Nonmulberry Silk Fibroin Scaffold Shows Superior ... - Semantic Scholar

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May 16, 2015 - Marko Loparic , Subhas C. Kundu , * Sourabh Ghosh , * and Asok Mukhopadhyay *. DOI: 10.1002/adhm.201500283. Recent years have ...

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Nonmulberry Silk Fibroin Scaffold Shows Superior Osteoconductivity Than Mulberry Silk Fibroin in Calvarial Bone Regeneration Neety Sahu, Prakash Baligar, Swati Midha, Banani Kundu, Maumita Bhattacharjee, Snehasish Mukherjee, Souhrid Mukherjee, Florian Maushart, Sanskrita Das, Marko Loparic, Subhas C. Kundu,* Sourabh Ghosh,* and Asok Mukhopadhyay*

Recent years have witnessed the advancement of silk biomaterials in bone tissue engineering, although clinical application of the same is still in its infancy. In this study, the potential of pure nonmulberry Antheraea mylitta (Am) fibroin scaffold, without preloading with bone precursor cells, to repair calvarial bone defect in a rat model is explored and compared with its mulberry counterpart Bombyx mori (Bm) silk fibroin. After 3 months of implantation, Am scaffold culminates in a completely ossified regeneration with a progressive increase in mineralization at the implanted site. On the other hand, the Bm scaffold fails to repair the damaged bone, presumably due to its low osteoconductivity and early degradation. The deposition of bone matrix on scaffolds is evaluated by scanning electron and atomic force microscopy. These results are corroborated by in vitro studies of enzymatic degradation, colony formation, and secondary conformational features of the scaffold materials. The greater biocompatibility and mineralization in pure nonmulberry fibroin scaffolds warrants the use of these scaffolds as an “ideal bone graft” biomaterial for effective repair of critical size defects.

N. Sahu, Dr. P. Baligar, S. Mukherjee, Prof. A. Mukhopadhyay Stem Cell Biology Laboratory National Institute of Immunology Aruna Asaf Ali Marg, New Delhi-110067, India E-mail: [email protected] Dr. S. Midha, Dr. M. Bhattacharjee, S. Mukherjee, S. Das, Dr. S. Ghosh Department of Textile Technology Indian Institute of Technology Delhi, Hauz Khas, New Delhi-110016, India E-mail: [email protected] Dr. B. Kundu, Prof. S. C. Kundu Department of Biotechnology Indian Institute of Technology Kharagpur, West Bengal-721302, India E-mail: [email protected] F. Maushart, M. Loparic Biozentrum and Swiss Nanoscience Institute University of Basel Klingelbergstrasse 70, 4056 Basel, Switzerland

DOI: 10.1002/adhm.201500283

Adv. Healthcare Mater. 2015, 4, 1709–1721

1. Introduction

An “ideal bone graft” biomaterial for effective repair of critical size defects should be osteoconductive (capable of directing/guiding new bone growth), as well as osteoinductive (encourage the osteoprogenitor cells to differentiate into active osteoblasts and produce bonespecific mineralized matrix) in nature. However, the currently used biomaterials as bone substitutes for reconstruction of cranial and other bone defects have several limitations, such as poor osteoinductivity, difficulty in shaping the graft, cell-mediated immune response to “foreign” biomaterial, which ultimately cause poor osseointegration.[1,2] The ceramic bone grafts, although frequently used, suffer from high brittleness, poor tensile strength, variable rates of resorption, and have adverse effects on bone remodeling in vivo.[3] These bone grafts are also characterized by low porosity, ultimately resulting in loose and fractured implants. Similarly, the poor mechanical strength of hydroxyapatite restricts its use in load bearing applications.[4] Compared to other biodegradable biopolymers, silk fibroin protein is a promising candidate for tissue engineering of bone due to its tunable mechanical properties and architecture, biocompatibility, tailor-made biodegradability, and ability to support bone formation.[5,6] Mulberry silk, Bombyx mori (Bm) fibroin scaffolds have shown promising results in vitro for adhesion and subsequent differentiation of cultured mesenchymal progenitors (MSCs) into osteoblasts, and formed bonelike mineralized structures.[7] The flexibility of silk polymer processing allows simulation of porous morphologies of cortical and trabecular bones and, as a result, facilitates osteogenic differentiation of seeded MSCs.[8] However, Bm fibroin scaffolds by themselves, when implanted without any cells, showed insufficient infiltration of host cells required for substantial in vivo healing of critical-sized calvarial defects in rat.[9] Implantation of Bm scaffolds with un-differentiated human MSCs has been found to form bone-like matrix only in the periphery of

