Open access peer-reviewed chapter
Abstract
Adult articular cartilage presents poor intrinsic capacity for regeneration, and after injury, cellular or biomaterial-based therapeutic platforms are required to assist repair promotion. Cartilage tissue engineering (CTE) aims to produce cartilage-like tissues that recreate the complex mechanical, biophysical and biological properties found in vivo. In terms of biomaterials used for CTE, three-dimensional (3D) self-assembling peptide scaffolds (SAPS) are very attractive for their unique properties, such as biocompatibility, optional possibility of rationally design cell-signaling capacity, biodegradability and modulation of its biomechanical properties. The most attractive cell types currently used for CTE are autologous chondrocytes and adult stem cells. The use of chondrocytes in cell-based therapies for cartilage lesions is limited by quantity and requires an in vitro 2D expansion, which leads to cell dedifferentiation. In the present chapter, we report the development of heparin-, chondroitin sulfate-, decorin-, and poly(ε-caprolactone)-based self-assembling peptide composite scaffolds to promote re-differentiation of expanded human articular chondrocytes and induction of adipose-derived stem cells to chondrogenic commitment.
Keywords
- 3D cell culture
- cartilage
- self-assembling peptide scaffold
- biomimetic materials
- tissue engineering
1. Introduction
Articular cartilage is an avascular connective tissue, composed of chondrocytes as practically unique cell type. Articular chondrocytes synthesize, maintain and remodel the highly specialized extracellular matrix (ECM) [1], which in turn allows to withstand the mechanical requirements of the joints [2]. It is currently believed that due to its avascular nature, cartilage tissue lacks an intrinsic capacity for regeneration in response to disease or injury, leading to long-term pain, degeneration and loss of function [3]. Cartilage tissue engineering (CTE) aims to produce cartilage-like tissue substitutes by combining the appropriate cells, scaffolds and bioactive molecules to assist repair cartilage lesions [4, 5].
Cell types currently used for CTE include autologous articular chondrocytes (ACh), which already possess the desired phenotype, and mesenchymal stem cells (MSC), from bone marrow (BMSC) or adipose tissue-derived (ADSC), which can be induced to undergo chondrogenic differentiation [6, 7]. Autologous chondrocytes would be the ideal cell source for cartilage repair due to their intrinsic properties regarding cell function and immune compatibility. However, cell accessibility from a patient biopsy is limited, and once isolated, chondrocytes needs to be extensively expanded in 2D monolayer [1]. During expansion process, chondrocytes rapidly undergo extensive loss of the original tissue-specific phenotype, downregulating the expression of chondrogenic markers, such as collagens and glycosaminoglycans while acquiring a fibroblast-like phenotype [8, 9].
Three-dimensional (3D) culture platforms are currently used to restore or maintain chondrogenic phenotype, since it recreates more closely the complex cellular microenvironment found
Synthetic hydrogels are good candidates for CTE since they possess unique properties, such as more than 95% of water content (which mimics the native cartilage ECM), biocompatibility and capacity of rationally design chemical signaling and biochemical properties. One of the best examples is the self-assembling peptide scaffold RAD16-I, commercially available as Puramatrix™. RAD16-I is a short peptide constituted by the sequence AcN-(RADA)4-CONH2, which alternates hydrophilic and hydrophobic amino acids (Figure 1A) [16]. The peptide undergoes self-assembly into a nanofiber network with antiparallel β-sheet configuration under physiological conditions (Figure 1B) [17]. The nanoscale architecture of the fiber network (around 10 nm diameter and 50–200 nm pore size) allows the cells to experiment a truly 3D environment (Figure 1C). Besides, biomolecules in such nanoscale environment diffuse slowly and are likely to establish a local molecular gradient. Non-covalent interactions allow cell growth, migration, contact with other cells, shape changes and a properly exposition of membrane receptors. Moreover, since stiffness can be controlled by changing peptide concentration these hydrogels can be tuned up to embed cells but not to entrap them [18].
