A multidisciplinary perspective on the latest trends in artificial cartilage fabrication to mimic real tissue

Applied Materials Today 29 (2022) 101603 Available online 27 July 2022 2352-9407/© 2022 The Author(s). Published by Elsevier Ltd. This is an open access article under the CC BY-NC-ND license (http://creativecommons.org/licenses/bync-nd/4.0/). A multidisciplinary perspective on the latest trends in artificial cartilage fabrication to mimic real tissue Amir A. Elhadad a , Ana Alcudia b , * , Bel´ en Begines b , Eva M. P´ erez-Soriano a , Yadir Torres a a Departamento de Ingeniería y Ciencia de los Materiales y del Transporte, Universidad de Sevilla, Escuela Polit´ ecnica Superior, Calle Virgen de ´ Africa 7, Sevilla E-41011, Spain b Departamento de Química Org´ anica y Farmac´ eutica, Facultad de Farmacia, Universidad de Sevilla, C/ Profesor García Gonz´ alez, 2, Seville E-41012, Spain ARTICLE INFO Keywords: Cartilage Osteoarthritis Osteochondral defects Polymer Scaffold Tissue engineering ABSTRACT Articular cartilage (AC) has a low ability for endogenous repair, and current treatments are not completely efficient in generating strong and healthy repair tissue. As a result, there is a great deal of interest in cartilage repair and recovery of joint function due to the aging population and the increasing frequency of cartilage degradation and osteoarthritis. Thus, the search for better and more efficient strategies to regenerate articular cartilage to repair the joints in patients with osteoarthritis is a critical motivation. Biomaterials, which are made of natural or synthetic polymers, or a mixture of the two, serve as a scaffold for tissue regeneration. Cartilage tissue engineering (TE) seeks to develop a treatment that can effectively repair damaged AC caused by trauma or disease. In cartilage TE, scaffolds that provide a 3D environment like native articular cartilage for cells to proliferate, attach, and hence ensure tissue regeneration success are needed. For optimal cell activities, which is associated with proper cartilage regeneration, these scaffolds must meet all necessary criteria, including adequate architectural, physicochemical, and biological properties. This study briefly explores the evolution of scaffolds constructed of synthetic polymers and hybrid materials for cartilage tissue engineering. Initially, general information on AC is presented, as well as a review of medical approaches to treating cartilage injuries or disease. Following that, the requirements for scaffolds in tissue regeneration, specific materials, and techniques for producing them are briefly discussed. 1. Introduction The primary function of AC is to smooth bone surfaces and produce a lubricated joint surface, allowing for low-friction load transmission [1]. Therefore, a healthy smooth cartilage surface is essential for pain-free synovial joint mobility [1]. However, because the AC lacks blood vessels and has no nutrient supply, damaged lesions take a long time to heal. AC degeneration is a significant problem around the world; in Europe, over the age of 60, 9.6% of men and 18% of women are influenced and experience symptomatic osteoarthritis (OA) [2]. People with symptomatic OA have mobility impairments in 80% of cases, 25% are unable to perform their primary daily tasks [3]. In 2008, OA was among the most costly diseases treated in US clinics, costing $40 billion in total national hospital expenses, accounting for 3.5% of national hospital bills and 70% of all arthritis-related inpatient hospitalizations, primarily due to arthroplasty surgery [4]. Several studies have recently associated symptomatic hip or knee OA with a higher risk of all-cause mortality [5–7]. Damage to the cartilaginous layer of the joint has been reported to cause serious morbidity for hundreds of years [8], and yet despite the substantial research already ongoing on the subject, effective healing remains a significant issue for clinicians. All these studies suggest that symptomatic OA has a significant socioeconomic impact. The goal of this study is to provide an overview of articular cartilage, its structure, diseases, and current treatment options. The prerequisites and materials for scaffolds in cartilage TE, as well as general techniques for scaffold preparation, are then described. The last section of the article discusses synthetic polymer scaffolds and their pairings with natural materials (hybrid scaffolds) as well as 3d-bioprinting technology. 