Silk fibroin scaffolds for common cartilage injuries: Possibilities for future clinical applications

European Polymer Journal 115 (2019) 251–267

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European Polymer Journal journal homepage: www.elsevier.com/locate/europolj

Silk fibroin scaffolds for common cartilage injuries: Possibilities for future clinical applications

T

⁎

Mehdi Farokhia, , Fatemeh Mottaghitalabb, Yousef Fatahic,d, Mohammad Reza Saebe,f,g, Payam Zarrintaje,f,h, Subhas C. Kundui, Ali Khademhosseinij a

National Cell Bank of Iran, Pasteur Institute of Iran, Tehran, Iran Nanotechnology Research Centre, Faculty of Pharmacy, Tehran University of Medical Sciences, Tehran, Iran c Department of Pharmaceutical Nanotechnology, Faculty of Pharmacy, Tehran University of Medical Sciences, Tehran, Iran d Universal Scientific Education and Research Network (USERN), Tehran, Iran e Departments of Resin and Addidtives, Institute for Color Science and Technolog, P.O.BOX 16765-654, Tehran, Iran f Advanced Materials Group, Iranian Color Society (ICS), P.O. Box 1591637144, Tehran, Iran g Color and Polymer Research Center (CPRC), Amirkabir University of Technology, P.O.BOX 15875-4413, Tehran, Iran h Polymer Engineering Department, Faculty of Engineering, Urmia University, Urmia, Iran i 3Bs Research Group, I3Bs – Research Institute on Biomaterials, Biodegradables and Biomimetics, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, University of Minho, AvePark, 4805-017 Barco, Guimaraes, Portugal j Department of Bioengineering, Department of Chemical and Biomolecular Engineering, Department of Radiology, California NanoSystems Institute, University of California-Los Angeles, Los Angeles, CA, United States b

A R T I C LE I N FO

A B S T R A C T

Keywords: Articular cartilage Silk fibroin Intervertebral disk Anterior cruciate ligament Meniscus Osteochondral defects

Regenerating chondral and osteochondral injuries is a main challenge in orthopedics. Some therapeutic strategies such as joint preservation operations, non-operative management, palliative surgery, and arthroplasty are the common clinical methods for repairing the cartilage defects. These treatments show often satisfactory as short-term outcomes and without clear long-term prospects. Over the past decade, the development of tissue engineering technologies offers a new therapeutic option to treat patients suffering from chondral lesions. Silk fibroin is a potent and advanced biomaterial for regenerating both soft and hard tissues. Fibroin scaffolds possess superior mechanical strength, suitable bioactivity, elasticity, degradability, and tailorable chemical structure. Due to the important properties as natural biomaterials, the fabrications of various types of scaffolds/matrices for regenerating the tissues like cartilage for regeneration and repairing the defects are possible. This review highlights the investigations on silk-based biomaterials for cartilage tissue engineering. The possibilities for future clinical application of silk fibroin based constructs in repairing intervertebral disk, anterior cruciate ligament, meniscus, and osteochondral defects are evaluated in detail.

1. Introduction Traumatic injury in weight-bearing articular cartilage (e.g., hip, ankle, and knee) causes problems that may also affect the surrounding ligaments and meniscus [1]. Sixteen to forty-six percent of patients with severe articular cartilage trauma suffer from a rupture of the anterior cruciate ligament (ACL) [2]. Furthermore, the osteochondral defects (OCD) are usually associated with subchondral bone and articular cartilage lesions as some consequences of ageing, trauma, or disease. The OCD commonly has a slow progression causing long-term disability, pain, and joint failure [3]. It is estimated that OCD will be the major cause of disability in 2030, with 35% incidence in the population

⁎

[4]. Occurrence of the OCD is related to age, sex, and geographical region. Although to treat the defects through orthopedic surgeries are progressing; however, the challenges are still remained for cartilage repair due to its structural complexities and low metabolic activities. The articular joint is first formed by undifferentiated mesenchymal stem cells (MSCs) with a complex cartilage matrix deposition by chondrocytes. The cartilage matrix consists of collagen type II (> 60%), proteoglycans (25%), and glycoproteins (15%) [5]. Due to the avascular nature of cartilage matrix, MSCs cannot be transported from blood vessels to injured cartilage, thus reducing the self-healing ability. Articular cartilage also face limited cellular turnover because the cells are entrapped in a dense extracellular matrix (ECM) formed by

Corresponding author. E-mail address: [email protected] (M. Farokhi).

https://doi.org/10.1016/j.eurpolymj.2019.03.035 Received 23 February 2019; Received in revised form 19 March 2019; Accepted 20 March 2019 Available online 20 March 2019 0014-3057/ © 2019 Elsevier Ltd. All rights reserved.

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highly expressed in all regions of the cartilage [25,26]. A similar pathway was identified for Smad-2 and Smad-3 in mediating TGFβ signaling [21]. Other signal transduction pathways like independent non-canonical Wnt and β-catenin-dependent canonical signaling pathways also play crucial roles in growth, development, and maintenance of chondrocytes during cartilage regeneration. β-catenin-dependent signaling induces endochondral ossification and axial growth. Overexpression of this signaling pathway could hinder the early formation of cartilage and the organization and functionalization of growth plate in mice. While non-canonical Wnt signaling triggers the columnar organization of growth plates by chondrocytes [20]. This signaling pathway is also related to other signals such as epidermal fibroblast growth factor receptor (EGFR) pathways [27], BMP/TGFβ [28], hedgehog [29], and retinoid [30]. These pathways play critical roles in cartilage and skeleton development. Activation of FGF-2 is another pathway that stimulates the proliferation of chondrocytes and prepares enough chondro-progenitor cells for differentiation. However, the exact mechanism of FGF-2 activity is still unclear [31,32]. It seems that overexpression of TGFβ1 as a result of FGF-2 activation promotes the proliferation of chondrocytes [33]. Furthermore, downregulation of TGFβ2 by the FGF-2 cascade might result in more differentiation of chondrocytes [34].

chondrocytes. These properties are responsible for low regeneration capacity of cartilage [6]. Various strategies have been performed to promote cartilage repair; however, none of them had optimal outcomes. Surgeons mainly focus on pain relief from the clinical symptoms rather than stimulating cartilage tissue repair [7]. Pridie’s resurfacing surgery technique is the first approach that helps to restore the surface of normal articular cartilage [8]. In this procedure, the sub-chondral bone is disturbed and the bleeding from the bone marrow triggers the healing mechanisms in the injured site of the cartilage. The current method is based on the biological capacity of articular cartilage healing to reduce patient symptoms [9]. Furthermore, cartilage tissue engineering pursues to design new therapeutic strategies. In this approach, the proliferation and differentiation of chondrocytes at the site of injury are promoted by using suitable biomaterial scaffolds incorporating genes, growth factors, and drugs [10]. Proper engineered cartilage should interface with the adjacent cartilage tissue and subchondral bone. The mechanical behavior of the engineered cartilage should also be similar to the adjacent tissue to maintain the functionality of the joint structure [11]. Many natural and synthetic polymers are used for fabricating artificial cartilage with appropriate biocompatibility, biodegradability, and structural characteristics. In recent years, the efforts focus more on using natural polymers because of their useful properties in terms of versatility, biocompatibility, functionalization, and biodegradability. Many natural polymers have some motifs, which act as ligands for cell surface receptors [12]. The unique properties of silk fibroin (SF) including tunable biodegradability, biocompatibility, low immunogenicity, outstanding structural integrity, and high mechanical strength make it a potentially useful natural polymer for tissue engineering [13,14] and drug delivery [15–17] applications. Recently, Cheng et al. published a general overview of using SF based scaffolds for cartilage tissue engineering [18]. A systematic review article also describes the application of Bombyx mori (B. mori) SF for regeneration of cartilage tissue [19]. Based on our knowledge, there is no comprehensive review article with focusing on the potential of SF-based constructs for repairing cartilage diseases. Recently, many researches have focused on the applications of SF-based constructs for common cartilage problems like intervertebral disc degeneration, cruciate ligament rupture, meniscus tear, and osteochondral defects. Therefore, we decide to collect all the recent in vitro or in vivo findings for these issues in a review article.