Figure 1.
Peptide RAD16-I self-assembles into a nanofiber network. (A) Molecular model of peptide RAD16-I. Since the scaffold contains no signaling moieties, the environment is not instructive for cells. R = Arg; A = Ala; D = Asp. (B) Molecular model of the nanofiber developed by self-assembling RAD16-I molecules. The nanofiber is formed by a double tape of assembled RAD16-I molecules in antiparallel β-sheet configuration. (C) RAD16-I nanofiber network viewed by SEM. The nanoscale architecture of the fiber network) allows the cells to experiment a truly 3D environment white bar represents 200 nm. Adapted from Semino [
Since the peptide scaffold does not contain signaling motifs, the environment can be considered non-instructive, from the point of view of cell receptor recognition/activation. However, the self-assembling peptide scaffold RAD16-I can be functionalized by solid-phase synthesis by extending at the N-termini with signaling motifs, such as ECM ligands for cell receptors, to trigger different cellular responses [16, 19]. Several studies showed the capacity of RAD16-I to support cell maintenance of multiple cell types, including endothelial cells [20], hepatocytes [19, 21], fibroblasts [22], embryonic [23] and somatic stem cells [24, 25].
In the present chapter, we report the development of new bicomponent scaffolds based on the self-assembling peptide RAD16-I, for guiding chondrogenic differentiation of both adipose-derived stem cells (ADSC) and expanded dedifferentiated human articular chondrocytes (hAChs).
On one hand, we took advantage of the versatility of RAD16-I to specifically add molecular cues for guiding chondrogenesis in order to develop more biomimetic scaffolds. Thus, the first approach was based on the addition of heparin (Hep) moieties to the peptide scaffold, forming a stable electrostatic-based composite made of heparin-self-assembling peptide hydrogel. The advantage of this bicomponent scaffold is its natural capacity to retain heparin binding domain (HBD) containing growth factors (GFs), and thus, protecting them from degradation or denaturation [25]. Therefore, the non-instructive RAD16-I scaffold provides the structural 3D environment while the heparin moiety the binding structure to HBD-containing GFs. Our second approach, was based on mimicking the native articular cartilage ECM while providing signaling moieties presented in mature cartilage. Glycosaminoglycans (GAGs) and proteoglycans (PGs) are structural components of the native cartilage ECM and influence the regulation of cell proliferation, migration and differentiation [26]. In particular, chondroitin sulfate (CS, a sulfated GAG usually found as a constituent of PGs) and decorin (a small PG, consisting of a protein core linked to a GAG chain, consisting of chondroitin sulfate or dermatan sulfate) [27, 28] molecules were added to the RAD16-I scaffold by mixing the components, obtaining a chondroitin sulfate- and decorin-based self-assembling peptide composite scaffold.
Finally, we combined the self-assembling peptide RAD16-I with a woven poly(ε-caprolactone) (PCL). 3D weaving can be used to create porous structures arranged in multiple layers of continuous fibers in three orthogonal directions [29]. Such scaffolds were engineered with predetermined properties aiming to reproduce the mechanical features of native articular cartilage. Moreover, PCL is a Food and Drug Administration (FDA) approved biomaterial, biocompatible and biodegradable, widely used for medical applications [30, 31]. Our strategy was based on combine these two biomaterials to promote the attachment and differentiation of embedded cells, providing at the same time a biomimetic mechanical environment of the native mature cartilage [32].