2. Structure and function of articular cartilage AC represents flexible avascular connective tissue that wraps the articulating ends of bones in the synovial joint. It is a dense structure of * Corresponding author. E-mail address: [email protected] (A. Alcudia). Contents lists available at ScienceDirect Applied Materials Today journal homepage: www.elsevier.com/locate/apmt https://doi.org/10.1016/j.apmt.2022.101603 Received 14 May 2022; Received in revised form 14 July 2022; Accepted 20 July 2022 Applied Materials Today 29 (2022) 101603 2 collagen and elastic fibres embedded in a firm gelatinous substance and characterized by a high fluid content and low oxygen supply [9]. It differs from the bone in various respects; it is an avascular kind of supportive connective tissue with a less ordered microarchitecture than bone [10]. AC withstands and redistributes the compressive, tensile, and shear stresses that occur during joint articulation [11], allowing the tissue to last for tens of millions of loading cycles during the course of a lifetime [10]. AC is made up of a dense extracellular matrix (ECM) that is randomly interspersed with highly specialised cells called chondrocytes and collagen [1]. Water (60–85% by weight), a crosslinked network of type II collagen (15–22%), and proteoglycans (4–7%) are the main components of cartilage ECM, non-collagenous proteins, and glycoproteins are also present, although in lesser amounts [12]. These components work together to maintain water within the ECM, which is essential to retain its mechanical characteristics. The avascular tissue AC has a thickness of 3 to 4 mm and is made up of four layers: a superficial zone (SZ), a middle zone, a deep zone, and a calcified zone (Fig. 1a). Each zone is responsible for reducing the imposed weight or sliding motion [13]. Collagen fibrils cells (mainly type II and IX collagen) are firmly packed and oriented parallel to the articular surface in the upper 10% of cartilage thickness (thin superficial zone) (Fig. 1b). The integrity of this layer, the orientation of collagen fibres, are critical for protecting the AC from severe tensile loads [9,14]. The middle (transitional) zone is adjacent to the superficial zone and acts as an anatomical and functional link between the two. It is the thickest region, accounting for 40–60% of the overall thickness of the AC volume [14]. In this layer, glycosaminoglycan-containing ECM surrounds chondrocytes and collagen fibrils (GAGs) [15]. The deep zone, which contacts the subchondral bone, represents around 30% of the AC volume. The deep zone is in charge of supplying the most compressive force resistance due to its high PG content and the fact that the collagen fibrils are positioned perpendicular to the articular surface. This layer is made up of hydroxyapatite (HAp), which acts as a bone anchor and is coupled with collagen and chondrocytes to form the articular surface in a vertical direction [15]. Proper lubrication is one of the most important components for optimal articular cartilage performance, as it protects cartilage from wear, but a reduction in lubrication leads to cartilage degeneration [16,17]. Healthy AC is essential for joint operation because it offers the low friction and load-bearing surface that joint articulation requires [18]. Nonetheless, individuals with joint injuries are more likely to have faster cartilage degradation and wear, as well as joint dysfunction owing to post-traumatic osteoarthritis (PTOA) [19]. Synovial fluid is a biochemical lubricant that, due to its biochemical components, has shear-thinning properties (Fig. 1c) [20]. As a pressurised interstitial fluid or boundary lubricant, it facilitates low friction during joint movement that affects shear kinematics and cartilage sliding in response to loading and/or motion [21]. AC can support imposed loads over its lifespan, which is anticipated to be 100–200 million loading cycles [22]. In the human body, cartilage comes in a variety of forms, and these variations affect the structure and function of the cartilage (Fig. 2). 2.1. Hyaline cartilage The most widespread form of cartilage in adults is hyaline cartilage [23]. It has a smooth surface and a pale blue-white color that makes it easier for tissues to slide/glide while also offering flexibility and support. There are no nerves or blood vessels in hyaline cartilage, which is mainly primarily made of type II collagen and proteoglycans [23,24]. The hyaline cartilage is generally moist on the surface, but as it ages, it becomes dryer, thinner, and more yellow [23,24]. 