silk fibroin from the Bombyx mori silkworm, considering its unique mechanical and biological properties. The significant studies on silk fibroin-based structures, namely hydrogels and scaffolds, towards bone, cartilage, and osteochondral tissue repair and regeneration are overviewed herein. The developed biomimetic strategies, processing methodologies, and final properties of the structures are summarized and discussed in depth.
Keywords
1 Introduction
Tissue engineering (TE) field has been evolved as a way to compensate the limited supply of donor tissue and organ transplants related to a high morbidity and mortality [1]. TE approaches involve different research areas simultaneously, including cell biology, materials science, and clinical evaluation, with the final purpose of creating a suitable microenvironment that mimics the tissue at the host site in a desired and faster regeneration [2]. Such microenvironment is typically composed by three-dimensional (3D) porous scaffolds , on which cells are stimulated to grow and organize to form an extracellular matrix (ECM) used to initiate the regenerative process [3]. These 3D constructs provide the chemical and mechanical support for in vitro ECM formation , being gradually degraded, resorbed, or metabolized after in vivo implantation. Therefore, apart from an essential biocompatibility, the scaffolds must possess an equivalent degradation profile to the host tissue, while keeping the mechanical properties and structural integrity promoted by the forming ECM [4, 5].
Natural- and synthetic-based polymers, ceramic materials, and composites have been proposed for scaffolding strategies in TE approaches [6, 7]. Natural-based polymers have emerged as preferred sources for the development of scaffolds with better biocompatibility and lower risk of metabolized degradation products, while the synthetic polymers are more stable and easier to process and modify [8, 9]. For example, collagen, gelatin, and chitosan include some of the most investigated natural polymers in TE field. However, these scaffold materials may present poor mechanical properties associated to a rapid degradation profile. Structural proteins like elastin, fibrin, silk, and albumin have also been used as sutures, and more recently for scaffolds production, and as drug delivery agents [10, 11]. Among them, special interest has been attributed to silk protein produced by a wide range of arthropods and lepidopteran insects, including spiders, scorpions, mites, flies, and silkworms, which possess a large molecular weight of 200–350 kDa, or more. It has been used for centuries in textiles production and as clinical sutures (good skin affinity). Their availability for large-scale processing was also economically advantageous for use in TE applications [12]. From the different sources of silk proteins, Bombyx mori silk produced by silkworms became the most investigated for diverse TE applications, holding impressive mechanical properties, biocompatibility, biodegradability, low immunogenicity, and suitable processability [12]. It is synthesized in a liquid state in the epithelial cells of the insects’ glands, secreted in the lumen and converted (spun) into a liquid-to-solid state (fibers) when in contact with the external air, being mechanically drawn in the form of cocoons. The spun fibers are composed by two animal-based proteins: a core protein named fibroin, surrounded by a glue-like protein named sericin [13]. Even it has been found that sericin may contain some biocompatibility problems [12], several studies have been proposing both silk fibroin (SF) [14,15,16,17] and silk sericin [18,19,20] for diverse biomedical and TE applications. SF has been recognized for presenting favorable biocompatibility, tunable molecular structure, and remarkable mechanical properties with controllable degradation rates, and for that reason it remains the most extensively studied silk protein as promising candidate for several structural [21], biomedical [16, 17], and TE applications [14, 15, 22]. Until now, several forms have been used to fabricate scaffolds made of SF, including films [23], membranes [16], fibers [16], textiles [15, 22], sponges [14], and hydrogels [24], used for the regeneration of soft tissues, like skin [23], ligament and tendon [25], and different hard tissues, including bone [15], cartilage [26], and osteochondral (OC) tissue [14].
The natural OC tissue existing in all human body joints is composed by two main tissues, the articular cartilage and subchondral bone, connected by a stable interface that unifies both elements as a single and complex tissue [27]. The great challenge of OC TE is the design of structures able to meet all the physicochemical and biological requirements for the repair/replacement of bone and cartilage tissues, and at the same time ensure a good compatibility between these two phases.