2. Materials and methods
2.1 2D culture of ADSC and hACh
ADSC (PT-5006, Lonza) were cultured in 175 cm2 T-flasks (<6th passage) in Adipose-Derived Stem Cell Basal Medium (ADSC-BM) (PT-3273, Lonza) supplemented with Adipose-Derived Stem Cell Growth Medium (ADSC-GM) SingleQuots (PT-4503, Lonza). hACh cells (CC-2550, Lonza) were cultured at 10,000 cells/cm2 from passages 2–6 in 25, 75, and 175 cm2 T flasks. The growth medium consisted of Chondrocyte Basal Medium (CBM) (CC-3217, Lonza) plus SingleQuots of Growth Supplements (CC-4409, Lonza) containing R3-IGF-1, bFGF, transferrin, insulin, FBS, and gentamicin/amphotericin-B. Cultures were maintained in the incubator in humidified atmosphere at 37°C and 5% CO2.
2.2 3D culture of ADSC and hACh in RAD16-I composites scaffolds
ADSC and hACh 3D cultures were maintained under control or chondrogenic conditions. Control medium was prepared with DMEM High Glucose, GlutaMAX (61965, Gibco), ITS + Premix 100x (354352, BDBioscience), 100 U/mL Penicillin/100 μg/mL Streptomycin (P11-010, PAA), 40 μg/mL l-Proline (P5607, Sigma) and 1 mM Sodium Pyruvate (11360, Life Technologies). Cultures for chondrogenic differentiation were induced at day 2 with chondrogenic medium (control medium supplemented with 10 ng/mL TGFβ1 (GF111, Millipore), 25 μg/mL l-ascorbic acid 2-phosphate (A8960; Sigma) and 100 nM Dexamethasone (D8893; Sigma)). Chondrocytes were also cultured in expansion medium (see Section 2.1). 3D cell cultures were maintained in the incubator at 37°C and 5% CO2, and medium was changed every other day. Cultures were maintained for 4 weeks in the described serum-free media under control or chondrogenic conditions (in the presence of stimulating factors to induce chondrogenic differentiation) [33, 34]. After 4 weeks, 3D constructs were analyzed for morphology, gene and protein expression, glycosaminoglycans production and mechanical properties.
2.3 ADSC 3D culture in RAD/heparin composite scaffold
RAD16-I (PuraMatrix™, 354250, Corning) and composites RAD/Hep were prepared at a final concentration of 0.3% (w/v) RAD16-I. The composites were prepared by combining 95 μL of 0.5% (w/v) RAD16-I and 5 μL of heparin sodium salt solution (H3149, Sigma) in a concentration range between 0.01% and 1% (w/v). The mixture was then diluted with 10% sucrose (S0389, Sigma) to a final concentration of 0.3% (w/v) RAD16-I. To obtain RAD and RAD/Hep 3D cultures, ADSC were harvested by trypsinization and resuspended to 4 × 106 cells/mL in 10% sucrose. The 0.3% (w/v) RAD16-I peptide solution was mixed with the cell suspension (1:1) to obtain a final concentration of 0.15% (w/v) RAD16-I and 2 × 106 cells/mL. Then, 80 μL of the cell-peptide mixture (160,000 cells) was loaded into individual wells of a 48-well culture containing 150 μL of medium, which induced the self-assembly of the peptide. The plate was placed in the incubator for 20 min at 37°C and 5% CO2, and then 650 μL of fresh medium was added to the 3D cell cultures. ADSC 3D cultures were maintained during 2 days under control medium. Cultures for chondrogenic differentiation were induced at day 2 with chondrogenic medium.
2.3.1 ADSC and hACh 3D culture in RAD/CS and RAD/Dec composite scaffold
RAD16-I (PuraMatrix™, 354250, Corning) and composites RAD/CS and RAD/Decorin were prepared at a final concentration of 0.3% (w/v) RAD16-I. The composites were prepared by combining 95 μL of 0.5% (w/v) RAD16-I and 5 μL CS or Decorin at a concentration of 0.2% (w/v). The mixture was then diluted with 10% sucrose (S0389, Sigma) to a final concentration of 0.3% (w/v) RAD16-I. To obtain RAD16-I, RAD/CS and RAD/Decorin 3D cultures, cells were harvested by trypsinization and resuspended to 4 × 106 cells/mL in 10% sucrose. The 0.3% (w/v) RAD16-I peptide solution was mixed with the cell suspension (1:1) to obtain a final concentration of 0.15% (w/v) RAD16-I and 2 × 106 cells/mL. Then, 80 μL of the cell-peptide mixture (160,000 cells) was loaded into individual wells of a 48-well culture containing 150 μL of control or expansion medium, which induced the self-assembly of the peptide. The plate was placed in the incubator for 20 min at 37°C and 5% CO2, and then 650 μL of fresh medium was added to the 3D cell cultures.