2.2. Elastic cartilage Elastic cartilage contains collagen and a large amount of elastin, which are arranged into a threadlike matrix in which dispersed chondrocytes reside [25]. It has a typical dull yellowish hue and greater elasticity than hyaline and fibrous cartilage due to its high elastin content [26]. The principal protein is elastin, which is mostly found in structures that are commonly deformed; for example, the larynx, ear, epiglottis, and eustachian tube [27]. Fig. 1. (a) The relative positions of the cartilage layers on the joint surface; (b) Major ECM components and the distribution of chondrocytes. Left: Zonal distribution of proteoglycans, counterions, as well as water molecules. Right: The structure of collagen fibres, as well as the shape and arrangement of chondrocytes; (c) The human femorotibial knee is shown in two states: unloaded (up) and loaded (down). A.A. Elhadad et al. Applied Materials Today 29 (2022) 101603 3 2.3. Fibrocartilage Fibrocartilage has a higher proportion of Type I collagen fibres and a lower proportion of proteoglycan [26] if compared to other cartilage tissues. As a result, it can withstand high levels of tension and compressive forces, making it an excellent connection in high-stress locations of the body such as between bones, ligaments, and tendons [26]. 3. Cartilage diseases/injury Cartilage may be damaged directly by disuse and overuse, as well as impact forces induced by falls, vehicle accidents, and sports injuries. Joint discomfort, inflammation, swelling, stiffness, or grinding around the joints, are all possible symptoms. Damage, on the other hand, may be severe enough to cause temporary disability; but, if impact forces exceed a certain threshold, irreversible damage of cartilage tissue with a total disability may occur [28,29]. Irreversible AC damage is the hallmark of osteoarthritis (OA) and a chronic systemic inflammatory disease (rheumatoid arthritis (RA)) [28]. The most common form of joint disease, OA, is accompanied by joint discomfort and function loss., which can lead to disability. Primary osteoarthritis and secondary osteoarthritis are the two types of OA. Primary osteoarthritis is a kind of articular degeneration that has no obvious reason. Untreated cartilage lesions cause the articular surfaces to become fibrillated and rough and can cause OA (Fig. 3) [30] that affects shear deformation during joint movement [31]. Secondary OA develops due to a pre-existing medical disease, overstress from physiological activity or obesity, abnormal force distribution across the joint, infection, or injury [32–34]. AC breaks down as a result of OA, causing the joint bones to rub against each other, resulting in severe pain, stiffness, as well as movement impairments [33]. The healing process of injured lesions is prolonged due to the avascularity of AC tissue and a lack of nutrition supply [1]. Depending on whether the damage only affects the AC layer or affects the cartilage and some of the underlying bone, injuries to the cartilage surface are classified as chondral defects or osteochondral defects (Fig. 5) [35]. Despite their distinctions, both types of defect lead to similar symptoms and require Fig. 2. Microscopic views of the three forms of cartilage (elastic, hyaline, and fibrous; up) and disposition in the human body (down). Fig. 3. Detailed anatomy of a normal joint, an osteochondral defect and osteoarthritis in the knee. A.A. Elhadad et al. Applied Materials Today 29 (2022) 101603 4 complete joint arthroplasty as a final treatment option if they are not properly treated [35,36]. 4. Cartilage regeneration Various treatments have been found to be effective, including marrow stimulation (e.g., microfracture, drilling, and abrasion arthroplasty), AC grafts, and cell-based therapies [37,38]. Despite their merits, these approaches have certain drawbacks, which will be discussed in the following sections. 4.1. Bone marrow stimulation techniques Bone marrow stimulating techniques are the most popular employed chondral defect repair method, including, e.g., microfracture, drilling, and abrasion arthroplasty [39]. Microfracture is a method to gain access to bone marrow mesenchymal stem cells (BMSCs) by using an awl to make small micro-holes in the subchondral bone under the defect, causing bleeding [39,40]. However, it induces fibrocartilage to fill the cartilage defect, which is less durable and more prone to re-injury than native hyaline cartilage [41]. Pridie drilling, which uses a drill instead of an awl, has the disadvantage of generating thermal osteonecrosis. Due to the heat generated by drills and high-speed burrs, the subchondral bone may be damaged. [37]. However, the fibrocartilage structure formed by either of these techniques is mechanically inferior to the adjacent hyaline cartilage and frequently fails over time [42]. Abrasion arthroplasty is a minimally invasive debridement (removal) that includes drilling the subchondral bone, which is located below the cartilage and triggers the bleeding and healing response [43]. Over the debrided area, a blood clot is formed, transforms into fibrous tissue and replaces the worn-out articular cartilage [43]. 