OC defects or damage can happen in any joint in human body, and affect both the articular cartilage, the underlying subchondral bone, as well its interface [27]. Moreover, the OC tissue repair involves a deep understanding of how the OC interface is combined in terms of structure, and mechanical and biological properties. Over the past years, several studies have been reported towards the repair/regeneration of these tissues, by creating single-phase structures able to fit into the defected area [26, 28,29,30,31,32].
Considering the heterogeneity of OC tissue, innovative 3D structures comprising different mechanical properties and biological performances, according to the target OC tissue layer, are strongly required. The recent development of bilayered scaffolds and improved multi-phased, or stratified, scaffolds with distinct subchondral bone and cartilage layers have been applied for this purpose [33, 34]. In general, OC TE strategies can be categorized into monophasic, biphasic, and triphasic depending on the physicochemical and cellular/biological characteristics of the scaffolds (Table 14.1) [35]. Bilayered hydrogels [36] and complex bilayered scaffolds have been reported for OC regeneration applications [37]. Recent studies in the field are leading to promising approaches to use SF-based biomaterials for OC tissue repair and regeneration strategies [38, 39]. The most recent and relevant studies focused on SF-based structures, namely hydrogels and scaffolds, targeting bone, cartilage, and OC tissue repair and regeneration are overviewed. Additionally, it is summarized developed strategies, processing methodologies, and final properties of the structures.
2 Characteristics of Silk Fibroin
Silk fibroin (SF) from Bombyx mori is composed by three protein components. A heavy (H) and a light (L) chain polypeptides of ~ 350 kDa and ~ 26 kDa, respectively, form the H-L complex linked by a single disulfide bond at the C-terminus of the H-chain (Fig. 14.1a). This H-L complex is also non-covalently linked to a glycoprotein (P25) of ~ 25 kDa in a ratio of 6:6:1 to form micellar units [40]. The H-chains are composed by hydrophobic domains containing highly ordered amino acid sequences repeats, capable of organizing themselves together into β-sheet or crystalline structures through intramolecular or intermolecular forces, including hydrogen bonding, van der Waals forces, and hydrophobic interactions, forming the basis for the tensile strength of SF (Fig. 14.1b) [41, 42]. The L-chains are smaller and less ordered than the H-chains, relatively elastic, and its sequence is not involved in the formation of the crystalline region in SF (amorphous region) (Fig. 14.1c) [40]. The hydrophobic repetitive domains that compose a H-chain are also interspaced by hydrophilic regions (Fig. 14.1a) [43]. These repeating units are known as Ala-Gly-Ser-Gly-Ala-Gly, where glycine (Gly, ~ 43–46%), alanine (Ala, ~ 25–30%), and serine (Ser, ~ 12%) are the three simplest and most abundant amino acids. The next most abundant amino acids in H-chain are tyrosine (~ 5%), the larger amino acid with a polar side chain (semi-crystalline), valine (~ 2%), followed by aspartic acid, phenylalanine, glutamic acid, threonine, isoleucine, leucine, proline, arginine, lysine, and histidine, present in much smaller percentages (less than 2%) [40, 44].
The complexity of SF can be demonstrated by their four different structures (silk I, silk II, silk III, and random coil), formed through different physicochemical stimuli, and that can be transformed to each other under proper conditions [45]. Silk I is formed alternatively by α-helix and β-sheet main conformations, while silk II is rich in β-sheet content and corresponds to the main structural configuration of SF providing high mechanical and physicochemical properties. Silk III is formed by a threefold α-helix crystal structure, and the random coil structure usually exists in the SF solution [44, 46]. By controlling the crystalline and amorphous domains of SF structures (size, number, distribution, orientation, and spatial arrangement), it is possible to produce SF-based matrices with distinct mechanical properties, degradation profile, and aqueous processability, which makes this protein attractive for distinct biomedical and TE applications [12].