2.3.2 hACh 3D culture in PCL, PCL/RAD and RAD scaffolds
In the case of PCL scaffold, a cell suspension of 25 × 106 cells/mL was seeded onto the surface of 5 mm × 0.75 mm woven PCL scaffolds (500,000 cells/scaffold). After 2 h, 100 μL of expansion or control medium were slowly added into the well and after 4 h, 700 μL were finally added. For PCL/RAD composites, cells were harvested and resuspended to 50 × 106 cells/mL in 10% (w/v) sucrose. Then, cells were equally mixed with 1% (w/v) RAD16-I and seeded onto the woven PCL scaffold disks (500,000 cells/scaffold). Then, 40 μL of expansion or control medium was added and the gel was spontaneously formed inside the PCL scaffolds, where the cells were embedded. After 30 min, 60 μL of medium was added in the well, and after 2 h, 700 μL was finally added. 3D cell cultures were maintained in the incubator at 37°C and 5% CO2, and medium was changed every other day.
3. Representative results
Autologous chondrocytes are one of the most attractive cell types for CTE, due to their intrinsic properties regarding cell function, since they are found in the native cartilage. Chondrocytes are characterized by a rounded morphology, the production of tissue-specific ECM components such as collagen type I and II and glycosaminoglycans (GAGs). One of the main challenges in CTE is to obtain enough cell mass to develop a tissue construct with the desirable biological and biomechanical properties. Particularly, articular chondrocytes are obtained by invasive techniques and cell number in patient biopsies is limited. Therefore, after isolation, chondrocytes need to be expanded in 2D monolayer [1]. The expansion process leads to a rapid downregulation of chondrogenic markers, such as Collagen type I (COL1) collagen type II (COL2) and Aggrecan (ACAN) [8, 9]. Moreover, the use of extensively passaged cells leads to some degree of hypertrophy, decreased biochemical content and compromised mechanical properties [1], which is not a good indication for cartilage substitute applications.
Mesenchymal stem cells (MSCs) have generated great interest as an alternative cell source to autologous chondrocytes. MSCs are pluripotent cells with a high proliferative capacity that can be differentiated, under the appropriate microenvironment, to numerous cell lineages, such as osteogenic, adipogenic and chondrogenic [35]. MSCs can be isolated from bone marrow, adipose tissue and other sources. In particular, the adipose tissue provides and abundant reservoir of mesenchymal stem cells (adipose-derived stem cells, ADSC), which can be obtained by non-invasive surgical techniques. ADSC can undergo chondrogenic commitment in the presence of TGF-β, ascorbate, and dexamethasone combined with a 3D culture environment [35].
Three-dimensional scaffold-based cell cultures are currently used in CTE to reestablish chondrogenic phenotype of dedifferentiated chondrocytes, since they mimic more closely the natural tissue environment. On the other hand, differentiation of ADSC to cartilage-like tissue has been achieved in various 3D scaffold systems such as alginate [36], agarose [37] and collagen [38]. We report here the development of new bicomponent scaffolds based on the self-assembling peptide RAD16-I, for guiding chondrogenic differentiation of both adipose-derived stem cells (ADSC) and expanded dedifferentiated human articular chondrocytes (hAChs).