4.2. Osteochondral grafting (OCG) Osteochondral grafting (OCG) is a method of treating cartilage injuries that expose underlying bone [44]. The OCG is commonly used by orthopedic surgeons to attack healthy tissues and deliver a small amount of transplant material. Common complications, such as postoperative infection and higher patient mortality, may restrict the use of autografts after bone transplantation [45]. Osteochondral allografts from cadaver donors are a popular alternative to autografts, but they bear the risk of disease transmission, and immunogenic reactions, as well as a higher cost [46]. These restrictions have prompted researchers to consider the use of biomaterial scaffolds as a possible alternative tissue engineering (TE) approach. 4.3. Cartilage tissue engineering The science world has recently been interested in cartilage tissue engineering [47]. Because AC is avascular, chondrocyte migration is restricted, which decreases self-recovery [1,11,30]. As a result, the basis of artificial cartilage inspires researchers to develop and produce materials that mimic biological cartilage. Artificial cartilage structures consisting of hydrogels or polymers that mimic the functional and tribological responses of native AC tissue are used to repair articular cartilage injuries. Owing to the complexity of tissue and its specific tribomechanical characteristics, developing materials for cartilage replacement presents various challenges, including selecting suitable cell sources and scaffolds, building biomechanically adequate tissues, and integrating them into native tissue [38]. During the proper growth of living tissue, cells adapt their biological activity and behavior in response to mechanical stimuli. Biomechanical signals control the physiological response of cells through a mechanism known as mechanotransduction [48]. Additional characteristics that distinguish AC include surface lubrication and a low coefficient of friction, which are necessary for engineered cartilage to retain the tissue’s bulk mechanical integrity [49]. 4.3.1. Cell-mediated osteochondral defect repair For the regeneration of the joint cartilage surface, a variety of cell types are being studied. a) Chondrocytes The AC is mainly made of a dense extracellular matrix (ECM) that is randomly interspersed with highly exclusive cells termed chondrocytes [1]. These cells reside in the ECM and operate along with it to keep cartilage tissue functioning. Because of these characteristics, chondrocytes are key cells in the development of cell-based cartilage repair and regeneration technologies. Articular chondrocytes (Fig. 4) are obtained from the non-load-bearing area of the knee joint and expanded in vitro for approximately four to six weeks to obtain enough cells before being implanted. Eventually, after enough cultured and expanded chondrocytes have been obtained, they are applied to the damaged region. In their new environment, transplanted cells proliferate and deliver new articular cartilage. However, limited biomechanical efficacy prior to autologous chondrocyte implantation (ACI) regeneration, cell leakage from the graft site and fibrocartilage formation are also concerns [50,51]. As a result, research on different cell sources is continuing in the search for excellent outcomes. a) Mesenchymal Stem Cells (MSCs) Because of its inherent chondrogenic characteristic, accessibility, cell homing capacity, and immunomodulatory role, mesenchymal stem cells (MSCs) have emerged as a promising cell source for AC tissue engineering (Fig. 5) [52]. They are distinguished by their long-term self-- renewal ability, fibroblast-like morphology, ability to differentiate along many cell lineages, including adipocytes, chondrocytes, or osteocytes, and expression of specific surface molecules [9,53]. MSCs are collected from a variety of sources, including bone marrow, fatty tissues, and the umbilical cord, and then injected into defective areas of cartilage either immediately (route 1), after collection (i.e., fresh non-cultured and expanded cells) or amplified in vitro and cultured for 2–4 weeks (route 2), and then injected into the damaged zone of the cartilage tissue [54]. In route 3, the isolated cells are amplified and activated toward chondrogenesis using growth factor and then injected into the damaged zone of the cartilage tissue. In fact, new cartilage formed by MSCs implanted in cartilage defects has been observed to be too thin to resemble mature cartilage [55]. The results also indicated that the defective areas were repaired by