3. Aspects of molecular mechanisms of cartilage repair In adults, articular cartilage is exposed to different pathogenic elements e.g., mechanical fatigue, and inflammation that lead to matrix degradation and chondrocyte death. However, small, acute, and full thickness joint injuries have some repair capacity in humans and animals, especially in young [35,36]. Any fracture in articular cartilage results in post-traumatic arthritis without any clear mechanism. Generally, the trauma increases the level of systemic and local pro-inflammatory components such as interleukin 6 (IL-6) and IL-8 [37–39]. But no fibrin clot formation, angiogenesis, and recruitment of inflammatory cells are observed in cartilage defects, unlike that observed during wound injury [40]. After injury, some antiangiogenic molecules like secreted protein acidic and rich in cysteine (SPARC), chondromodulin-1, thrombospondin-1, type XVIII derived endostatin, and collagen type II derived N-terminal pro-peptide (PIIBNP) inhibit angiogenesis and thus restrict cartilage regeneration [41]. Conversely, the BMP groups, especially BMP-7 is known to stimulate matrix synthesis and cartilage regeneration. So, BMP-7 plays important roles during both cartilage and bone repair [42]. Similarly, BMP-2 has a significant role in cartilage regeneration, with the ability to enhance the expression of aggrecan and collagen type II in IL-1-induced cartilage injury in a mice model. Inhibiting the function of BMP-2 decreased the production of proteoglycans [43] and thus enhance cartilage damage [44].

2. Cell signaling and cartilage formation For optimal cartilage repair, it is important to activate key signaling pathways and biological cues involved in cartilage tissue interaction and regeneration. A variety of local and systemic factors have been introduced for cartilage repair including transforming growth factor β (TGFβ), Wnts, Indian hedgehog, bone morphogenetic protein (BMP) superfamily, thyroid hormone, parathyroid hormone-related protein (PTHrP), platelet derived factors (PDGFs), insulin like growth factors (IGFs), fibroblast growth factors (FGFs), and different vitamins [20]. Among them, TGFβs are the key regulator that controls the proliferation and differentiation of chondrocytes [21]. During cartilage synthesis, MSCs migrate and localize in a suitable region or niche and form condensed blueprints for future bone structures [22,23]. During chondrogenesis, pre-chondrocytes secrete specific ECM components like collagen type II and undergo changes in morphology from fibroblastlike to round shape [23]. TGFβs have many isoforms, which are involved in MSCs condensation. TGFβ prompts the expression of RY-box containing gene 9 (Sox-9) that is responsible for collagen type II and aggrecan formation in the early condensation phase of MSCs [24]. This biochemical also stimulates the Smad signaling pathway in the targeted cell. Smad-2 and Smad-3 are overexpressed in the growth plate. Generally, Smad-2 is more highly expressed in proliferative and pre-hypertrophic cells, whereas Smad-3 is highly expressed in pre-hypertrophic and hypertrophic chondrocytes. In contrast, Smad-4 showed

4. Current advanced therapies for cartilage regeneration Pridie's resurfacing technique was the first approach that helped to restore the surface of normal articular cartilage [8,45]. In this procedure, the sub-chondral bone is disturbed and the bleeding from the bone marrow is stimulated to start the healing mechanisms in the injured site of the cartilage. The current method is based on the biological capacity of articular cartilage healing with an ability to reduce patient symptoms [9]. Since this development, various techniques have been used for cartilage regeneration as summarized in Table 1. 5. Tissue engineering based strategies for cartilage repair Cartilage tissue engineering uses various scaffolds and offers many advantages compared with scaffold-free approaches, including better control of filling the injured cartilage site, less complications at donor sites, fewer challenges during implantation, and a shorter-time recovery. Moreover, generating a 3D structure supports chondrocytes to 252

European Polymer Journal 115 (2019) 251–267

[53]

[54]

Risk of graft fracture during harvest, Femoropatellar pain, graft cartilage necrosis, Pseudarthrosis in graft

Not enough data to indicate the effect of lavage in inducing the repair and activity of defected cartilage Inducing osteonecrosis and chondrolysis, generating reactive synovitis, promoting articular cartilage degeneration

reduce dedifferentiation and to produce higher content hyaline-like cartilage [56,57]. A successful tissue engineered construct for cartilage is able to create the liquid and solid phases of the connective tissue, biomimetic zonal and regional cartilage structure, and integrate with the surrounding native tissues [58,59]. Therefore, it is important to select the appropriate biomaterials that improve regeneration of defect cartilage. To date, various synthetic and natural polymers have been developed for cartilage tissue engineering. For designing an appropriate scaffold for optimal tissue repair, it is necessary to consider the structure and function of the normal tissue. Collagen type II and glycosaminoglycan (GAG) have important effects toward preserving chondrocytic phenotype and stimulating chondrogenesis in vitro and in vivo [60]. Otherwise, chondrocytes can de-differentiate or create fibrocartilaginous matrices enriched by collagen type I that lead to failure in the formation of hyaline cartilage [61]. Different types of collagen are extracted from animal specimens and depending on type and purity can induce some inflammatory responses in the host [62]. In order to overcome this limitation, some studies have used recombinant human type II collagen [63]. Monomeric type-I and type-II collagen scaffolds that bypassed the immunogenic responses associated with fibrillar collagens were used to prevent OCD [64]. Table 2 and 3 summarize reports using synthetic or natural polymers for cartilage tissue engineering. 6. Characteristics of silk protein: structure, crystallinity, and mechanical properties