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the mice femoral defect,[7] whereas scaffolds with osteogenically predifferentiated MSCs could promote substantial amount of bone formation.[9] Various attempts have been made to further enhance the osteogenic potential of silk matrices either by altering the surface chemistry of the matrix or by employing genetically manipulated cells. The inclusion of polyaspartic acid during scaffold preparation subsequently followed by controlled deposition of calcium phosphate lead to premineralized fibroin scaffolds, which have been found to increase the expression levels of bone morphogenetic protein (BMP-2).[10] Canine bone marrow stromal cells, over-expressing BMP-2 and cultured on apatitecoated silk scaffolds resulted in enhanced bone formation in nude mice, as compared to the control cells.[11] The expression levels of important osteogenic genes, such as osterix and osteocalcin, have been found to increase with culture time when Bm scaffolds are seeded with induced pluripotent stem cells over-expressing a transcription factor Special AT-rich sequencebinding protein 2.[12] This construct helped in correction of 4 mm critical sized calvarial bone defects within 5 weeks of implantation. However, mulberry silk fibroin lacks arginine–glysine– aspartate (RGD) attachment site which is a known cell adhesion ligand peptide motif.[13] It has been found that silk fibroin obtained from silkworms not feeding on mulberry leaves, such as indian tropical tasar silk, Antheraea mylitta (Am), or chinese oak tasar silk, Antheraea pernyi, naturally possesses RGD motifs.[14] Attempts have been made to incorporate this motif into mulberry silk by covalent attachment to fibroin macromolecules. A marked improvement in the osteoblast attachment and differentiation was noticed after introducing the RGD motif in Bm matrix.[15] However, there is a concern that in chemically decorated silk, this extrinsically added RGD motif may not be accessible for cellular interactions due to entrapment during extensive chain entanglement and β-sheet formation.[16] The natural presence of RGD motif in nonmulberry silk Am imparts pronounced cell attachment, osteogenic differentiation, and mineralization.[17] However, the translation of bone defect healing with Am scaffolds to clinical trials seems to remain a challenge, as of now, due to the limited number of in vivo studies performed using Am silk-based biomaterials. The present study explores the strategy of implanting cellfree silk fibroin scaffolds into a critical size cranial defect in rats. It is a pioneering study that attempts to quantitatively compare the extent of osteogenic generation in rat cranial defect mediated by the Bm and Am scaffolds. The use of cell-free scaffolds as biomaterials is of great significance in clinical trials. This eliminates the chances of graft rejection and complicated cell isolation protocols from patients, and limits the expenditure incurred in current good manufacturing practice (cGMP) compliance.

2. Results 2.1. Pore Geometry of Scaffolds Further analysis revealed that the average pore sizes of Am and Bm scaffolds were 76.37 ± 5.63 µm and 73.12 ± 9.51 µm, 1710

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respectively, suggesting nominal differences in pore size of the two scaffolds (Figure 1A). The porosity of scaffolds was 60% for Bm and 66.66% for Am, as determined by liquid displacement method.

2.2. FTIR The attenuated total reflectance-Fourier transform infrared (ATR-FTIR) spectroscopic analysis was carried out for these scaffolds to determine the secondary structures. Both the scaffolds revealed prominent amide I, II, and III peaks at 1621, 1516, and 1232 cm−1, respectively (Figure 1B). To further enumerate the types of secondary structures, the spectra were deconvoluted by Fourier self-deconvolution method and the results were appended in Table 1. In the present study, the Am scaffold was found to possess 2.6 times greater β-sheet crystal content as compared to Bm scaffold.

2.3. In Vitro Biodegradability of Am and Bm Scaffolds Exposure of the two scaffolds to protease XIV in vitro revealed that Bm degraded much faster in the initial 5 d of incubation, whereas Am scaffold retained more than 75% of the mass under identical experimental conditions (Figure 1C).

2.4. Am Facilitates Osteoblast Colony Formation The colony forming ability of rat primary osteoblasts and adipose-derived (AD)-MSCs cultured on Am and Bm silk fibroins was conducted to examine the extent of biological mimicry of extracellular matrix (ECM) in the two scaffolds for bone tissue engineering. Am fibroin coated tissue culture plates showed significantly (p < 0.01) more colonies both for osteoblasts and AD-MSCs as compared to Bm fibroin (Figure 2). These results indicated an intrinsic propensity of Am fibroin compared to Bm that supported better colony forming efficiency in osteoblasts and related cells. The test was conducted in triplicates for each cell and scaffold type.

2.5. In Vivo Degradation of Am and Bm Scaffolds Both types of scaffolds were well tolerated by the host animals within the peritoneum cavities, with no visible signs of rejection within the first month. Over long term, no obvious signs of degradation were noticed in the Am scaffolds up till 12 months, whereas the Bm scaffolds started to gradually degrade within 3 months of implantation. No remnants of Bm scaffold could be detected in subsequent samples over long-term examination (Table S1, Supporting Information). Microscopic examination showed thinning of the biomaterial and lodging of about four folds more immune cells on Bm scaffold as compared to Am (Figure 3A). Further, to confirm their relative degradation and osteogenic potential in vivo, we implanted these scaffolds in critical size calvarial bone defects of rats. Histological analyses suggested

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FULL PAPER Figure 1. Physical characterization of Am and Bm. A) Representative SEM showing porous architecture of scaffolds. B) ATR-FTIR for quantification of secondary conformation of scaffolds. C) Degradation of fibroin scaffold materials.

that like peritoneum cavity, the Bm scaffold degraded much faster than Am in cranial defects (Figure 3B). In addition, most of the defect site in Bm constructs was infiltrated with a dense network of mononuclear cells encapsulating the scaffold (Figure 3B). The presence of soft tissue, i.e., fibrous connective tissue and a chronic inflammatory

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response (which was later confirmed by CD11b staining) could be observed, interspersed with some blood vessels (Figure 3B). While some signs of ossified-like matrix (as validated by von Kossa (VK) and Alizarin Red S (ARS), Figure 4) were evident in 1 month Bm constructs, the ossification could not be further detected in 3 month Bm explants. Moreover, within 3 months

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www.MaterialsViews.com Table 1. Secondary structural conformations of Am and Bm silk fibroin scaffolds. β-sheet [%]

α-helix [%]

β-turns [%]

Random coils [%]