fibrocartilaginous tissue as well as the high cost and time consuming of such a treatment [54]. Because of these limitations, it is challenging to establish appropriate standards for the use of MSCs in cartilage TE. 4.3.2. Scaffolds for cartilage treatment Scaffolds are used to provide biological and physical support for cell formation and transplantation into an organism. Tissue bioengineering (TE) combines the use of a temporary supporting structure (scaffold) that offers a favourable environment for cells and ensures tissue regeneration success [56]. In addition to mechanical considerations, this scaffold is seeded with cells with the addition of growth factors that allow cells to differentiate and produce tissue-specific proteins [57]. According to O’Brien [58], ideal scaffolds must meet several design criteria for tissue engineering. •Cell attachment, proliferation, and maintenance of differentiated cell signaling should all be achievable. •The scaffolds must be biodegradable without inducing any inflammation or toxicity in vivo induced by the polymer or its degradation breakdown by-products. These by-products must be able to be removed from the body without interfering with other organs. A.A. Elhadad et al. Applied Materials Today 29 (2022) 101603 5 •The scaffold should biodegrade at a controlled rate to allow new tissue to grow in the implantation site. •The mechanical integrity of the scaffold that has been implanted must be maintained. The ability of any scaffold to allow cell infiltration and vascularization is dictated by a combination of adequate mechanical characteristics and a proper porosity design that allows for cell adhesion. Cell ingrowth and matrix formation, along with mass transport of nutrients and waste materials into and out of the tissue, are all facilitated by the scaffold’s porosity. •These scaffolds must also withstand physiological parameters such as pH and body temperature. Along with the above-mentioned criteria, cartilage tissue engineering has centered on the bulk mechanical properties of natural cartilage, such as compressive aggregate modulus and tensile strength [59]. In general, it requires the choice of the appropriate scaffold, control of cell differentiation, and the application of mechanical limitations tailored to the desired tissue type [57]. There are various materials that can be used to create scaffolds, including natural polymers such as collagen chitosan, silk fibroin, alginate, cellulose, gelatin, and hyaluronic acid (HA) as well as synthetic polymers such as polyglycolic acid (PGA) and polyethylene glycol (PEG) [60,61]. Cartilage tissue grown in vitro from cell-seeded scaffolds has certain limitations. When compared to natural cartilage, tissue is generally of inferior quality than real cartilage, with lower collagen content and a fraction of mechanical aspects [62]. This could be due to a lack of weight-bearing during tissue development, which has been found to impact gene expression, matrix formation, and nutrient transport. 4.3.2.1. Natural scaffolds. Chitosan, collagen, alginate, silk fibroin, hyaluronan, and gelatin are natural polymers that have been employed in TE for cartilage regeneration [63]. They promote the development of ECM due to their enhanced biocompatibility, great biodegradability, achievable mechanics, and favourable cellular adhesion. Fig. 4. Schematic diagram of autologous chondrocyte implantation (ACI). Fig. 5. Schematic diagram of treatment for repairing cartilage defect based on stem cell implantation. Route 1: isolated cells are directly injected into the damaged zone of cartilage. Route 2: isolated cells are amplified in vitro and then injected into the damaged zone of the cartilage tissue. Route 3: isolated cells are amplified and activated toward chondrogenesis using growth factor and then injected into the damaged zone of the cartilage tissue. A.A. Elhadad et al. Applied Materials Today 29 (2022) 101603 6 a) Chitosan (CS) Due to its content that is similar to the ECM components of the original tissue, chitosan (CS), a naturally occurring polymer, has emerged as an attractive biomaterial for the application of TE to bone, skin, and cartilage without complications [61,64,65]. CS offers a favourable environment for chondrocyte proliferation, ECM formation, and chondrogenesis in cartilage regeneration, since it is a natural substance with structural similarities to sulfated glycosaminoglycans [66, 67]. Because it increases calcium/phosphate ion deposition and improves biomineralization capabilities, it has been employed to boost the biomineralization of composite materials [68,69]. CS has the potential to increase osteogenesis, strength, and hardness of the scaffold