3.1 Bicomponent scaffolds made out of heparin/self-assembling peptide hydrogels
In this section, we report the development of a nanofiber scaffold with growth factor binding affinity. The strategy consisted of adding heparin moieties to the RAD16-I peptide scaffold by mixing the two components, forming a stable composite hydrogel scaffold with a natural capacity to retain HBD-containing growth factors. To evaluate the functionality of this approach for CTE applications, ADSC were cultured in the new bicomponent scaffold and induced to chondrogenic differentiation using TGFβ-1, l-ascorbic acid 2-phosphate and dexamethasone as inductors in serum-free media. 3D cultures were maintained for 4 weeks in chondrogenic or control medium, and analyzed for proteoglycan production, protein expression and mechanical properties.
During ADSC culture in the peptide scaffold RAD16-I combined with increasing concentrations of heparin (RAD/Hep), constructs cultured under chondrogenic medium—unlike constructs under control medium—became highly stained with toluidine blue, indicating a significant production of proteoglycans (Figure 2A). This result correlated with the aggrecan (
Figure 2.
Chondrogenic capacity of ADSC in RAD/heparin composite scaffold. ADSC were encapsulated within the RAD16-I peptide scaffold combined with increasing concentrations of heparin and cultured for 4 weeks under control and chondrogenic medium. (A) Toluidine blue staining of 3D ADSC constructs cultured under control and chondrogenic medium. 3D construct view scale bars = 500 μm and section close up scale bars = 100 μm. (B) Aggrecan gene expression levels of chondro-induced ADSC. Constructs cultured with control medium did not express aggrecan after 4 weeks of culture. Ct values relative to ribosomal protein L22 (RPL22) were obtained and reported as fold increase (ΔΔCt) relative to 2D cultures. (C) Protein expression characterization of ADSC cultured in RAD/Hep composites and in 2D monolayer. Western blot results of collagen type I, II and X when ADSC were maintained in control and chondrogenic medium in RAD16-I scaffold and RAD16-I/Hep composites. Actin expression was used as an internal control. (D) Mechanical characterization of 3D constructs cultured for 4 weeks in chondrogenic medium compared to chicken and calf articular cartilage. ADSCs cultured with RAD16-I and RAD/Hep composite scaffolds were analyzed for storage modulus (G′, A), loss modulus (G″, B), complex modulus (G*, C) and tan(delta). Significant differences are indicated as * for p < 0.05, ** for p < 0.01, and *** for p < 0.001, one-way ANOVA, N = 2 n = 3). Adapted from Fernández-Muiños et al. [
Moreover, mechanical characterization was performed over 3D chondro-induced constructs. 3D constructs, presented a storage modulus (G′) in the same order of magnitude to chicken or calf articular cartilage, but the full mechanical response of the constructs was different from native cartilage as evidenced by tan(delta) (Figure 2D).
3.2 Bicomponent scaffolds made out of chondroitin sulfate or decorin and self-assembling peptide hydrogels
The next strategy was based on mimicking the native cartilage ECM by adding chondroitin sulfate or decorin molecules to the nanofiber scaffold, generating thus chondro-favorable biochemical cues in the 3D environment. Previous work has evaluated the influence of CS to guide chondrogenesis in different hydrogel scaffolds such as chitosan [39], PEG [40], or collagen type I [41], but less is known about the ability of decorin to promote chondrogenesis commitment. In the present work, we studied the influence of both CS and decorin molecules on chondrogenesis in a nanometric 3D system. The capacity of these bicomponent scaffolds to foster chondrogenic differentiation was evaluated in two different scenarios: re-differentiation of expanded hAChs and induction of ADSC to chondrogenic commitment. Cells were seeded in RAD16-I/CS, RAD16-I/Dec and RAD16-I scaffold alone and maintained for 4 weeks in chondrogenic or control medium. Moreover, chondrocytes were also cultured in expansion medium, which contains GFs that could affect the fate of the 3D culture. 3D constructs were analyzed for morphology, gene and protein expression, proteoglycan synthesis and mechanical properties.