Am

22.8

8.66

55.5

12.6

Bm

8.04

13.15

61.5

16.91

Test sample

of implantation, the residual biomaterial was minimally noticed in the case of Bm scaffold and the presence of soft tissue in growth near the defect margins was conspicuous (Figure 3B, inset). Abundant lymphocytes, macrophages, and multinucleated giant cells (Figure 3A,B); typical cells involved in chronic inflammation, closely associated with the Bm surface could be identified by H&E staining. As macrophages are the primary immune cells involved in the degradation of foreign objects in the host tissue, therefore we examined the fibroin protein explants for macrophage infiltration via immunohistochemical analysis. To confirm this, 1 month and 3 months old explants were harvested for both type of fibroin scaffolds and subsequently stained for CD11b antigen, a typical marker of monocytes/macrophages. On visual examination, the representative images showed higher infiltration of immune cells in Bm over Am scaffold (Figure 3C). Upon manually counting the total number of CD11b+ cells per field of view for quantitative analysis, it was further validated that significantly (p < 0.05) higher CD11b+ cells per field were present in Bm over Am scaffold, corresponding to an increase of 1.6 fold after 1 month and 1.3 fold after 3 months in explants. Moreover, the total number of multinucleated giant cells (cells with >3 nuclei were counted; Figure S2, Supporting Information) infiltrating the defect site were quantitatively analyzed from H&E stained specimens (Figure 3B, inset). The results revealed that Bm implanted calvarial defects demonstrated maximum number of giant cells counted per field (3-5 cells in 1 month and 1-2 cells in 3 month

Bm sites) as opposed to an occasional giant cell found in the case of Am constructs at any given time point. Thus, the overall results suggested that Bm scaffold was more sensitive to biodegradation, in which presumably macrophages and multinucleated giant cells played a major role. On the contrary, new bone formation, with a typical lamellar structure, was conspicuous within Am scaffolds in 3 months. Fibrous connective tissue formation was seen interspersed within the defect where new bone was being formed. Further, the infiltrated mononuclear cells (indicative of inflammation) gradually declined over the 3 month period in Am matrix. This was confirmed by the relatively lower abundance of CD11b+ cells in Am implanted defects (Figure 3C). Ossified-matrix like tissue (seen as dark patches) within the defect, was juxtaposed to the Am surface. Mineralization of this tissue was further validated by von Kossa and Alizarin Red S staining.

2.6. Neo-Bone Formation in Silk Scaffolds 2.6.1. Radiological Analysis Representative X-ray radiological images of normal calvarial bone, bone with artificially created defect, and an implanted are shown in Figure 4A. The constructs were harvested after 1, 3, and 6 months of implantation from rat calvariae, and the neo bone regeneration within the scaffolds was examined using X-ray radiological imaging. Negligible bone regeneration was observed in the inner edge of the controls without any implanted scaffolds. The fibroin scaffold, by virtue of being radiolucent, facilitated the assessment of the extent of ossification by visualization of the mineralized matrix deposition. There was no obvious mineralization observed in the case of Bm scaffold even after 6 months postimplantation (Figure 4B, right panel). On the other hand, Am scaffolds showed gradual filling of the cranial defects with considerable deposition of bone-like matrix (Figure 4B, left panel).

2.6.2. Histological Analysis

Figure 2. Colony forming efficiency. Bar diagram shows significantly more number of colonies in Am (p < 0.05) than Bm coated wells. TCP: Tissue culture plate; OB: Osteoblasts; MSCs: Mesenchymal stem cells.

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In order to validate the X-ray radiological observations and to further study the distribution of the bone matrix, histological studies were performed. Using von Kossa and Alizarin Red staining, a thin septa of bone-specific ECM deposition was observed juxtaposed to the margins of the residual Bm scaffold after 1 month of implantation (Figure 4C). Surprisingly, no staining with either of the dyes was noticed in the same scaffold 3 months postimplantation (Figure 4C). This nominal evidence of mineralization along the Bm scaffold margins was probably even below the detectable range of X-ray radiography (Figure 4B) and therefore could not be observed. This may be explained by the relatively lighter shade of staining observed in Bm defects over the heavily ossified Am defects. A gradual progression of mineralized ECM synthesis was seen in Am scaffolds over 3 months period (Figure 4C). This was further confirmed by collagen specific picrosirius red staining of constructs and examined by light microscopy. In 1 month constructs (both Am and Bm), numerous fiber-like

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FULL PAPER Figure 3. Histology and immunohistochemistry of silk scaffolds harvested at different timepoints. A) H&E stained sections of scaffolds harvested from peritoneal cavity 3 months postsurgery. Presence of partially degraded implants in Am and Bm (white arrowheads), while numerous immune cells (black arrowhead) interspersed with some blood vessels (black arrows) observed in Bm defects. B) Scaffolds harvested from cranial defects 1 and 3 months postsurgery. Inset shows soft tissue ingrowth in Bm, black arrows indicate blood vessels. C) CD11b+-stained Am and Bm scaffolds for comparative analysis of immunological responses. D) Bar diagram showing significantly more macrophages in Bm (p < 0.05) over Am scaffold. M: Month. Scale Bars = 10 µm (A), 100 µm (B), 50 µm (C).

structures (Figure 5A, arrows) were interspersed throughout the sections that showed a faint pink-red staining; typical of newly formed, immature collagen deposition. The fibers were juxtaposed to the material surface, organized in a circumferential pattern. No significant differences were observed in terms of the percentage collagen proportionate area between the two constructs after 1 month of implantation (Figure 5B). However, after 3 months of implantation, conspicuous collagen deposition with mature cross-linked fibers marked by intense red staining was evident in Am constructs (Figure 5A, arrows). Whereas, the percentage proportionate area of collagen associated with Bm constructs observed at the same time point was significantly lesser than Am constructs (p < 0.01) (Figure 5B). Moreover, within the Bm group, the intensity of collagen specific staining diminished over time (from 1 to 3 months), concomitantly with the progressive increase in CD11b+-stained macrophages, suggestive of an immunoinfiltration response (Figure 4D). In Bm implanted defects, 6 month samples could not be processed for histological analysis due to excessive degradation of the material, while the Am scaffolds were highly ossified and hence brittle which made it difficult to obtain sections without

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decalcification. Furthermore, Am scaffolds were well-integrated with the native host bone to the extent that detachment of these specimens required application of force at the time of implant recovery. In vivo, the Bm scaffold could support limited growth of osteoblastic cells (as suggested by the abundance of immunofiltrate and nominal evidence of ossification in the defect site) which led to restricted secretion of ECM up till 1 month postimplantation. Bm scaffold could support limited growth of osteoblastic cells and secretion of ECM up till 1 month postimplantation. On the other hand, the Am scaffold was capable of promoting new bone formation to a greater extent than Bm scaffold. This may be due to the higher structural stability of Am that allowed osteoblastic cells to proliferate, facilitate extensive ECM secretion followed by enhanced ossification over Bm constructs. The inadequate regeneration of bone on the Bm scaffold was most likely due to the extensive infiltration of phagocytes and other immune cells, as compared to Am (as confirmed by CD11b staining, Figure 3C) which may have been responsible for their subsequent degradation under the effect of matrix metalloproteinase secreted by the macrophages.