Mild cutting instruments only, no effect in stimulating cartilage regeneration

Silk proteins are generated by various species of silkworms and spiders. Different silkworms produce silk proteins in fiber forms with various properties in terms of structures and functions. Some nonmulberry silkworms such as Chinese oak tasar silk of Antheraea pernyi or Indian tropical tasar silk of Antheraea mylitta contain the arginine-glycine-aspartic acid (RGD) motif that plays an important role in cell adhesion [82]. Additionally, the presence of higher amounts of basic (arginine), acidic (aspartic acid), and polar (serine) amino acids in the structure of SF extracted from mulberry silkworms (B. mori, African wild silk moth Gonometa postica, and G. rufobrunnea) increase the opportunity for surface modification and consequently the attachment of cells [83]. The SF fibers isolated from B. mori silkworms have a 10–25 mm diameter with light (∼26 kDa) and heavy (∼390 kDa) protein chains and a glycoprotein P25 (∼30 kDa) [13]. A disulfide bond connects the heavy and light subunits, while P25 is linked by hydrophobic interactions. The heavy chains contain hydrophobic Gly-X (X being Ala, Ser, Thr, Val) domains that form anti-parallel β-sheets. The hydrophilicity of the light chain makes it moderately elastic [84,85]. Structurally, SF consists of crystalline and amorphous domains. The crystalline domains contain Gly-Ala repeats interspersed via tyrosine and serine amino acid containing motifs. Bulky side chain amino acids like aspartic acid constitute the amorphous domains. Three conformations including silk I, II, and III are considered for crystalline structures of SF from B. mori. Silk I is the water-soluble form of SF with a high α-helix content that can easily be transformed into silk II with an anti-parallel β-sheet structure [86–88]. The flexibility and mechanical strength of SF fibers come from the unique structural conformation [89]. It was previously reported that degradation of silk fibroin during 12 weeks in phosphate buffer saline could transition from β-sheet to random coil that decreased the Young’s modulus and adhesion force of the construct [90]. Generally, the tensile strength, breaking elongation, and toughness of SF fibers are estimated to be 0.5 GPa, 15–62%, and 104 J kg−1, respectively [91]. Various methods have been used to stimulate the β-sheet crystallization in silk-based scaffolds. Methanol and ethanol are two common inorganic solvents that promote β-sheet formation ∼36% and 55%, respectively [92–94]. However, the highest βsheet crystallinity (∼60%) is induced by steam autoclaving in high pressure and temperature [95]. Water annealing in room temperature is another method [96] that induced β-sheet (30%) formation mixed with

Laser ablation/Laser Chondroplasty

Closed-needle-hole technique that irrigates or lavages the joint with sodium chloride, Ringer and lactate Use of laser to remove debris Joint lavage

Experimental method for treating painful joint like structural lesions

Using cylindrical osteochondral autografts from low-weight bearing areas of knee for injured area Mosaicplasty

A two-step procedures that include biopsy preparation from articular cartilage (first stage) followed by cell expansion and implantation in the defected site (second stage) Using mature hyaline cartilage to cover the surface of focal chondral injuries Autologous chondrocyte implantation (ACI)

Osteochondral allograft transplantation (OAT)

Fibrin clot formation by inducing the release of bone marrow products like growth factors, stem cells, and platelets that remodel to fibrocartilage Microfracture

Using own tissues from the individuals that cause less infection during transmission compared with allografts Simple, one-step surgical procedure, ability to transfer the living autologous cartilage, in-expensive with low morbidity Useful for individuals suffering from osteoarthritic or lesions with painful joints, e.g., knee Tissue coagulation in aqueous environment unlike electrothermal cutting instruments that produce temperatures > 250 °C

[54,55]

[49–52]

Treatment of lesions with more than 10 cm2 diameter by returning the hyaline-like cartilage

Applicable for defects that less responded to debridement or reparative strategies, most useful for knee and less for humeral head and talus Useful for femoral condyle transplants, ankle, shoulder, and elbow. Best approach for defects with size 1–4 cm2 Applicable for patella, tibia, and femoral condyle

High risk of donor site morbidity and multiple plugs in lesions with size more than 8–9 cm2

[46–48]

Fracture of subchondral bridges, high growth rate of bony structures, insufficient marrow products in incomplete microfracture, hypertrophic overgrowth, fragmentation of fibrocartilagesubchondral bone junctions The need for two-step procedures, long recovery time, slow tissue maturation by implanted chondrocytes Simple, cost-effective, low rate of morbidity, and short-term clinical outcomes in 67–80% of patients Appropriate for treating full-thickness chondral defects with small diameter (≤2-3 cm2), infrequent treatment of large defects (≥2.5–3 cm2) in older individuals

[46,47,49]

Refs. Disadvantages Advantages Applications Description Technique

Table 1 Current strategies for treating cartilage defects in the clinic.

M. Farokhi, et al.

253

Selective laser sintering

–

Molding

PCL6

PEG7

UHMWPE9/ PGA10

254

3D porous structure 3D structure

3D porous structure

3D porous structure Hydrogel

3D porous structure 3D porous structure 3D porous structure

Scaffold type

2

Polylactic acid. Biomaterial-associated multinucleated giant cells. 3 Polyglycolic acid. 4 Adipose derived stem cells. 5 Poly(lactic-co-glycolic acid). 6 Polycaprolactone. 7 Polyethylene glycol. 8 Human bone mesenchymal stem cells. 9 Ultra High Molecular Weight Polyethylene. 10 Polyglycolic acid. 11 Poly (vinyl alcohol). 12 Polyurethane. 13 Mesenchymal stem cells. 14 Transforming growth factor beta-3.

1

3D printing

Solvent casting/ particulate leaching

PLGA5

PU12

Molding

PGA3/PLA

Salt-leaching

3D printing

PLA1/bioglass

PVA11/PLGA

Processing methodology

Material

MSCs13

Chondrocytes

BMSCs

hBMSCs8

Chondrocyte

–

Significantly enhanced wear resistance of scaffolds after in vitro cell culture Stimulating self-aggregation of MSCs by inducing influx of TGF-β3 14 or small molecule drug Y27632, promoting MSCs differentiation and cartilage matrix production

–

Increasing the hydrophilicity, mechanical behavior, and water uptake of PCL by surface functionalization using collagen and gelatin Stimulating chondrogenesis by hBMSCs using glucosamine-modified PEG hydrogels specially at 5 mM and 10 mM concentrations

Structural integrity of scaffolds, morphology changes with decreased mechanical strength with progressive degradation Mesh or wrap distribution of PLA in the scaffold under the influence of PGA fibers confirmed by SEM –

– ASCs4

In vitro key findings

Seeded cell

Table 2 Synthetic polymers applicable for cartilage tissue engineering.

High performance of 3D printed scaffolds in inducing cartilage repair in rabbit model

Formation of cartilage tissue eight weeks after subcutaneous implantation of 10 mM glucosamine-modified hydrogels seeded with hBMSCs, downregulation of fibrosis and hypertrophy by increasing concentration of glucosamine in scaffolds Formation of mature cartilage around the implanted scaffold after eight weeks of implantation, growth of the new cartilage within the inter-connective pore of the scaffold –

Higher vascularization and BMGC2 content after incorporation of bioglass to PLA scaffolds Well integrated newly formed cartilage with adjacent tissue and subchondral bone three months after implantation Formation of a new hyaline cartilage enriched with collagen type II and glycosaminoglycan, integration with subchondral bone, moderate inflammatory response Inducing new cartilage formation eight weeks post-implantation

In vivo key findings

[72]

[71]

[70]

[69]

[68]

[67]

[66]

[65]

Ref.