3 Silk Fibroin-Based Hydrogels
Hydrogels formation follows several distinctive requirements that mimic the natural ECM microenvironment of tissues [47]. The hydrophilic nature of the ECM is represented by the hydrophilic crosslinking of polymer-based hydrogels, formed by the reaction of one, or more monomers connected by hydrogen bonds or/and van der Waals interactions between the chains [48]. One of the most important advantages of hydrogel systems is their aqueous environment that not only protects cells, but can also be sensitive for drugs and biological agents incorporation, transport, and delivery at the injury site [47]. Moreover, they can present tailored mechanical properties, degradation profiles, and swelling abilities according to their final applications [24, 49]. Facing the traditional scaffolding strategies, hydrogel networks have also been proposed as injectable systems not only for TE strategies, but also for other clinical applications [50].
The naturally derived hydrogels have desirable biological properties as compared to synthetic hydrogels; however, they can present rapid degradation profiles for hard tissues regeneration, not to mention the chemical and molecular instability, which usually limits the reproducibility of natural-based materials [51]. Such limitations can be overcome through hydrogels formed from regenerated aqueous SF solutions, that when submitted to different physical and chemical treatments, including mechanical agitation, ultrasonication, thermal treatment, pH adjustments, organic solvents, crosslinking using ionic species (Ca2+ ions) or biological agents (enzymes), acquire a sol-gel transition (Fig. 14.2) [24, 52,53,54]. During the structural transition process, an interconnected network is formed in the aqueous solution either by the β-sheet aggregates formation (transition from random coil to β-sheet) [55] or by the crosslinking of fibroin molecules [56]. Yan et al. [50] proposed a horseradish peroxidase (HRP)-mediated crosslinking approach to produce novel SF hydrogels in a random coil conformation, that can undergo intrinsic conformational changes from amorphous to β-sheet (Fig. 14.3). These hydrogels can be adjusted in terms of gelation time, mechanical properties, and degradation profile, only by changing the SF concentration and the crosslinker (HRP/H2O2) ratio. This will allow producing different hydrogel networks according to its final application. Moreover, the enzymatic crosslinking of SF was conducted at physiological conditions, envisioning their application as injectable systems for drug delivery purposes. In a recent study [57], the potential applications of SF-based hydrogels were magnified by the production of agarose-SF sponges, processed by freeze-drying agarose-SF blended hydrogels.
The traditional approach for TE and regenerative medicine involve the culture of cells, withdrawn from the host tissue, into a pre-established structural matrix and subsequent implantation into the defect site [1]. Structures designed to mimic the natural ECM microenvironment of the replacing tissues usually involve the engineering of matrices at several levels, depending on the physical, chemical, and biological properties of the host tissue, and SF has been proven to be a suitable material to engineer different tissues [4, 12, 58]. Combining polymeric scaffolds for osteogenesis induction with a cartilaginous-like hydrogel matrix has been one of the most studied strategies for OC tissue engineering [37, 59]. Furthermore, the mechanical strength and high biocompatibility of SF make this material a rational choice as injectable fillers or as scaffolds for bone TE purposes [60, 61]. Fini et al. [60] have tested the in vitro behavior of injectable SF hydrogels, through osteoblast culture, and after in vivo implantation into critical-sized defects of rabbit distal femurs. The proposed SF hydrogels were obtained by adding citric acid into aqueous SF solution, and a commercial synthetic poly(d,l lactide-glycolide) (PLGA) copolymer was used as control material [62]. In vitro tests showed a significant increase of cell proliferation in the SF hydrogels and a higher bone remodeling capability was observed after in vivo implantation into the femoral defects, as compared to the synthetic PLGA control. Sonication-induced SF hydrogels were also proposed by Zhang et al. [63] as injectable bone replacement biomaterials (Fig. 14.4). These hydrogels were combined with the osteogenic-related growth factors, vascular endothelial growth factor (VEGF) , and bone morphogenic protein-2 (BMP-2) , and evaluated as vehicles for the encapsulation and release of biological agents. The dual factors were slowly released from the injectable SF hydrogels promoting angiogenesis and new bone formation after in vivo implantation in rabbit maxillary sinus. The authors concluded that the proposed SF hydrogels can be used as injectable matrices in a minimally invasive approach to fill and regenerate bone tissue. Moreover, the possibility of being used as vehicles to deliver multiple growth factors was also a great achievement.