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Figure 4. Bone regeneration on Am and Bm scaffolds. A) X-ray radiograph showing unoperated calvaria (left) and 5 mm bilateral defects on day 0 of surgery (right image). Centre image shows surgical site on rat calvarial surface with implanted scaffold on one side. B) Radiological analysis of 1, 3, and 6 months postsurgery showing gradual increase in neo bone formation in Am constructs (left panel: a–c). No bone formation occurs in Bm constructs 6 months postimplantation (right panel: d–f). C) Histological analysis of bone matrix formation 1 and 3 months postsurgery stained with von Kossa and Alizarin Red S. Enhanced matrix deposition in Am within 3 months (arrowheads) (left panel: a–d). In Bm constructs, some deposition observed after 1 month, which disappeared after 3 months (right panel: e–h).

fibrous ECM was observed (Figure 6H), albeit only in peripheral regions.

2.7. SEM of Excised Constructs Scanning electron microscopy (SEM) images of Am construct retrieved after 3 months (Figure 6A–D) showed clear cellular ingrowth (Figure 6D) on surface of the construct as well as within the pores. Osteoprogenitor cells entrapped within the pores formed clear attachment to Am matrix and produced osteoid matrix. By contrast, SEM studies confirmed the presence of fibrous capsules on the Bm construct (Figure 6E–H). There was a significant amount of fibrous ECM detected on both the silk fibroin matrices (Figure 6D,H). In some locations of Bm constructs, mineralized matrix deposition along with

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2.8. Atomic Force Microscopy (AFM) of Excised Constructs The AFM images of empty Am fibroin scaffold demonstrated an aggregated rough surface morphology, where occasional fibrillar structures of about ≈0.25 nm diameter were visible (Figure 7A). In case of native calvarial bone, pristine collagen fibrils with an average diameter of 120 nm (range from 50 to 200 nm) with repeats of ≈67 nm along the length of individual

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FULL PAPER Figure 5. A) Picrosirius red stained sections of Am and Bm rat calvarial constructs 1 and 3 months post-implantation. Enhanced collagen deposition observed in Am within 3 months (arrows) (top panel). In Bm constructs, some immature collagen deposition observed after 1 month, which diminished after 3 months (bottom panel). B) Bar diagram showing significantly more percentage of collagen proportionate area in Am (p < 0.01) over Bm scaffold. M: Month.

fibrils (Figure 7B) were clearly visible. Most of the collagen fibrils were randomly oriented, while some of them were arranged in parallel bundles with an average bundle diameter of 600 nm. These findings are in accordance with earlier reports on high resolution imaging of bone.[18] Six months postimplantation, explants exhibited collagen fibrillar-like morphologies similar in appearance to the native bone collagen structures, though the size of the fibril bundles and sub-fibrillar organization was found to vary (Figure 7C). Individual collagen fibrils were highly oriented and formed a parallel arrangement and formed larger bundled fibers

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compared to the native bone sample. Average diameters of such fused parallel bundles were measured to be around 1.7 µm.

3. Discussion Conventionally, osteoblasts are obtained from autologous bone biopsies and expanded in vitro for bone tissue engineering. However, the lower yield from explants, time-consuming culture process, donor site morbidity, and the associated financial

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Figure 6. SEM of constructs. A–D) Am constructs showing colonization by osteoprogenitor cells 3 months postsurgery. Micrographs “B” and “D” represent magnified images from selected regions of “A” and “C.” E–H) Bm constructs depict the formation of fibrous capsule. Micrographs “F” and “H” represent magnified images from selected regions of “E” and “G.” Fibrous ECM detected on both the constructs (“D” and “H”). Arrowhead marks indicate mineralized ECM matrix depositions (F, H).

burden have prompted the need for biomaterials that can act as bone substitutes with inherent osteoinductive and osteoconductive properties. Biocompatibility and the ability to process tailormade scaffolds with desired mechanical properties, in order to match the features of the target bone defect, are the two major criteria for the selection of biomaterials for in situ bone tissue engineering.

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Silk fibroin protein is one of the biomaterials currently being used as a biomimetic matrix for various tissue engineering applications, including regeneration of bone.[19] Earlier reports demonstrated that implantation of Bm scaffolds without cells or with undifferentiated stem cells failed to recruit host osteoprogenitor cells to heal bone defects.[7,9,20] Our earlier studies demonstrated that both 2D and 3D matrices of Am fibroin

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FULL PAPER Figure 7. AFM analysis of constructs. A) Representative image of Am construct, prior to implantation, showing aggregated rough morphology with occasional fibrillar structures. B) Native calvarial bone showing collagen fibers on surface. C) Am scaffold, 6 months post implantation, shows collagen fibers and branching, which assemble to form mature fiber bundles, within ridges of silk matrix. Compact, well-organized lamellar architecture of collagen fibers, with nested arches, simulate typical features of matured bone tissue. Bm matrix could not be collected 6 months post implantation due to extensive degradation.

facilitated osteogenic differentiation of human MSCs.[17] These findings encouraged us to conduct a comparative analysis between the performance of the two scaffolds, i.e., Am and Bm fibroins for in vivo bone regeneration. Not many studies have been reported in the literature directly comparing the osteogenic potential of the two silk sources, therefore, to the best of our knowledge this is the first, concrete in vivo evidence dictating the superiority of Am scaffolds over their Bm counterparts.