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European Polymer Journal 115 (2019) 251–267

Nanofibers Hydrogel

Hydrogel

Hydrogel

Hydrogel

Hydrogel

Hydrogel

Hydrogel

–

–

–

–

–

–

–

Fibrin

Agarose

Hydroxypropyl-methyl cellulose

Alginate

Chitosan

Elastin

HA

255

2

Bone marrow-derived mesenchymal stem cells. Stromal cell-derived factor-1. 3 Infrapatellar fat pad. 4 Platelet-rich plasma. 5 Human adipose stromal cells. 6 Periodontal ligament stem cells. 7 Gingival mesenchymal stem cells. 8 Human bone marrow mesenchymal stem cells. 9 Demineralized bone matrix. 10 Hyaluronic acid. 11 Sulfated-glycosaminoglycan. 12 Mesenchymal stem cell.

1

Gelatin

Oriented and random structure

Temperature gradient-guided thermal-induced phaseseparation Electrospinning

Collagen

Scaffold type

Processing methodology

Materials

Table 3 Natural polymers applicable for cartilage tissue engineering.

MSCs12

Chondrocyte

Chondrocytes

PDLSCs6, GMSCs7, hBMMSCs8

hASC5

Chondrocyte

IFP3

Chondrocyte

BMSCs

1

Seeded cell

Increasing HA10 concentration enhanced expression of cartilage gene markers, improved sGAG11 deposition, and reduced unwanted fibrocartilage phenotype HA hydrogel (1% (w/v)) offers permissive microenvironment for encapsulation of MSCs at a high density (60,000,000 ml−1) with acceptable mechanical properties (∼1 MPa)

Higher spreading of cells on DBM9 incorporated chitosan scaffolds

Significantly higher rate of chondrocyte-related gene expression, glycosaminoglycan, and collagen contents in PRP4–agarose gel compared with pure agarose gel after 28 days Improved mechanical strength by adding laponites (nanoreinforcement clay) to cellulose hydrogels without interfering with diffusion of O2 and viability of cells after gel formation Expression of collagen II and Sox-9 after 4 weeks postdifferentiation

Potential of electrospun nanofibers modified by gelatin-tyrosine to increase proliferation, spreading, and attachment of chondrocytes Functionalize fibrin hydrogels with 2% ECM particles without disturbing the stability of the hydrogel

Good mechanical properties of oriented collagen scaffolds than random samples, supporting cell proliferation

In vitro key findings

–

Successful formation of new cartilage tissue containing ECM components like collagen and glycosaminoglycan using composite scaffolds Appearance of ectopic cartilage repair inside and around the implanted area, higher chondrogenesis of PDLSCs compared with GMSCs and hBMMSCs Successful repair of defect cartilage by implantation of scaffolds after 24 weeks, no inflammation after implanting allogenic chondrocytes –

Ability of freshly isolated stromal cells to induce cartilage repair when loaded on fibrin hydrogels functionalized with ECM particles –

[81]

[80]

[79]

[78]

[77]

[76]

[75]

[74]

–

Ref. [73]

2

Oriented scaffolds conjugated with SDF-1 repair of cartilage defects in rabbit model

In vivo key findings

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markers displayed the potential of AF-like tissue in inducing the differentiation of hMSCs into chondrogenic lineage. The scaffold had also high compressive modulus (499.18 ± 86.45 kPa) because the concentric rings of lamellar structure surrounded the gelatin core NP substitute [108]. This was comparable to the compressive modulus of the native human AF tissue (380 ± 160 kPa) [109]. In another study, three sources of SF proteins including B. mori silk as a mulberry silk and non-mulberry silk e.g., Antheraea assamensis, and Philosamia ricini, were used to fabricate full-thickness disc-like angle-ply AF lamellar sheets [110]. The porcine AF cells were proliferated along the direction of the lamellar pores. In Antheraea assamensis SF constructs, the mechanical properties of the scaffold increased due to high deposition of COLIA1, while the amount of sulfated glycosaminoglycan (sGAG) decreased at the same time. This phenomenon may be related to the specific response of a cellular population to the stiffness of a particular substrate. Furthermore, the mechanical strength of the lamellar layer gradually increased from the inner to outer layers. The cells existed in the outer layer had a fibroblast-like phenotype and they were oriented along the axis; while the cells in the inner layer had chondrocyte-like morphology with more expressed cartilage markers like aggrecan and collagen type II (COL2A1) [110]. To prepare advanced biomaterials with more regeneration capacity, some of the bioactive biomolecules can be incorporated into the bulk materials to mimic and activate the native process of tissue repair. For this, chondroitin sulfate (CS) was incorporated into multi-lamellar silk scaffolds to promote the regeneration of AF region in disc tissue. This construct was comparable to the hierarchal structure of AF region and provided suitable substrate for cell growth, attachment, and matrix production in direction of silk fibers. The high expression of RhoA was observed in CS conjugated SF construct confirming positive control over the chondrocyte specific genes expression through highly enhanced gene expressions of collagen II, biglycan, and Sox-9. Sox-9 can affect the expression of collagen type II, which is generally decreased by degeneration and aging of disc tissue. Thus, incorporating CS into SF substrate helped to synthesis Sox-9 and consequently produced collagen type II which stimulated better regeneration of the defected disc tissue [111]. Other study showed that adding CS to oriented lamellar silk fibrous structures had no significant effect on producing collagen type II and GAG. This increased the stiffness of the silk fibers and made it suitable for AF tissue repair. CS has also a crucial role in morphogenesis of IVDs because it is highly distributed during different phases of embryo development, postnatal growth, and aging steps. CS is linked to GAG chains in the ECM of AF tissue [112]. Silk scaffolds containing genipin-enhanced fibrin hydrogels showed suitable characteristics for AF tissue regeneration [113]. The hydrogels tolerated the applied forces and maintained the silk fleece-membrane within the implanted region for two weeks. No herniation was also detected after applying loading regimes. A decreased in GAG content was observed in the defected discs, which were under static loading indicating an irreversible recovery after inducing the invasive defect [113]. It was also reported that SF functionalized with tripeptide Arg-Gly-Asp (RGD) peptide did not promote the AF cell attachment and tissue formation. The enhanced expression of aggrecan and collagen type II were compared with the cells seeded on the unmodified constructs [114]. Rather than engineering AF tissue, it is still challenging to regenerate NP part of IVD by tissue engineering strategies. It is necessary to develop appropriate structures for NP repair. Accordingly, highly porous SF scaffolds containing interconnected macro-pores were prepared by combining freeze-drying and paraffin-sphere-leaching methods [115]. The porosity and pore size of the scaffolds were 92.38 ± 5.12% and 165.00 ± 8.25 μm, respectively. The scaffolds also had suitable mechanical strength and interconnected pores for cartilage studies. After 3 weeks of cell culture, the scaffolds had a compressive elastic modulus of 0.10 MPa which was comparable to the compressive modulus of the native NP (0.003–0.31 MPa) [116]. Six weeks after implantation, large number of cells and ECM components