Considering the high incidence of articular cartilage-related injuries [64], hydrogels are particularly desired matrices for cartilage TE, since these are water-swollen materials capable of retaining and absorbing large volumes of water and maintaining sufficient mechanical properties to support loading forces. Moreover, hydrogels have shown to be capable of encapsulating cells, biomolecules, and growth factors, for controlled drug delivery approaches after implantation in cartilage defects [65, 66]. SF-based hydrogels have also shown great potential for cartilage regeneration applications [67]. Sonication-induced SF hydrogels were proposed by Chao et al. [68] as an alternative approach to the commonly accepted agarose hydrogels [69] that yield the ability to sustain immature chondrocytes with biomechanical properties comparable to the native cartilage tissue. However, the non-degradability and lack of possibilities to modify the agarose’s structure, composition, and mechanical properties increased the author’s interest of using biocompatible, biodegradable, and highly tuned SF hydrogels to prepare cartilaginous constructs. These hydrogels presented a variety of structural and mechanical properties according to the SF extraction method, concentration, and gelation conditions. Moreover, the rapid encapsulation of chondrocytes and full maintenance of cell viability for 42 days with ECM formation (collagen and glycosaminoglycans) suggests that these hydrogels can be used as 3D models of cartilage tissue formation and maturation. In a different study, Park et al. [70] proposed novel sonication-induced SF composite hydrogels with fibrin/hyaluronic acid for nucleus pulposus cartilage formation. The authors demonstrated that the composite hydrogels allowed the chondrogenic differentiation in five different groups made of fibrin/hyaluronic acid and different SF concentrations. Importantly, the mechanical strength measurements also showed that SF induced a stronger mechanical support for cartilage tissue on composite hydrogels than fibrin/hyaluronic acid hydrogels alone. Yodmuang et al. [71] have proposed SF-based composites made by combining silk microfibers with SF hydrogels as a potential support material for cartilage TE. SF fiber-agarose hydrogel composite materials were used as control condition, showing that the 100% SF-based composites presented better and similar mechanical properties to those of native cartilage tissue (Fig. 14.5). The SF fiber reinforcement significantly influenced the mechanical and biological properties of composite materials supporting chondrocytes maturation and cartilage matrix deposition. Once again, the versatility of SF as a composite material came to overcome the limitations presented by the “gold standard” agarose-based biomaterials. The same recognition was done by Singh et al. [57] that demonstrated higher levels of cartilaginous tissue formation (glycosaminoglycans and collagen matrix deposition) within microporous hydrogels of SF blended with agarose, as compared with the microporous hydrogels of agarose used as control.
Lately, bilayered hydrogels combining different polymeric materials [36, 72], or encapsulating growth factors and/or cell populations, have also shown promising results in OC tissue repair /regeneration [73,74,75]. SF biomaterials can be particularly attractive for these strategies due to the self-assembly properties and controlled processing of the β-sheet crystalline content , which enable to modulate the degradation rate and mechanical properties of SF structures according to the target OC tissue [76]. In a recent study, a 3D bio-printing method was used to create SF-gelatin (SF-G) bioinks incorporated with human mesenchymal stem cells [77]. Enzymatic and physical crosslinking methods were applied after cell incorporation for post-printing stabilization, showing that both developed constructs supported multilineage differentiation and specific tissue formation according to the applied crosslinking method. These results provide a proof-of-concept for the fabrication of 3D heterogeneous tissue constructs using different crosslinking methods of SF-G hydrogels with different mechanical properties and biological effects, especially required for OC tissue regeneration. Moreover, the possibility of creating a printed construct for a target tissue in a patient-specific approach would be the answer for personalized OC therapy and regeneration.
4 Silk Fibroin-Based Scaffolds
Scaffolds are 3D porous matrices developed to provide a defined microenvironment that promotes tissue repair and regeneration. Ideally, scaffolds should be able to: (1) stimulate cell-biomaterial interactions, cell attachment, growth, and migration, (2) facilitate transport of mass, nutrients, and regulatory factors for cell survival, proliferation, and differentiation, (3) afford structural mechanical support, as tensile strength and elasticity, (4) degrade at a controlled rate, and (5) present minimal degree of inflammation or toxicity in vivo [78]. Further, scaffolds have desired characteristics for cell transfer into a defect site and to limit cell loss, instead of simple cells injection to the defects.