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Though qualitative, the study provides valuable preliminary evidence on the extent of the in situ osteogenic potential of the two silk types over long-term implantation, warranting further work for detailed quantitative evidence as well as the underlying molecular mechanisms that are responsible for generating superior bone formation in Am. However in the current study, the processing of the two silk types differs in the way that Am scaffold has been directly isolated from glands while Bm

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has been regenerated from silk fibroin solution. This is due to the poor solubility of Am cocoon in organic solvents or chaotropic reagents due to strong inter and intramolecular interactions between fibroin chains and hydrophobic nature of this variety of silk.[21] In the first stage of study, Am and Bm scaffolds were implanted in peritoneum cavities of Lewis rats to generate insights upon the immunological responses and degradation rate of the two scaffolds in vivo. In the second stage, Am and Bm scaffolds were implanted in critical size, nonload bearing calvarial defects of rats to evaluate their respective osteogenic potential. The primary reasons for selecting rat cranium as an animal testing model were twofold: (i) it renders the ability to create critical size defects, which is ultimate clinical target for such biomaterials, (ii) cranial reconstruction is considered a severe test for evaluating the potential of osteogenic substrates. This is due to its poor regenerative properties as a result of the lack of sufficient vasculature, a prerequisite for osteogenesis to occur which makes it a challenge for even smaller defects to regenerate spontaneously within this tissue. Am fibroin scaffold showed remarkably efficient bone regeneration in calvarial defects compared to Bm fibroin scaffold. New bone formation on an empty scaffold may be mediated by the migration and infiltration of the osteocytes or osteoblast precursor cells from the surrounding healthy bone tissue. This study suggests the recruitment of precursor cells from the surrounding native bone to the empty silk scaffolds, prior to proliferation and ossification. SEM studies confirmed cellular infiltration and mineralized matrix deposition. Histological analysis confirmed deposition of higher extent of bone-specific matrix in Am compared to Bm scaffolds. AFM results showed highly parallel arrays of collagen fibrils on Am, leading to the formation of bundled lamellar collagen fibers with nested arches, which would contribute to bone toughness as well as simulate the basement matrix in scaffolds similar to the architectural hierarchy of natural bone tissue. In fact, these structural cues were supported by osseointegration of bone with the host bone. On the contrary, Bm scaffolds did not support the formation of ossified bone except the thin septa of mineralization lining the margins of residual material within the defect site as evidenced by histology. In vitro studies have demonstrated spontaneous crystallization of HA on Bm surface upon exposure to simulated body fluid,[22] fetal bovine serum,[23] and calcium chloride solution.[24] Moreover, the β-sheet crystalline regions were proven to act as nucleation sites which subsequently formed HA-nanocrystals, similar to the role of collagen type I in native bone tissue.[23] This inherent mineralization of Bm may have been responsible for the small patches of mineralization found juxtaposed to the material surface (either along the margins or covering the material surface) in von Kossa and Alizarin Red S stained specimens. This was further validated by the collagenspecific staining of constructs which clearly demonstrated the diminishing levels of collagen deposition in Bm constructs, concomitant with the macrophage infiltration overtime from 1 to 3 months. This evidence suggests that the inherent osteogenic potential of Bm may be sufficient for provoking some nominal traces of mineralization in the initial period of implantation, which subsequently declined due to the gradual take over by the infiltrated mononuclear inflammatory cells, however, if

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coupled with additional pro-osteogenic agents (such as bonespecific growth factors or bioactive molecules), the osteogenesis in Bm may be triggered. In context of the above argument, we believe that the difference in the outcome of our results may be partly attributed to the presence of RGD motifs in Am fibroin, which was lacking in Bm fibroin. The presence of these cell adhesive RGD-peptide motifs has previously been identified.[25] Bone sialoprotein (BSP) has been shown to have osteoinductive properties; in which RGD motifs mediate cell attachment through α5β1 integrin.[26] Fifteen amino acids stretch of unique BSP, containing RGD motif, promoted focal contact formation, adhesion, and spreading of osteoblasts on quartz surfaces.[27] In an in vitro comparative study on Am and Bm, both silk proteins were immobilized on titanium surfaces and evaluated for osteogenesis and immunological responses. After 7 d of culture with osteoblast progenitors, increased gene expression profiles of osteogenic differentiation markers (alkaline phosphatase and osteocalcin) were found in Am compared to Bm, indicating the superiority of Am in bone formation in vitro.[28] The higher colony number and intense staining of von Kossa and ARS in 1 month explants confirmed improved cytocompatibility and matrix synthesis on Am scaffolds. These findings were in corroboration with earlier reports, which showed that induction of in vivo bone formation by expanded human MSCs is directly correlated to the clonogenicity of the progenitor cells.[29] The BMP-2 signaling pathway, which plays an important role in osteogenesis, is intimately related to the integrins. Since BMP-2 receptor overlaps with the integrins, the BMP-2 signaling pathway could be triggered by the engagement of α5β1 integrin with RGD motifs present in Am fibroin protein.[30] Furthermore, it was revealed that Bm scaffolds degraded much early as compared to their Am counterparts. A proper balance of biomaterial degradation and cellular ingrowth-cummatrix deposition would maintain tissue integrity during in vivo bone regeneration. When the scaffold degradation is too fast to be able to match matrix deposition, the structural integrity of the neo-tissue is lost, further resulting in collapse of the 3D architecture, which in turn leads to mechanical failure of the graft and poor bone regeneration. However, if the degradation is too slow, the impaired ECM synthesis and fibrotic encapsulation may impede the process of tissue regeneration.[31] Earlier studies reported long term degradation of Bm scaffolds in vivo.[32] The scaffolds made of 6 wt% Bm fibroin protein using aqueous-based processing methods completely degraded between 2 and 6 months in subcutaneous implantation site.[33] Whereas on the other hand, organic solventbased Bm fibroin scaffolds persisted beyond 1 year, although both scaffold types had comparable β-sheet content.[33] The kinetics of degradation of protein scaffolds is largely dependent on its structural features and immunogenicity. It has been shown that the composition of the secondary structures of the protein, particularly the β-sheet content, provides the structural stability against enzymes like collagenase IV, protease XIV, and α-chymotrypsin.[34] When fibroin protein is dissolved to prepare scaffolds, it assumes random coil structure which is more susceptible to proteolytic cleavage. In native silk fibroin, about 45%–60% of the secondary structure is β-sheet type as a result of the stabilization process during postprocessing of silk fibroin.[35] However, the extent of this