a high amount of silk I helical structure [97]. To fully control the structure and characteristics of silk, a single systemic process was used through temperature-controlled water vapor annealing from 4 °C to 100 °C. By using this method, various types of silk conformations including silk I, and silk II with controllable β-sheet can be produced that can also regulate the biological response, mechanical strength, and degradability [92]. In different studies, heat-treatment of silk fibers and films were studied [98–100]. The results showed that by increasing the temperature, color, mechanical strength, and crystallinity of the silk were changed. It was also reported that heating the casted silk films up to 140 °C, the conformation of silk was highly transformed to β-sheet which is common in B. mori cocoon fibers. Moreover, heating the oriented silk fibers over transition temperature, induced the β-sheet conformation over α-helical structure [101]. Therefore, by changing the temperature, it is possible to control the structure and mechanical strength of the silk-based substrates. 7. Current strategies to promote intervertebral disc regeneration using silk fibroin-based scaffolds Degeneration of intervertebral disk (IVD) is a multifactorial disease that causes many complications for patients. Millions of people suffer from degenerative disk disease with lower back pain around the world. Many parameters affect the initiation and progression of degenerative disk disease including aging, malnutrition, mechanical damage, and genetic disorders; however, the pathogenesis of the disease is not fully understood [102]. IVD is comprised of nucleus pulposus (NP) and annulus fibrosis (AF), which contain different cellular populations and biomolecules with diverse mechanical properties. The AF comprises 15–25 concentric layers reinforced by collagen fibers that are aligned with transverse plane of the disc at an angle of ∼30° [103]. NP segment are consist of collagen type II, negatively charged proteoglycans, and high water content [104]. In situ induction of damaged IVD for regeneration is preferable; however, some artificial IVDs are fabricated in vitro and consequently implanted in an animal model [105]. Furthermore, the tissue-engineered IVD constructs are beneficial when compared with arthroplasty or prosthetic discs in terms of biomechanical similarities to the defected disks. Various natural and synthetic biomaterials are used to mimic the anatomical structure of AF. Accordingly, using lamellar scaffolds for engineering AF structures have reputation because they can mimic the hierarchical structure of AF, while providing comparable mechanical properties to the native AF. Recently, many investigations are focused on the engineered IVD based on SF scaffolds. Fig. 1 illustrates the use of SF scaffolds for promoting IVD repair. A lamellar SF scaffold was prepared by freeze-drying to mimic the lamellar units of AF tissue [106]. The inter-lamellar spaces and the average pore sizes of lamellar scaffolds were 150–250 µm and 100–250 µm, respectively. The proliferation and attachment of the porcine AF cells were induced by seeding on the lamellar SF scaffolds leading to AF-like tissue formation. The cells also covered the walls of the scaffolds and penetrated into the pores after 2 weeks. In contrast, the cells only covered the surface of non-lamellar SF scaffolds as a control group. Moreover, the lamellar scaffolds had non-significant weaker ultimate tensile strength (UTS) and linear elastic modulus than non-lamellar ones after 2 weeks [106]. The mechanical stability of AF is an important issue because IVDs are exposed to fatigue failure particularly in high bending movement. AF failure or disturbing its collagenous network are the main causes of disk herniation [107]. Silk-based disc-like angle-ply scaffolds containing concentric lamellar sheets were also applied to mimic the complicated structure of AF [108]. The direction of lamellar layers exhibited the alignment of native AF layers to prepare angle-ply scaffold. The seeding of porcine AF cells on the lamellar structures increased the expressions of collagen Iα 1 (COLIA1), Sox-9 (SRY box-containing), and aggrecan to ∼11, ∼5.1, and ∼6.7-fold, respectively. The high-level production of these 256

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Fig. 1. Artificial intervertebral disk based on silk fibroin as a candidate for the treatment of degenerative disk disease.

Fig. 2. Histological assessment for evaluating the deposition of collagen type II and proteoglycans. (A) hematoxylin and eosin (H&E) (A,B), safranin O (C,D), and immunocytochemical staining (E,F) at 3 and 6 weeks. Copyright © 2014 Elsevier B.V. Reprinted with permission from Zeng et al. [115].

recovering the hydrogel during load bearing. Implantation of hydrogels in rabbit animal model showed high biocompatibility and stability in vivo over a three-month [117]. The blended silk with polyvinyl alcohol cryogel was also introduced as a replacement for NP. This cryogel enhanced cell proliferation and good physical features in terms of compressive modulus and water content which were useful for NP replacements [118]. The mentioned studies were focused on the application of AF or NP in degenerative disk disease, separately; however, artificial discs should have AF and NP parts, concomitantly. For this, a silicone NP construct

like collagen type II and proteoglycans were deposited within the pores of the scaffolds, which confirmed the results obtained from the in vitro study (Fig. 2) [115]. Using injectable hydrogels is a less-invasive strategy for repairing NP defects, but it seems challenging for optimization. It was reported that liquid silk fibroin/polyurethane (SF/PU) hydrogels in an injectable form could be a suitable choice for NP structure [117]. The hydrogels showed high modulus under confined compression. It was decreased when applying unconfined compressions. Moreover, the hydrogels exhibited acceptable fatigue resistance as an important factor for 257

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the in vivo repair of ligament in an early stage of ACL regeneration [127]. After 2 weeks, acceptable rate of well-perfused cells and vessels found in ACL tissue grafted by ETEL, which reached to lower regeneration during 4–12 weeks. During this time, higher mature ligament and more regenerated ligament-bone interfaces observed in animals treated with ETEL compared with control group. Compared to control group, the modulus and stiffens of ETEL treated group were significantly higher after 12 weeks post-operation [127]. In another study, roped shape SF/poly(lactic-co-glycolic acid) (PLGA) constructs were used to repair ACL injuries [128]. After 26 weeks, a direct insertion site-like structure was formed in the ligament-bone junction in defected ACL region after implanting scaffold-BMSC constructs in a rabbit model. Re-construction of ACL tissue was confirmed by expression of collagen type I and II and tenascin-C. Scaffold-BMSC also had higher tensile strength (101.2 ± 13.4 N) than bulk scaffolds without BMSC on the new ligament-bone interface region [128]. Fan et al. fabricated the structure of ligament ECM by combining microporous silk sponges into knitted silk meshes seeded with MSCs [129]. After 24 h implantation in rabbit model, MSCs with fibroblast-like morphology were dispersed in the repaired ACL. Moreover, the major ECM components of ligaments e.g., collagen type I, III, and tenascin-C were produced. The host knee joint may provide an autologous bioreactor by secreting numerous cytokines and physiologically relevant cyclic loading to improve the differentiation capacity of MSCs towards chondrogenic cells [129]. In another study, the long-term effects of knitted silk-collagen sponges on repairing the ACL was evaluated in a rabbit model [130]. Two months post-implantation, the scaffolds improved the spreading and attachment of spindle-shaped cells compared with pure silk scaffolds without collagen. The collagen phase of the scaffolds maintained the internal space within the micro-porous knitted silk sponge [131]. This structure supported the tissue in-growth, proliferation, and migration of MSCs from the tibial plateau/femoral condyle, and also fibroblasts/synovial cells from the joint cavity [132–134]. Six months post-transplantation, the animals with knitted silk-collagen grafts showed high expression levels of ligament-relevant genes. Eighteen months after, the mature structures of ligaments and direct ligament-tobone repair were also observed [130]. In another study, silk-based scaffold synthesized by rolling the micro-porous silk mesh around braided silk cord [135]. The scaffolds supported the growth and differentiation of MSCs into fibroblast. Then, the scaffolds were implanted in a knee joint of a pig model due to its similarity to the human knee joint. After 24 weeks post-implantation of MSC seeded scaffolds, ECM production and proliferation of fibroblastlike cells was noted (Fig. 4). The tensile strength of the regenerated tissue provided a suitable mechanical requirement for daily activities [135].