SEM images were obtained at week 4 of culture to assess cell morphology and their interaction with each scaffold (Figure 3). Articular chondrocytes cultured in expansion medium possessed a spherical morphology with possible cell-matrix interactions and thorough ECM components. Nanofibers and putative matrix components were detected on the surface of constructs cultured in control medium. Moreover, grooves with visible fibers were observed on the surface of constructs cultured in chondrogenic medium, fact that suggested the presence of secreted matrix components. On the other hand, adipose-derived stem cells under chondrogenic induction looked elongated and anchored to the scaffold surface, while nanofibers and possible ECM components synthesized by the cells were observed in control medium (Figure 3). No significant differences in cell morphology were detected between RAD, RAD/CS or RAD/Dec scaffolds in any cell type.
Figure 3.
SEM images of hACh and ADSC at week 4 of culture in RAD, RAD/CS and RAD/Dec scaffolds. Two images per condition are shown. Adapted from Recha-Sancho and Semino [
Chondrogenic markers expression were studied at gene and protein level in hACh 3D constructs cultured in chondrogenic and expansion medium, and compared to their 2D counterparts.
Figure 4.
Chondrogenic capacity of dedifferentiated hACh in RAD/CS and RAD/Dec 3D composite scaffolds. hACh were encapsulated within the RAD16-I peptide scaffold combined with chondroitin sulfate and decorin, and cultured for 4 weeks under expansion, control and chondrogenic medium. (A) Gene expression levels of chondrogenic and hypertrophic markers. hACh were analyzed by qRT-PCR for collagen type I (
Interestingly,
Toluidine blue staining was performed in hACh 3D constructs to qualitatively assess the production of GAGs. Constructs under chondrogenic induction became highly stained, indicating a significant production and accumulation of GAGs by the cells (Figure 4C). Constructs cultured under expansion medium showed less staining, while constructs under control medium became weakly stained.
The mechanical properties of hACh 3D constructs cultured under chondrogenic medium were assessed at week 4 by dynamic mechanical analysis (DMA) and compared to calf and chicken articular cartilage (Figure 4D). hACh constructs exhibited lower storage modulus values (G′) than did the native cartilage samples. The viscous components (G″) and the complex modulus (G*) displayed a more similar tendency to cartilage controls. Nevertheless, all samples presented G′ values higher than G″ values, indicating that the constructs were more elastic than viscous. Tan(delta) showed that 3D constructs were comparable to chicken cartilage but differed from calf cartilage.
Chondrogenic and hypertrophic markers were studied in ADSC 3D constructs in the three scaffold types and compared to 2D monolayer culture. Results show that the gene expression of
Figure 5.
Chondrogenic capacity of ADSC in RAD/CS and RAD/Dec 3D composite scaffolds. ADSC were encapsulated within the RAD, RAD/CS and RAD/Dec composite scaffolds and cultured for 4 weeks under control and chondrogenic medium. (A) Gene expression levels of chondrogenic and hypertrophic markers. ADSC were analyzed by qRT-PCR for collagen type I (
Chondro-induced ADSC produced sulfated glycosaminoglycans, as reveals the intense staining by toluidine blue (Figure 5C, up). No calcium mineralization, an indicator of hypertrophy, was detected by Von Kossa staining (Figure 5C, down).
The mechanical properties of ADSC cultured under chondrogenic conditions in RAD, RAD/CS and RAD/Dec were assessed by dynamic mechanical analysis (DMA) at week 4 (Figure 5D). The constructs presented a storage modulus (G′), viscous component (G″) and complex modulus (G*) closely related to chicken and calf cartilage. However, samples presented values of G′ much higher than G´´ so that the constructs were more elastic than viscous. Tan(delta) showed that the full mechanical response of the constructs was very similar to chicken cartilage but differed from calf cartilage.