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4. Conclusions The above reported findings tempted to speculate the application of acellular Am scaffold as a promising platform for bone tissue engineering as it promotes effective migration and ingrowth of osteoprogenitor cells, bone-specific matrix accumulation, and effective bone regeneration compared to Bm fibroin scaffold. The aqueous-derived A. mylitta scaffold does not require functionalization such as grafting of RGD motifs for bone tissue repair in vivo. Further research is warranted to extend the application of this scaffold for bone regeneration in critical size load bearing defects by optimizing their 3D architecture, porosity, and incorporation of osteogenic components such as tethering of bone-specific growth factors for sustained release from these pro-osteogenic Am scaffolds.

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β-sheet formation depends upon the manufacturing process of the scaffold.[36] Both in vitro and in vivo examinations showed faster degradation Bm scaffolds compared to Am. As protease XIV, the degradative enzyme used in in vitro study, extensively degrades the fibroin β-sheet crystals, a possible contributing factor to faster degradation of Bm could be the relatively higher percentage of β-sheets in the Am scaffold as compared to Bm.[37] The low β-sheet content in the Bm scaffold could be either due to their low protein concentration (2%) or the fabrication method adopted in the present study, which resulted in faster degradation of the material.[38] Interestingly, though Am scaffolds were manufactured by an aqueous-based process and then subjected to ethanol treatment similar to Bm scaffolds, it was relatively stable in the rat peritoneum cavity as compared to the cranial bone defects. Therefore, it could be concluded that the site of implantation is another important factor that may govern the rate of degradation of silk fibroin scaffolds. Furthermore, we cannot exclude the possibility that the striking difference in degradation and regeneration capacity of Am and Bm fibroin scaffolds may be partly be due to the fact that the latter was reconstituted by dissolving cocoons, while the Am was directly isolated from silk gland. As discussed earlier, Am cocoons are difficult to dissolve due to the strong molecular interactions between fibroin chains and also due to their hydrophobicity.[21] Natural fibroin biopolymers are enriched with alanine and glycine repeat sequences, which rapidly form β-sheet crystals.[39] The dissolution and reconstitution process used to extract fibroin from the Bm cocoons might have disrupted some of the β-sheet crystalline domains in the regenerated Bm scaffolds, leading to their faster degradation in situ. A greater invasion of macrophages or monocytes and multinucleated giant cells was detected in Bm scaffold than Am. Through a direct co-culture model it was demonstrated that macrophages exhibit significantly lower levels of interleukin 1β, tumor necrosis factor α and nitric oxide (cytokines involved in inflammatory responses) when cultured in association with osteoblasts compared to macrophage only controls, which was most likely regulated by cellular cross-talk.[28] This may explain the reduced level of inflammatory cells observed in Am implanted constructs. The complete ossification of the Am implanted defect within 6 months suggests heavy infiltration of osteogenic cells and their subsequent differentiation into bone matrix. The abundance of these osteogenic cells within the site may be in part responsible for suppressing the activity of macrophages in Am site. However, the complete opposite was observed in the case of Bm implanted defects. Moreover, the secretion of matrix metalloproteinases (MMPs) by the macrophages infiltrated within the Bm implanted defects may have affected matrix properties of the implant, leading to the degradation of protein molecules.[40] Hence, one shortcoming of the present study is the lack of comparative evidence of the physical features of these matrices that can putatively modulate the immune response, such as wettability, surface roughness, and stiffness. However, innate immune responses to silk biomaterial are found to be largely governed by the protein conformation rather than by such physical parameters.[41]