wrapped with SF scaffold containing bone marrow stromal cells (BMSCs) was investigated to fabricate natural IVD-like structure [119]. The seeded BMSCs survived at least for 4 weeks on the constructs and produced GAGs. The ratio between collagen types I and II changed after 4 weeks. First, expression of collagen type I dominated, and then collagen type II increased after 4 weeks [119]. These data suggested that extensive ECM remodeling happened during this period with a resulting matrix composition similar to the inner AF. The increase in collagen type II is a sign of chondrogenic differentiation of BMSCs. In another study, a biphasic composite based on silk and fibrin/hyaluronic acid gels was prepared to mimic the structure of AF and NP, respectively [120]. The results confirmed that biphasic scaffolds supported more cell growth and enhanced specific expression level of chondrocyte markers in comparison to the cells proliferated on porous constructs with the fibrin/hyaluronic acid gel. Therefore, the biphasic scaffold was useful to construct both AF and NP segments of IVD in vitro [120]. The aim of using bio-engineered scaffolds for disk repairs is to restore the morphological and functional features of the native tissue. A tissue-engineered artificial IVD should be biocompatible with moderate porosity, which should also provide similar morphological features to the native tissue with suitable structural integrity and mechanical strength. It seems that the various structures based on SF polymers have excellent characteristics to fabricate AF and/or NP scaffolds with the hope to apply for treatment of IVD related diseases in human in near future. 8. Management of anterior cruciate ligament regeneration using silk fibroin based scaffolds Anterior cruciate ligament (ACL) failure is a frequent problem that affects the normal function of knee joints. In the United States, about one in 3000 persons suffer from ACL failure [121] and annually more than 100,000 ACL reconstructions are performed [122]. To date, there is no clinically effective treatment for this issue due to low regeneration capacity of ACL. The tissue engineering strategies might be helpful. Artificial ligaments are the one alternative for traditional treatment of ACL. Collagen is the main components of the ligaments and collagenbased scaffolds had been used for ACL reconstruction. However, the low mechanical strength of collagen restricts its therapeutic applications [123]. Furthermore, some artificial ligaments are commercially available and are used in clinic e.g., Leeds-Keio®, Proplast®, Gore-Tex®, and Stryker-Dacron® [124]. Ligament advanced reinforcement system (LARS) comprising polyethylene terephthalate (PET) is another artificial ligament. However, the presence of PET in LARS, as a hydrophobic structure, limits the ingrowth of autologous tissue. SF alone (Fig. 3) or incorporated with PET can be better structured for fabrication of ACL because they have light weight (1.3 g/cm3), possess high mechanical strength (> 4.8 GPa), and good elasticity (> 35%) [125]. A study showed that silk could increase the ligamentization of artificial PET ligament in a canine anterior cruciate ligament reconstruction model [126]. The hybrid suspensory ligament made from silk/PET had silk in the weft yarn and PET in the warp yarn, while PET suspensory ligament had PET in weft and warp yarn, concomitantly. The hybrid silk/PET suspensory ligament was mechanically enough resistant to regenerate the ACL in a dog model. Collagen deposition and autologous tissue in-growth were higher in animals implanted with silk/PET suspensory ligaments than those made of PET suspensory ligaments, which indicated the potentiality of silk in inducing PET ligamentization [126]. Vascularization is an essential process for optimal tissue regeneration specifically for repairing massive and avascular tissues such as ACL. Recently, it is interested to induce angiogenesis by using engineered scaffolds in the defected site. To promote angiogenesis, many researchers focus on using exogenous vascular cells or angiogenic factors. In a study, the vascularized ectopic tissue engineered ligament (ETEL) based on silk/collagen scaffolds were prepared for mimicking

9. Application of silk fibroin for treatment of meniscal rapture Annually more than 400,000 meniscus surgeries take place in Europe and over 1 million in the United States. This usually occurrs through sports and daily activities [136]. More than 37% of sport injuries are also related to knee-joint defects. Another main knee injury (∼24% incidence) is lesions in medial meniscus that affect more than 80% of individuals. The absence of blood vessels in the fibrocartilaginous tissue of meniscus limit the regeneration capacity of this tissue [137]. Menaflex® Collagen Meniscus Implant (CMI) and Actifit® polyurethane meniscal scaffold are two commercial artificial meniscus. They are used clinically, without wide application for all cases [138,139]. Artificial silk based meniscus is examined in different studies indicating proper characteristics in terms of mechanical strength, structural, and biological properties for meniscus repair (Fig. 5) [140,141]. The friction coefficient of silk substitutes against cartilage was 0.056, which was more than native meniscus. It was still acceptable for meniscus replacements [142]. 258

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Fig. 3. Silk scaffold as substitutes for a ruptured anterior cruciate ligament.

model [149]. Based on the macroscopic assessment, no observable immunogenic response was detected around the implanted region and most scaffolds were located in the defect site. Though, the remaining meniscus rims had no solid connection with the implanted scaffolds and some of them also ruptured at the middle part [149]. Totally, meniscus tissue engineering is in its infancy. This needs further investigation for developing suitable scaffolds that mimic the native tissue. A useful scaffold should have similar mechanical characteristics in terms of compressive modulus (75–150 kPa), and tensile modulus (75–150 MPa) to the native meniscus [150,151]. The degradation rate of the scaffold should be extended to 9–12 months for optimal regeneration [152]. FibroFix™ is a silk-based scaffold mimicking the cartilage tissue with suitable compressive modulus comparable to native meniscus [153]. However, to enhance the regeneration capacity of the developed scaffolds, it is necessary to improve the fixation methods and evaluation of the performance of the implants in longer periods.

To regenerate the injured meniscus, it is necessary to consider the structure and anatomical position of the native tissue. The native wedge-shape semilunar disc shapes of meniscus adhere to transverse, medial collateral, and femoral ligaments [143]. Recently, a trilayered wedge-shaped silk meniscal structure was designed to mimic the native structure of the meniscus [144]. Similar to the native meniscus, the fibroblasts were seeded outside of the scaffold and the chondrocytes inside the scaffold. The immunohistological and histological analyses exhibited the maintenance of chondrogenic phenotype with up-regulation of GAG and collagen type I and II. High collagen deposition indicated cellular proliferation and ECM production through different layers of the silk scaffolds, which was similar to the native meniscus (Fig. 6) [144]. In another study, autoclaved eggshell membrane powder reinforced three-dimensional porous silk fibroin-polyvinyl alcohol scaffolds containing unique combination of biomolecules (UCM) were used to repair meniscal injuries [145]. About 3% of autoclaved eggshell membrane in scaffolds provided an appropriate substrate for growth and ECM secretion of primary human meniscal cells. Subcutaneous implantation of the scaffolds in a rabbit model also exhibited low immunogenic response with acceptable biocompatibility and vascularization based on immunohistochemical assessment [145]. The regeneration capacity of the injured meniscus in a rabbit meniscectomy model was also evaluated by implanting freeze-dried collagen-coated sponge silk scaffold [146]. The presence of collagen in the outer layer remarkably enhanced the initial friction. After 3 month postimplantations, more cartilage tissue ingrowth with less cartilage wear was observed in silk-collagen group compared with pure silk sponges [146]. It was also reported that SF-BMSCs porous scaffolds revealed acceptable capacity in regenerating meniscus injuries after 6–12 weeks in rabbit model. At the implantation site, high cartilage-like structure, collagen type I and II, and GAG were formed [147]. The ability of the silk scaffolds in regenerating partial meniscus injuries was also examined in a sheep model. The scaffolds withstood the mechanical loads and caused no inflammatory response over 6 months. However, in some cases, the scaffolds were not stable in the defected region causing gap between the host and scaffold or full loss of the implant. To better integrate the implants with the host tissue, it is essential to develop better fixation methods [148]. For improving surgical fixation, SF fibers were incorporated into the porous substrate and were then used for substituting partial meniscal defects in a sheep