Layered scaffolds aiming bone and OC TE applications, where both underlying bone and cartilage tissues are damaged, are able to promote regeneration with specific properties and biological requirements [49, 55,56,57]. The strategy is the fabrication of stratified scaffolds consisting of separate osteogenic and chondrogenic regions, which can be manufactured in a single integrated implant, or fabricated independently and joined together with sutures or sealants. It is pretended to ensure a good compatibility between the regions by keeping the porous structure and the mechanical strength. Porosity and pore sizes, respectively, of 50–90% and 300 μm are required for an improved osteogenesis , whereas a pore size between 90 and 120 μm is recommended for chondrogenesis [79]. Several technologies have been applied to produce scaffolds with organized porosity and pore size, namely foam replica, salt-leaching/solvent casting, freeze-drying, phase separation, gas foaming, rapid prototyping, supercritical fluid technology, additive manufacturing, photolithography, microfluidics, and electrospinning [14, 26, 31, 33, 80,81,82,83,84,85,86]. These techniques also allow envisioning the encapsulation of pharmaceutical agents/drugs and cells.
Considering the unique properties of SF for biomedical applications as abovementioned, the fabrication of useful scaffold SF-based systems, as well as constructs, has been extensively investigated with very positive results, to repair and regenerate the bone, cartilage, and OC tissues [26, 31, 39, 87, 88]. Saha et al. [39] evaluated the osteo-/chondro-inductive ability of acellular mulberry and non-mulberry SF scaffolds as an implantable platform in OC therapeutics. It was shown that SF scaffolds of Antheraea mylitta (non-mulberry) were more chondro-inductive, while those of Bombyx mori (mulberry) were more osteo-inductive in similar conditions. The in vitro culture in chondro- and osteo-inductive media, showed that non-mulberry constructs seeded with human bone marrow stromal cells exhibited chondrocyte-like cells behavior up to 8 weeks of culture, whereas mulberry constructs seeded with human bone marrow stromal cells formed bone-like nodules. In vivo neomatrix formation on the scaffolds, absorbed with transforming growth factor β-3 or recombinant human BMP-2, was demonstrated after implantation for 8 weeks, in OC defects of the knee joints of Wistar rats. The neomatrix formed comprised collagen and glycosaminoglycans except in mulberry silk without growth factors, where a predominantly collagenous matrix was observed.
A different strategy was reported by Chen et al. [89] where they used SF sponge scaffolds seeded with rabbit bone marrow stromal cells (BMSC ) aiming to engineer a multilayered OC construct. BMSC-seeded scaffolds were first cultured separately in osteogenic and chondrogenic stimulation media. Then, the differentiated pieces were combined with RADA self-assembling peptides and subsequently co-cultured. It was shown that after co-culture, the GAG production in the chondrogenic region was downregulated compared with the chondrogenic control group, while the GAG production in the osteogenic region was greater than from the osteogenic control group. Furthermore, in the intermediate region of co-cultured samples, hypertrophic chondrogenic gene markers collagen type X and MMP-13 were found on both chondrogenic and osteogenic sections. However, significant differences of gene expression profile were found in distinct zones of the constructs co-cultured, and the intermediate region had significantly higher hypertrophic chondrocyte gene expression. Moreover, results showed that specific stimulation from osteogenic and chondrogenic BMSCs affected both layers inducing the formation of an OC interface. Another interesting work investigated the capability of regenerated SF/natural degummed silk fiber composite scaffolds , combined with fibrin glue, with and without autologous chondrocytes for the repair of OC lesions [90, 91]. The scaffolds showed very good mechanical properties and porosity due to the incorporation of silk fibers into the SF. In vivo biocompatibility tests of the scaffolds after implantation in OC defects of rabbit knees demonstrated good healing, regular chondrocyte arrangement, great connection to native tissues, and complete degradation 36 weeks post-implantation.