In the present work, we aimed to induce chondrogenesis differentiation of both expanded hACh and ADSC in 3D bicomponent scaffolds made out of chondroitin sulfate or decorin and self-assembling peptide hydrogels. The expression of chondrogenic markers such as
3.3 Bicomponent scaffolds made out of PCL and self-assembling peptide hydrogels
Self-assembling peptide hydrogels provide a soft and permissive microenvironment, allowing cells to migrate, extend cellular processes and contact with other cells. Nevertheless, the use of soft hydrogels for CTE can be challenging due to its low stiffness. One approach to address this issue is the use of composite scaffolds, comprising a microscale component to increase mechanical properties and a hydrogel component (of nanoscale dimension) to promote chondrogenesis. Woven 3D poly(ε-caprolactone) (PCL) resemble native cartilage mechanical properties and, due to its high wettability, can be infiltrated with a hydrogel matrix, such as fibrin, alginate, and poly-acrylamide [43, 44, 45]. In this study, we developed a unique composite scaffold by infiltrating a 3D woven microfiber poly(ε-caprolactone) scaffold with the RAD16-I self-assembling peptide nanofiber to obtain a multi-scale functional cartilage-like tissue. The chondrogenic capacity of this new bicomponent was evaluated with expanded dedifferentiated human articular chondrocytes.
The high wettability properties of the PCL scaffold (Figure 6A) allowed to easily introduce the cells suspended in the RAD16-I peptide solution between the interweaving fibers of PCL scaffold (Figure 6B, left). Areas of RAD16-I peptide deposition could be observed within the organized woven morphology of the fiber scaffold (Figure 6B, right). Thus, cells were seeded in the composite PCL/RAD and in the two scaffolds independently, PCL and RAD, and maintained for 4 weeks in expansion, chondrogenic and control medium. 3D constructs were analyzed for morphology, gene and protein expression, proteoglycan synthesis and mechanical properties.
Figure 6.
SEM characterization of PCL/RAD and hACh constructs. (A) Water (left) and 0.5% RAD16-I solution (right) contact angle. The liquid was totally absorbed by the PCL scaffold (contact angle << 90°), indicating high wettability. (B) Surface view of PCL and PCL/RAD structure by SEM. 0.5% RAD16-I was lyophilized within the PCL scaffold. (C) hACh at week 4 of culture in PCL, PCL/RAD and RAD scaffolds. hACh were seeded in each scaffold and cultured in expansion, control and chondrogenic medium. Two images per condition are shown. Adapted from Recha-Sancho et al. [
In order to evaluate cell morphology and their interaction with the scaffolds, SEM images of hACh cultured in PCL, PCL/RAD and RAD 3D scaffolds in expansion, control and chondrogenic medium were taken at week 4 of culture (Figure 6C). hACh seeded in PCL scaffolds looked elongated and growing on the surface of PCL fibers. Interestingly, more fibers were detected under chondrogenic induction, probably due to an increase in extracellular matrix components production by the cells. In PCL/RAD constructs, cells seemed to be attached to the PCL fibers, with a more spherical morphology than in PCL scaffold alone, while hACh in RAD scaffolds presented in general a spherical shape.
Chondrogenic and hypertrophic markers were studied at gene and protein level at week 4 of culture in 3D scaffolds and compared to 2D cultures.
Figure 7.
Chondrogenic capacity of dedifferentiated hACh in PCL, PCL/RAD and RAD scaffolds. hACh were seeded in each scaffold, and cultured for 4 weeks under expansion, control and chondrogenic medium. (A) Gene expression levels of chondrogenic and hypertrophic markers. hACh were analyzed by qRT-PCR for collagen type I (
The expression of
The production of sulfated glycosaminoglycans was qualitatively assessed by toluidine blue staining. Constructs under chondrogenic medium were the most strongly stained compared to expansion and control medium (Figure 7C).