10. The promotion of osteochondral repair using silk fibroin scaffolds Osteochondral defect (OCD) includes damage in the articular cartilage and subchondral bone. Currently, some short-term treatment strategies are clinically available for osteochondral repair including allografting, microfracture, and mosaicplasty. However, the long-term regeneration of OCD using these methods faces some drawbacks due to the graft rejection, donor site morbidity, unwanted formation of fibrocartilage tissue, and inadequate integration of donor with the host tissues [154]. Tissue engineering is considered as a potential alternative for repairing OCD. There are some commercial constructs for repairing osteo/chondral tissues including Ostecel comprising hydroxyapatite (HAp) and autologous MSCs, INFUSETM as a recombinant human bone morphogenetic protein (rBMP-2), VITOSS® consists of calcium phosphate (CaP) and bone bonding protein, and CORTOSS as a synthetic bone void filler. These structures are based on synthetic materials; thus, might fail to support long-term regeneration of osteochondral tissue due to provoking the host immune systems and unbalanced degradability with formation of new tissue [155]. Recently, some researchers have focused on using SF-based substrates applicable for OCD regeneration. For example, the potential of B. mori (mulberry silk) and Antheraea mylitta (non-mulberry silk) based silk scaffolds in 259

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Fig. 4. Immunohistochemical analysis of repaired ligament after 24 weeks: (A, B) collagen I, (C, D) collagen type III, (E, F) tenascin-C 24 weeks post-implantation. Note: The arrows and asterisks specify the sutures and degraded silk fibers, respectively. Copyright © 2009 Elsevier B.V. Reprinted with permission from Fan et al. [135].

cartilage-like tissue and well-organized subchondral bone structures (Fig. 7) [158]. Injection of parathyroid hormone-related protein (PTHrP) in the intra-articular space along with implanting collagen-silk scaffolds was beneficial for repairing OCD by improving chondrogenesis and increasing cartilage repair in a rabbit model after 4–6 weeks. PTHrP partially blocks the Wnt/β-catenin signaling pathway and thus had an inhibitory effect on the terminal differentiation. Additionally, the time interval between 4 and 6 weeks after surgery was the best time for injection of PTHrP together with implantation of collagen-silk scaffolds [159]. Concomitant regeneration of cartilage and subchondral bone injuries using mono-phase substrates is challenging because these scaffolds are unable to repair bone and cartilage, simultaneously. An ideal OCD construct should have a biphasic (bilayer) structure (Fig. 8) comprising a chondrogenic matrix with small pore sizes to form hyaline-like structures with suitable flexibility and stretch, and an osteogenic matrix with large pores to mimic the subchondral bone with good mechanical strength and bioactivity [160]. There are some commercial bilayered scaffolds such as Trufit® and MaioRegen® that are used in clinic for treatment of OCD [161–163]. In a study, a biphasic scaffold comprising two SF species (B. mori and A.

regenerating osteochondral defects was evaluated in vitro and in vivo [156]. Culturing hBMSCs on non-mulberry scaffolds for four weeks induced chondrocyte-like cell formation; however, mulberry ones promoted formation of bone-like nodules. Animal studies showed that cellfree silk scaffolds were able to form neo-osteochondral constructs by absorbing TGF-β3 or rBMP-2. The ECM formed in the mulberry scaffolds was mainly constructed from collagenous compartments, particularly collagen type I; while, proteoglycans and collagen type I and II were the main ECM components of non-mulberry scaffolds [156]. In a comparative study, the regeneration capacity of fiber scaffolds and fibrin glue with or without chondrocytes in repair of the rabbit knee osteochondral defects during 36 weeks showed that the scaffolds containing cells were more potent than those without cells. The regeneration capacity of SF and fibrin glue showed no remarkable differences [157]. Other study exhibited that the implantation of silkchondroitin scaffolds in injured osteochondral sites of rabbit promoted neo-tissue formation and more structural restoration compared with silk scaffolds alone [158]. After six weeks post-implantation, silk and silk-chondroitin scaffolds were filled with fibrous and fibrocartilage tissues. Twelve weeks post-implantation, the silk scaffolds were invaded by fibrocartilage and cartilage-like tissues; however, the silk-chondroitin scaffolds were able to repair the defect site by producing 260

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Fig. 5. Meniscus engineering using silk fibroin to treated defected meniscus as useful strategy.

Fig. 6. Representation of cells seeded on the medial and lateral layers of silk scaffold to mimic the native meniscus tissue and histological analysis of the matured tissue. Scale bar; 100 µm. Copyright © 2009 Elsevier B.V. Reprinted with permission from Mandal et al. [144].

model, better repair of cartilage and the subchondral bone tissue was observed in biphasic group compared with monophasic ones [164]. Incorporating osteoconductive materials into bilayered scaffolds could also be useful in facilitating the formation of subchondral bone. CaP based materials were good candidate to stimulate the regeneration of subchondral bone; however, this ceramic phase may stimulate hypertrophy of chondrocytes [165]. Porous bilayered constructs

assamensis) was fabricated for regenerating OCD defects [164]. The suitable mechanical strength of the scaffold supported osteochondral repair in a load-bearing site. The growth and attachment of cells also improved due to the inherent presence of RGD motifs in the structure of A. assamensis silk. Osteochondrogenic gene markers were also overexpressed by using the biphasic construct. After 8 weeks post-implantation of the scaffold in osteochondral bone defect in a rabbit 261

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Fig. 7. Histological analysis after 12 weeks post-operation. (A) H&E evidence form: (a, d) Non-treated group; (b, e) silk group, and (c, f) silk-CS group. Arrows show the rudimental construct. (B) Safranin O results from: (a, d) Non-treated group; (b, e) silk group, and (c, f) silk-CS group. (Aa–c) and (Ba–c): scale bar: 500 mm. (Ad–f) and (Bd–f): scale bar: 200 mm. Copyright © 2009 Elsevier B.V. Reprinted with permission from Zhou et al. [158].