In order to improve the rapid degradation and low mechanical strength of pure chondroitin sulfate (CS), Zou et al. [92] developed a scaffold combining SF with CS, using salt-leaching, freeze-drying, and crosslinking methodologies, for cartilage tissue repair. The scaffolds exhibited a porous and interconnected structure with pores sizes of approximately 100–300 μm (Fig. 14.6I). In vitro biocompatibility tests using human articular chondrocytes (hACs) cultured on the scaffolds showed the formation of clusters inside the pores of SF scaffold, but better adhesion in SF/CS scaffold (Fig. 14.6II). After 12 weeks post-implantation in a rabbit OC defect model, the defects in SF scaffolds were repaired with fibrocartilage tissue and cartilage-like tissue, generated a thinner layer compared to the surrounding normal cartilage tissue. The defects in SF/CS scaffolds were repaired by thicker cartilage-like tissue, and well-organized subchondral bone (Fig. 14.6III). It was also observed that SF/CS scaffolds maintained better chondrocyte phenotype than SF scaffold, and silk-CS scaffolds reduced chondrocyte inflammatory response that was induced by interleukin (IL)-1β, consistent with the well-reported anti-inflammatory activities of CS.
Alternative approaches for OC TE involve SF-based structures combined with other functional materials, as the case of calcium phosphates (CaPs), can significantly enhance its biofunctionalities, and hence improved advantages of the final composites [14, 31, 87]. Yan et al. [14] prepared bilayered scaffolds for OC defects regeneration, consisting of SF and SF/CaP, respectively, for cartilage and bone regions using salt-leaching/freeze-drying techniques (Fig. 14.7A). The scaffolds showed improved micro/macrostructure able to promote cell attachment and proliferation, as well as enhanced osteoconductivity and mechanical strength by the incorporation of calcium phosphate in SF. Also, good adhesion and proliferation of rabbit bone marrow mesenchymal stromal cells (RBMSCs) cultured on the scaffolds during 7 days were observed. The scaffolds were implanted subcutaneously and in critical sizes of OC defects in the rabbit knee for 4 weeks, showing a fully integration into the host tissue with no inflammation (Fig. 14.7B). Moreover, the ingrowths of the subchondral bone in the bottom domain and the regeneration of cartilage in the surface area of the implant were observed. The quantitative results of CaP content and porosity of different regions showed much higher void space in the defect controls than in the defects with implant, with CaP content of 20% higher in SF/CaP layer than in the SF. From Fig. 14.7C, it can be observed that the connective tissues were tightly integrated in the implants, and filled the inner pores of the scaffolds, with visible vessels formed inside the scaffolds, and some fibroblasts presented in the SF/CaP layer.
Recently, in a work developed by our group, biofunctional scaffolds composed of SF and β-tricalcium phosphate, incorporating different ions, reported enhanced mechanical properties, improved cell proliferation, and higher osteogenic potential, which can also be used to engineer the bone layer in OC applications [93]. The scaffolds showed macropores highly interconnected with a size around 500 μm, and microporous structure pores with a size range of 1–10 μm (Fig. 14.8A). The biomineralized SF scaffolds, after immersion in SBF for 15 days, showed globule-like structures of apatite crystals, while the incorporation of ceramic powders into SF leads to the formation of porous spherulites-like structures (Fig. 14.8B). Interestingly, in vitro assays using hASCs presented different responses on cell proliferation/differentiation when varying the ionic agents in the biofunctional scaffolds (Fig. 14.8C). The incorporation of Zn into the scaffolds led to improved proliferation, while the Sr- and Mn-doped scaffolds presented higher osteogenic potential as demonstrated by DNA quantification and ALP activity. The combination of Sr with Zn led to an influence on cell proliferation and osteogenesis when compared with single ions.