Mechanical properties of the scaffolds alone and hACh 3D constructs were assessed by dynamic mechanical analysis (DMA) at week 4 of culture, and compared to chicken and calf articular cartilage (Figure 8). The elastic component (G′, storage modulus) of scaffolds and 3D cultures was significantly lower than values of chicken and calf cartilage. Regarding the viscous component (G″, loss modulus), 3D constructs differed from calf native cartilage, while only PCL cellular scaffolds presented differences with chicken cartilage. All samples presented G′ values higher than G″ values, meaning that the material was more elastic than viscous. Because the complex modulus (G*) is the sum of both components, G* basically corresponds to the elastic component in this case and it presented the same pattern as the storage modulus (G′). Concerning tan(delta), which is the full mechanical response of the material, the scaffolds and cell constructs were closely related to both native cartilages, with exception of RAD constructs in chondrogenic medium, which presented differences with calf cartilage. Moreover, differences were observed between PCL/RAD and RAD constructs under the same medium. The combination of PCL scaffold and RAD hydrogel changed their viscoelastic nature after 4 weeks of culture with hACh, since tan(delta) values of the composite were increased compared to RAD scaffolds alone. This effect was not observed between composites PCL/RAD and PCL scaffold alone.
Figure 8.
Mechanical characterization of scaffolds alone and 3D constructs cultured for 4 weeks in expansion, control and chondrogenic medium compared to chicken and calf articular cartilage. hACh cultured in PCL, PCL/RAD and RAD scaffolds were analyzed for storage modulus (G′), loss modulus (G″), complex modulus (G*) and tan(delta). Significant differences are indicated as * for p < 0.05, ** for p < 0.01, and *** for p < 0.001, one-way ANOVA, N = 2 n = 3). Adapted from Recha-Sancho et al. [
In the present study we report the chondrogenic capacity of dedifferentiated hACh in a composite scaffold comprising a microscale woven 3D poly(ε-caprolactone) and the peptide nanofiber scaffold RAD16-I. PCL scaffold resembles native cartilage mechanical properties while the RAD16-I hydrogel provides a soft and permissive 3D environment. The expression of chondrogenic markers such as
4. Conclusions
We report evidences from our previous work which indicates the chondro-inductive capacity of newly developed biomaterials including heparin-, chondroitin sulfate-, decorin-, and poly(ε-caprolactone)-based self-assembling peptide composite scaffolds. In particular, we demonstrated that these biomimetic biomaterials fostered re-differentiation of expanded human articular chondrocytes as well as adipose-derived stem cells into chondrogenic lineage commitment. Moreover, both biological and biomechanical properties obtained of these cartilage substitutes were comparable to natural samples of chicken and calf counterparts. This clearly suggest that these newly class of biomaterials are promising for their future application in reparative and regenerative medicine platforms.
Acknowledgments
This research was supported by funding from the AO Foundation in Acute Cartilage Injury Collaborative Research Program (ACI CRP) under the project “Bioactive and biomimetic scaffolds for cartilage regeneration” (BIOCART) and OSTEOCHON3D, and grants from the NIH (AR50245, AR48852, AG15768, AR48182, AG46927).
Abbreviations
| 2D | two dimensional |
| 3D | three dimensional |
| ADSCs | adipose-derived stem cells |
| BMSCs | bone marrow mesenchymal stem cells |
| CS | chondroitin sulfate |
| CTE | cartilage tissue engineering |
| Dec | decorin |
| DMA | dynamic mechanical analysis |
| ECM | extracellular matrix |
| GAGs | glycosaminoglycans |
| G′ | storage modulus |
| G″ | loss modulus |
| G* | complex modulus |
| hACh | human articular chondrocytes |
| HBD | heparin binding domain |
| Hep | heparin |
| MSCs | mesenchymal stem cells |
| PCL | poly(ε-caprolactone) |
| PGs | proteoglycans |
| PGA | polyglycolic acid |
| PLGA | poly(lactic-co-glycolic acid) |
| PLA | polylactic acid |
| SAPS | self-assembling peptide scaffolds |
| SEM | scanning electron microscope |
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