5. Experimental Section Fabrication of Silk Fibroin Scaffolds: Am and Bm scaffolds were prepared following a published protocol.[42,43] The fibroin proteins were isolated as follows: (a) Am fibroin was collected by dissecting the mature live fifth instar larvae from Tasar Silkworm Farm of IIT, Kharagpur, West Bengal, India. To isolate fibroin, the silk glands were squeezed out and washed in deionized water to remove the traces of hydrophilic sericin. The gland fibroin protein was then solubilized in 1% (w/v) sodium dodecyl sulphate (SDS) solution in 10 × 10−3 M Tris (pH 8.0) and 5 × 10−3 M ethylenediaminetetraacetic acid (EDTA). Aqueous fibroin solution was obtained by dialyzing the above protein solution (MWCO 12000; Pierce, USA) against deionized water for 8 h, (b) For Bm, fresh cocoons collected from the Debra Sericulture Farm, West Midnapore, India were prepared following a standard protocol.[18] Briefly, Bm cocoons were cut into small pieces, boiled in 0.02 M Na2CO3 solution for 1 h, washed, and dried. The degummed silk fibroin fiber was then dissolved in 9.3 M LiBr (Merck, India) at 60 °C for 4 h before dialyzing against deionized water using Slide-a-Lyzer dialysis cassettes (MWCO 12000; Pierce) for 3 d. Each fibroin protein solution (2%) was separately casted in moulds, frozen at −20 °C, and subsequently lyophilized. The fabricated scaffolds were treated briefly with absolute ethanol to induce transition from random coil to β-sheet conformation, leading to insolubility of silk fibroin in aqueous medium. Physical Characterization: For physical characterization, SEM (Zeiss EVO50, Germany) was performed on the lyophilized scaffolds (n = 3/group) and the pore size was calculated from the micrographs using ImageJ software (NIH, Bethesda). The porosity measure was determined using the liquid displacement method.[44] In Vitro Degradation: Lyophilized scaffolds (n = 5) were sterilized in 70% ethanol overnight and subsequently exposed to 60 °C for 2 h to determine dry weights followed by incubation at 37 °C in phosphate buffered saline (PBS, pH 7.4) with or without 2 U mL−1 protease XIV (Sigma, St. Louise). The enzyme solutions were replaced (50%) daily over a period of 30 d. At each time point, the mass retention of the dried scaffold materials were taken after rinsing in distilled water, followed by drying at 60 °C until constant weights were obtained and compared with the initial dry weights. The percentage degradation was calculated by comparing initial dry weight of samples and weight loss due to material degradation. The scaffolds incubated in PBS without enzyme under the same conditions were served as controls. Colony Forming Unit: Primary osteoblasts were isolated from rats after collagenase 1 treatment and explant culture. Similarly, AD-MSCs were also isolated at NII. The cells were expanded from collagenasedigested stromal vascular fraction cells in culture medium consisting of Dulbecco’s Modified Eagle medium-F12 (DMEM-F12) Ham’s medium, 10% fetal bovine serum (FBS), 100 U ml−1 penicillin-streptomycin. For the experiment, six well tissue culture plates were coated with Am

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(2 wt%) or Bm (2 wt%) fibroin.[43] The cells [rat primary osteoblasts (passage 2, Figure S1A, Supporting Information) or adipose-derived mesenchymal stem cells (passage 5) were seeded (200 cells per well) and incubated in a hypoxic (5% O2) CO2 incubator for 7 d. Cells were fixed with 1 mL of ice-cold methanol for 30 min, followed by staining with 0.2% crystal violet solution for 2 h. After thoroughly washing the plates, the number of individual colonies (at least 30 cells per colony) was counted in each well (Figure S1B, Supporting Information). The experiment was repeated at three different times. Attenuated Total Reflectance-Fourier Transform Infrared (ATR-FTIR) Analysis: The ethanol treated scaffolds were dried in vacuum oven at 60 °C. The dried scaffolds were subjected to infrared spectroscopy in transmittance mode in the spectral range of 4000–400 cm−1 using Bruker Alpha P ATR-FTIR having Deuterated Triglycine Sulfate IR detector in transmission mode, with spectral resolution 4 cm−1 and number of scans 264. The IR spectra were deconvoluted in the region 1500–1700 cm−1 to generate insights about the composition of the secondary structure. The spectra were normalized and the relative area under each peak was used to determine the total crystallinity of the samples. Fourier self-deconvolution of spectra was performed using PeakFit Version 4.12 (Systat Software, Inc., CA). Fitting of the Fourier self-deconvolution (FSD) FTIR spectra with Gaussian profiles was done in the amide I region between 1595 and 1705 cm−1 and the β-sheet fraction was determined by integral of Gaussian profiles. The β-sheet content was estimated by considering the ratio of these bands to the total amide I bands. Bands between 1617 and 1637 cm−1 and between 1697 and 1703 cm−1 indicated the presence of β-sheet crystals.[45] Surgical Implantation of the Scaffolds: In order to compare the biodegradability of pure Am and Bm fibroin scaffolds in vivo, identical discs of ethanol-sterilized scaffolds were implanted in the peritoneum cavities of two groups of Lewis rats (n = 10), previously anesthetized by intermuscular injection of xylazine (11.5 mg kg−1) and ketamine (100 mg kg−1). The scaffolds were excised 0.5, 1, 3, 6, and 12 months postimplantation and processed for histological analysis by hematoxylin and eosin (H&E) staining. For bone remodeling studies, three to four weeks old Lewis rats were anesthetized as above and a calvarial bone defect model was established as reported earlier.[46] Briefly, a cranial skin incision was made along the midline of the skull bone. The subcutaneous tissue, musculature, and periosteum were shifted by blunt dissection and the calvarium was exposed. Bilateral, circular defects of 5 mm diameter were created with an electric drill rotating at slow speed with saline irrigation under a surgical microscope (Nikon, Japan). Scaffolds of 5 mm diameter and 1 mm thickness were cut by biopsy punch, sterilized by ethanol-treatment, air dried, and thoroughly washed by PBS immediately before implantation. Rats were divided into three groups (G): G1 (n = 9); controls, did not receive any scaffold, G2 (n = 9); received Am scaffolds, and G3 (n = 9); received Bm scaffolds. The scaffolds were implanted on both calvarial defects of rats, while in some rats one location was kept empty. After implantation, the cranial skin was repositioned and secured with 3–0 nylon sutures and covered with antibiotic impregnated dressing material to avoid contamination. Rats were kept in an isolator, fed with autoclaved acidified water, and irradiated food ad libitum. The implanted scaffolds were harvested after 1, 3, and 6 months of implantation. All experiments using rats were conducted following the protocols approved by the Institutional Animal Ethics Committee of the NII, New Delhi. X-Ray Imaging: Rat calvariae containing the scaffolds were harvested after 1, 3, and 6 months postimplantation and subsequently fixed in 10% buffered formalin prior to radiography. The calvarial defect areas were imaged using Kodak FX PRO Imaging System (Carestream Healths, Inc., Rochester, NY) to examine the extent of bone generation in implanted scaffolds. Three specimens were imaged for both Am and Bm implanted calvarial defects at each respective time point. Histology and Mineralization: Harvested implants (n = 3/group), 1 and 3 months postsurgery were dehydrated in ascending alcohol series, embedded in paraffin, and cut into 7 µm thick sections for H&E staining. Quantification of the total number of multinucleated giant cells was done by morphometric analysis of nuclei in 15 nonoverlapping

