**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 cartilagelike 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 *in vivo* [10, 11]. In terms of biomaterials used for CTE diverse possibilities in composition, structure, biodegradability and biomechanical properties exist. In general, biomaterials user for tissue engineering applications can be classified into natural or synthetic scaffolds. Natural scaffolds are commonly hydrogels made of natural materials such as Matrigel™, collagen type I, laminin and gelatin, which provide chemical cues, principally ECM binding motifs. However, due to its natural origin, they frequently contain undefined amounts of different constituents such as growth factors and cytokines which would be the main responsible of presenting variability from batch to batch [10]. Thus, due to its complex composition possible modifications to improve them are limited. On the other hand, synthetic scaffolds have minimal variation from batch to batch production, providing a reproducible cellular microenvironment. Moreover, they present lower biodegradability *in vitro*, fact that permits to maintain structural and mechanical properties for longer periods of time. Alike natural scaffolds, structural properties, such as matrix stiffness, can be modulated by increasing concentration. In the last decades, polymeric scaffolds, such as poly(lactic-co-glycolic acid) (PLGA) [12] and polylactic acid (PLA) [13, 14] as well as synthetic peptide nanofibers [15] have been developed to culture cells in 3D. Clinically used scaffolds are collagen type I/III and hyaluronic acid-based biomaterials, and others under consideration are for instance injectable fibrin gels, collagen type I or II and sponges, polylactic acid (PLA) and polyglycolic acid (PGA). As today, however, the best CTE product does not maintain their tissue properties after implantation, and the minimal medical standards are not yet achieved.

 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].

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 *Cartilage Tissue Engineering Using Self-Assembling Peptides Composite Scaffolds DOI: http://dx.doi.org/10.5772/intechopen.83716* 

#### **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 [17].* 

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 HBDcontaining 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].