As mentioned earlier, biomimetic OC multilayered systems, with specific microstructures and properties, have the potential clinical benefit in promoting bone and cartilage tissue repair and replacement. Taking an OC approach, Çakmak et al. [94] designed a SF-based trilayered scaffold suitable for both bone and cartilage, fabricated by salt-leaching process. For this purpose, the bone side was prepared with 4% (w/v) SF plus 5% (w/w) hydroxyapatite, the interface was obtained from 4% (w/v) SF, and for the cartilage layer were used arginine-glycine-aspartic acid-serine (RGDS)-containing peptide amphiphile hydrogels . The final mean pore size obtained for bone and OC interface layers was, respectively, 416 ± 87 μm and 194 ± 67 μm. Osteogenic and chondrogenic activity were evaluated by hBMSCs cultured in the SF scaffold in osteogenic media, while hACs were encapsulated and cultured inside the PA-RGDS in chondrogenic media, without using selective growth factors. After 2 weeks of growing separately, the bone and cartilage layers were combined with the interface layer by the soft silk scaffolds, followed by co-cultured in an OC cocktail medium. Results showed that the presence of hACs in the co-cultures significantly increases the osteogenic differentiation of hBMSCs, whereas hACs produces a significant amount of glycosaminoglycans (GAGs) for the cartilage region. Moreover, the effect of hBMSCs on chondrogenic differentiation of hACs was less effective than that of hACs on hBMSCs, and the hACs in the co-culture preserved the amount of synthesized chondrogenic ECM. Ding et al. [38] developed a trilayered scaffold combining SF/hydroxyapatite and paraffin-sphere leaching with a modified temperature gradient-guided thermal-induced phase separation (TIPS) technique . The bone layer and interface are constituted respectively by a porous and dense structure (Fig. 14.9). Live/dead tests indicated good biocompatibility for supporting the growth, proliferation, and infiltration of adipose-derived stem cells (ADSCs) in the scaffolds. Histological and immunohistochemical stainings confirmed that the ADSCs could be induced to differentiate towards chondrocytes or osteoblasts in vitro at chondral and bony layers in the presence of chondrogenic or osteogenic culture medium, respectively. Moreover, the intermediate layer could play an isolating role for preventing the cells within the chondral and bone layers from mixing with each other.
5 Concluding Remarks and Research Efforts
Many progresses have been made over the past few decades in order to fully treat and replace the damaged or non-functional OC tissues. TE is an essential approach for that, which can combine different biomaterials, bioactive molecules, and cells. Repairing OC lesions remains a formidable challenge due to the high complexity of native OC tissue, and the limited self-repair capability of cartilage. Innovative strategies , such as the ones aforementioned, present solutions for specific OC challenges, and take important roles in cell proliferation and differentiation, envisioning the formation of new tissues. Such approaches will provide the production of hybrid constructs that act as bioresorbable temporary implants and resemble the physical characteristics of the ECM . Hydrogels are defined as possessing high water content and viscoelastic nature. On the other hand, scaffolds have mechanical strength necessary to temporarily offer structural support until new tissue ingrowth.
Among all the natural biopolymers presently available, SF has shown remarkable potential for biomedical applications due to its favorable structural and mechanical properties, as well good biocompatibility, biodegradability, and thermal stability. The design of 3D structures involving SF results in biomaterials with structural integrity for self-healing and load-bearing for future applications in OC TE. Furthermore, multiphasic structures, with distinct subchondral bone and cartilage regions and well-integrated interface, can overcome the common problems experienced with monolayers scaffolds, and be an effective approach for the effective regeneration of OC tissue.
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Acknowledgments
The authors thank to the project FROnTHERA (NORTE-01-0145-FEDER-000023), supported by Norte Portugal Regional Operational Programme (NORTE 2020), under the PORTUGAL 2020 Partnership Agreement, through the European Regional Development Fund (ERDF). The financial support from the Portuguese Foundation for Science and Technology to Hierarchitech project (M-ERA-NET/0001/2014), for the fellowship grant (SFRH/BPD/113806/2015) and for the fund provided under the program Investigador for J. M. Oliveira (IF/00423/2012 and IF/01285/2015) are also greatly acknowledged.
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Ribeiro, V.P., Pina, S., Oliveira, J.M., Reis, R.L. (2018). Silk Fibroin-Based Hydrogels and Scaffolds for Osteochondral Repair and Regeneration. In: Oliveira, J., Pina, S., Reis, R., San Roman, J. (eds) Osteochondral Tissue Engineering. Advances in Experimental Medicine and Biology, vol 1058. Springer, Cham. https://doi.org/10.1007/978-3-319-76711-6_14
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