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fields of multiple scaffold sections. Only those cells comprising of >3 nuclei were considered as giant cells. To assess mineralization, the sections were separately stained for VK and ARS. For VK staining, the fixed tissue sections were deparaffinized, washed with distilled water, and subsequently stained with 1% silver nitrate solution for 1 h under ultraviolet light treatment. Next, following several changes of distilled water, specimens were treated with 5% NaS2O3 for 2 min to remove untreated silver and were given a final rinse of distilled water. For ARS staining (Sigma-Aldrich, St. Louise), deparaffinized sections were incubated in the stain solution (2%) at room temperature for 5 min, followed by washing with distilled water. To examine the collagen content in ECM of constructs, deparaffinized sections were stained with picrosirius red dye (1 mg mL−1) (Polysciences, Inc, Warrington, PA, USA) and counterstained with haematoxylin. The positively stained regions were delineated on the light micrographs using NIH Image J program to quantify the percentage collagen proportionate area which was measured in six randomly selected fields per image across multiple sections at 20× magnification. All the stained sections were examined under a bright filed microscope (Model IX51, Olympus Shinjuku, Tokyo) fitted with a DP70 digital camera using LCPlanFl 10×, 20×, and 60× objectives. Immunohistochemistry and Immunocytochemistry: Deparaffinized sections were processed for antigen retrieval by heat treatment with citrate buffer pH 6.0 at 90 °C for 15 min, followed by 3% (v/v) H2O2 treatment to deplete endogenous peroxidase. After several washes in PBS, the sections were incubated with primary anti-rat CD11b antibody conjugated with biotin (Biosciences, San Diego, CA) for 12 h at 8 °C. After primary antibody staining, the sections were given three consecutive washes with PBS containing 1% bovine serum albumin (BSA) and reacted with streptavidin conjugated horseradish peroxide (Jackson Immuno Research Laboratory Inc., West Grove, PA) for 1 h at room temperature. Finally, the sections were incubated with 3,3′-diaminobenzimidine (DAB) for 5 min (Vector Lab. Inc., Burlingame, CA) and counter-stained with hematoxylin, dried and mounted in DPX prior to imaging under a bright filed microscope (Olympus). For quantitative analysis, the number of CD11b+ cells per field (number of fields = 15) was manually counted from the stained sections for morphometric analysis. To confirm that the bone forming cells (osteoblasts) were isolated from the neonate bone, osteopontin staining was performed on cells (Methods, Supporting Information). Briefly, the 4% paraformaldehyde-fixed primary osteoblasts were cyto-spinned on poly-lysine coated glass slides. The cells were permeabilized using 0.1% Triton X-100 followed by incubation with osteopontin antibody (1:100 dilution; Santa Cruz Biotechnology, Santa Cruz, California) for 2 h and further incubated with Alexa Fluor 488 conjugated secondary antibody (1:400 dilution; Molecular Probes, Inc., Eugene, OR) for additional 45 min. The nuclei were stained with 4′, 6-diamidino-2-phenylindole (DAPI) and imaged using a fluorescence microscope (Olympus). Scanning Electron Microscopy (SEM): Both empty defects as well as explanted scaffolds were fixed in 4% paraformaldehyde for 4 h, osmicated in 1% osmium tetraoxide in sodium cacodylate buffer, and dehydrated in graded alcohol series. Following sample processing, the stub-mounted scaffolds were sputter-coated with gold and imaged using SEM (Zeiss EVO 50, Germany), at an accelerating voltage of 15 kV. Atomic Force Microscopy: The surface images of native calvarial bone, neo bone developed on Am matrix (6 months post-implantation) and control empty Am scaffold samples were recorded using an ARTIDIS (Nanosurf AG, Liestal, Switzerland) atomic force microscope. Standard triangular cantilevers DNP-S10 (Bruker, CA, USA) with sharp pyramidal tips (nominal tip radius 10 nm, nominal k = 0.06 Nm−1) were used for contact mode imaging of the surface. The spring constant of each cantilever was determined using the thermal noise method.[47] During measurement, each sample was kept immersed in PBS at room temperature. Imaging was done on several locations across the samples with a map size of 15 × 15 µm and a scan rate of 1 Hz. Images were analyzed using the Gwyddion (Petr Klapetek and David Necˇas, Brno, Czech Republic) software to enhance contrast and overlay topography and deflection data for 3D diagrams.

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Supporting Information Supporting Information is available from the Wiley Online Library or from the author.

Acknowledgements N.S. and P.B. contributed equally to this work. The work was funded by the Department of Biotechnology, Government of India, New Delhi, for which S.C.K. and A.M. are grateful. M.L. and F.M. were supported by Commission for Technology and Innovation (CTI) grant (Project 14540.2; 1 PFNM-NM). A.M., S.C.K., and S.G. acknowledge NII, IIT Kharagpur, and IIT Delhi for providing infrastructural supports to conduct this study. Received: April 17, 2015 Revised: May 16, 2015 Published online: June 17, 2015

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Statistical Analysis: Data were represented as mean ± SEM (standard error of mean). Student’s t-test was carried out to estimate the significance between the means of both groups and p value < 0.05 was considered as significant.

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