Fig. 8. Treatment of knee chondral defect using biphasic scaffold based on silk fibroin containing chondral and subchondral parts. 262

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ADSCs seeded on the bony layer and cultured in osteogenic medium was differentiated to osteoblast cells by production of calcium and collagen type I [172]. Electrospun nanofibers containing three gradient layers of CH, SF, and nHAp were also prepared to regenerate osteochondral injuries [173]. The scaffold was composed of three layers including CH/HAp composites at the bottom layer, integrated nanofibrous membrane at the middle layer, and CH/SF composites at the top layer to induce subchondral bone growth, calcification, and chondral repair, respectively. The layered construct supported the proliferation of both osteogenic and chondrogenic cells; while providing suitable substrate for mineralization of the neo-formed tissue [173]. Taken together, the regenerating of OCD is more complicated compared with osseous and chondral injuries alone. We need complex scaffolds with the ability to repair both cartilage and bony tissues in situ. The strategies to stimulate osteo-chondral regeneration first relied on single-phase scaffolds; while bilayer or trilayer scaffolds were later developed to promote cartilage and bone repair in different layers, simultaneously. Applying scaffolds with suitable biodegradability along with using tissue specific cells (e.g., chondrogenic, osteogenic, and/or stem cells) lead to the formation of osteochondral tissue in vitro and in vivo. Among the applied biomaterials, silk revealed outstanding potential in regenerating cartilage defects and subchondral bone. Some important growth factors could be incorporated into SF scaffolds that provide an environment for differentiating the stem cells into chondrogenic and osteogenic lineages.

containing fully integrated SF layer and silk-nano CaP were studied for OCD repair [166]. The fabricated scaffolds possessed good mechanical strength and stability, tailorable porosity, and suitable CaP deposition. rBMSCs seeded on silk-nano CaP layers in osteogenic medium was able to differentiate better into the osteogenic lineage compared with those seeded on silk layers. This observation was based on alkaline phosphatase (ALP) production which suggested that CaP promoted the osteogenic differentiation of rBMSCs. The bilayered scaffolds also had good integration with the host tissue after implanting in the knee of a rabbit model. During OCD repair, it is essential to stimulate the rapid formation of subchondral bone because it can provide suitable mechanical support for regenerating cartilage tissue and fix the scaffold in the injured site. Additionally, top layer of bilayered scaffold promoted cartilage formation; while the bottom section containing a silk-nano CaP layer stimulated in-growth of subchondral bone and some signs of angiogenesis [166]. Ruan et al. fabricated a biphasic scaffold containing SF/chitosan (SF/CH) and SF/CH/nHAp to induce chondrogenic and osteogenic differentiations of BMSCs, respectively [167]. The porosity and water uptake of the scaffold were 90% and 822%, respectively. The cell seeded on biphasic scaffold exhibited filling activities upon implantation in defected cartilage in the femur of rabbit model. Based on micro-CT evidences, the regeneration of chondral and subchondral bone tissues was achieved. Furthermore, the expression of collagen type I and II was increased by using the biphasic scaffold [167]. Electrospinning technique was also used to prepare bilayer scaffold with the ability to support the proliferation and growth of chondrogenic and osteogenic precursor cells [168]. The osteogenic phase was comprised of 70S bioactive glass (70SiO2·25CaO·5P2O5) and the chondrogenic phase was comprised of SF from B. mori and Antheraea assama species. The electrospun of non-mulberry nanofibers revealed better performance in terms of biological, mechanical, and physicochemical properties compared with electrospun mulberry nanofibers. The specific markers of cartilage tissues, such as Sox-9 and aggrecan, were upregulated on chondrocyte-seeded silk scaffolds which confirmed the benefits of using these materials for chondrocyte cultivation. Bone specific markers like bone sialoprotein (BSP) were over-expressed on osteoblast seeded bioglass mats in comparison to silk mats suggesting the advantages of using bioglass for cultivating osteoblasts. Runx-2 was highly expressed on bioactive glass part than silk mats. Runx2 is an early differentiation factor produced by committed osteoblast cells and indicated the onset of mineralization. Similarly, Christakiran et al. fabricated scaffolds based on layer by layer method [169]. The first layer was comprised of electrospun 70S bioactive glass and the second layer was consisted of (mulberry B. mori and endemic Indian nonmulberry Antheraea assama). The well-integrated layers of the biphasic scaffold was able to support the proliferation and maturation of both osteogenic and chondrogenic cells at the interface. The scaffold containing non-mulberry silk showed higher expression of osteogenic and chondrogenic gene markers than those comprising mulberry silk [169]. Another biphasic scaffold was prepared that consisted of a cartilage phase containing silk protein and a bone phase comprise of silk coated strontium-hardystonitegahnite ceramic scaffold. The biphasic scaffold was able to induce the chondrogenic or osteogenic differentiation of hMSCs seeded on the scaffolds [160]. To better mimic the native osteochondral structure, a trilayered scaffold containing chondral, calcified, and bony layers would be more advantageous for concomitant regeneration of cartilage, calcified cartilage, and bone, respectively [170,171]. Accordingly, a trilayered scaffold composed of SF and HAp was fabricated using a paraffin-sphere leaching and temperature gradient-guided thermal-induced phase separation method [172]. The chondral layer promoted differentiation of ADSCs towards chondrocyte-like cells by producing cartilage matrix containing GAG, collagen type II, and over-expressing chondrogenicrelated genes. The results revealed that the chondral layer induced chondrogenic differentiation of ADSCs in chondrogenic media containing TGF-b1 and insulin-like growth factor (IGF). Furthermore,

11. Concluding remarks and future perspectives The long-life expectancy, the increase in global population, and subsequently the increase in the number of cartilage injuries motivate scientists to develop new strategies for repairing the damage cartilage tissue. The progressive therapies using novel biomaterials with the ability of controlling the release rate of biomolecules and the accommodation of various cell sources are also increasing. Over the past decade, many constructs based on natural or synthetic polymers are used for cartilage and osteochondral regeneration. Some of these materials support the proliferation and spreading of chondrogenic and osteogenic cells with maintaining their normal morphologies. Among the numerous biomaterials used for this purpose, silk protein gains attention for cartilage, ligament, and musculoskeletal tissue engineering. To date, various strategies have been used to fabricate three-dimensional silk-based scaffolds for cartilage regeneration. Among them, hydrogels are more demanding for cartilage tissue engineering purposes due to suitable swelling capacity. In recent years, with the help of bio-printing, complex hydrogels containing bioactive molecules and cells are prepared for cartilage repair. The outstanding properties of silk fibroins in terms of biocompatibility, biodegradability, suitable biological, and mechanical strength make it a good choice as a bio-ink or printable scaffolds that preserve their shape for a long-term shape constancy [174]. Although, silk based structures have the ability to promote the proliferation and adhesion of chondrocyte cell; but there is no intrinsic sequence in backbone of silk fibroin that induces some chondrocyte cell signaling pathways. The possible molecular mechanism and signaling pathway of silk in promoting chondrocyte behavior is still unclear. This evaluation is needed in immediate future. To examine the safety and feasibility of silk fibroin based scaffolds for cartilage regeneration in animals are essential to validate the in vitro experimentations so far carried out. Small animal models (e.g., rabbit, rat, and mice) are not suitable for the demands due to variations in ECM structure and composition. The difficulties in creating animal models for cartilage defects are still not feasible. In such situation, larger animal models may be preferable for clinical investigations. Acknowledgments This work was supported by Pasteur Institute of Iran. 263

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Conflict of interest

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