produced, and has been employed with bioceramics such as calcium phosphate, calcium polyphosphate, or hydroxyapatite for cartilage TE [70]. CS-hyaluronic acid [71], CS-graphene oxide [72], CS-oligopyrrole [73], carbon nanotube [74], and CS-chondroitin sulfate [75] are some more combinations of CS employed in cartilage TE. Its hydrophilic nature facilitates cell adhesion, proliferation, and differentiation [76]. Scaffolds containing CS could be made in 2D (films and fibres) or 3D (gels and sponges) and provide the properties needed for wound healing [77]. CS is a chitin-derived linear cationic polysaccharide with remarkable biodegradability, biocompatibility, non-antigenicity, adsorption, antibacterial and antifungal qualities, and wound-healing capabilities (Fig. 6) [64,78,79]. CS is produced by partially deacetylating chitin with a deacetylation degree (DD) of 70% or higher, either chemically or biologically, or by mixing the two. Most commercially available CS has a DD of 70–90%; even DDs greater than 95% can be achieved by further deacetylation procedures. It is produced by partially deacetylating chitin, chemically or biologically, with a deacetylation degree (DD) of 70% or higher, or by combining the two [80]. Therefore, chitosan chains are formed as a copolymer mixture of 2-acetamido-2-deoxy-D-glucose and 2-amino-2-- deoxy-D-glucose units linked by β–(1→4) glycosidic bonds [80]. In animal metabolism, CS is considered biodegradable because it may be broken down by enzymes that hydrolyze glycosidic bonds, such as -GlcN-GlcN-, -GlcN-GlcNAc-, -GlcNAc-GlcN-, and -GlcNAc-GlcNAc- (Fig. 6). Furthermore, it has been established that the degree of deacetylation and molecular weight of CS determine its biodegradability in living organisms. The lower the deacetylation, the larger the molecular weight, which increases the chemical stability and mechanical strength while lowering the solubility in standard solvents [81]. Because of the availability of free amine groups in its carbon skeleton, CS can be chemically modified to include important features for biomaterial development, such as adding antibacterial medicines to improve its antimicrobial effect [76]. Furthermore, the enzymatic crosslinking of chitosan hydrogels can stimulate the chondrocyte and MSCs proliferation, preserve the chondrogenic phenotype and shape, and increase cartilaginous ECM deposition in vitro [67]. Due to the insolubility nature of CS in aqueous solutions, it has several drawbacks that limit its use in Fig. 6. Chemical structure of chitin, chitosan and potential metabolites. The partial deacetylation of the 2-acetamide-2-deoxy- D-glucose (GlcNAc), forming the chitin chains by β–(1→4) bonds, leads to chitosan. Additionally, the β–(1→4) links could be hydrolyzed by the points bonding a) two 2-amino-2-deoxy-D-glucose units (GlcN), b) one GlcN and one GlcNAc unit and/or c) two GlcNAc units. A.A. Elhadad et al. Applied Materials Today 29 (2022) 101603 7 living systems [82]. Extra-molecular hydrogen bonding produces a hard crystal structure. However, due to its molecular structure and pK a (6.2 to 7.0), it dissolves in acid solutions (pH <6) [83]. The acid most commonly used for this purpose is acetic acid (1–3%, v/v), and citric acid (3%–4%, v/v) although several other acids found in the human body, such as HCl, lactic acid, and pyruvic acid, may also solubilize CS in water [84]. Recently, water-soluble chitosan (WSC), such as chitosan salts, zwitterionic CS, and CS oligomers, has attracted attention [85]. The development of the WSC simplifies the procedure while avoiding the use of acidic solvents [85]. In-load bearing applications, (e.g., AC repair), the CS scaffolds has poor mechanical integrity in their wet form and, therefore, require extra modifications [86]. The application of a crosslinking agent or different reinforcing additives could achieve this change [87]. Another issue is enhancing the designed scaffold’s biological activity, particularly its pro-chondrogenic potential, by adding biological markers relevant to the cartilage-ECM [87]. a) Collagen Collagen is a major structural protein found in most hard and soft tissues in animals and plays a crucial role in tissue formation and maintenance [88]. It could be found in abundance in corneas, cartilage, bones, blood vessels, stomach, intervertebral discs, and tooth dentin. Collagen is the main component of numerous organs, and aids to serve a variety of functions in the body, including transferring force between bone and muscle, providing flexibility to harder tissues such as teeth, cartilage, blood vessels, skin, and muscles [89]. Collagen is essential for preserving the physiological and structural integrity of the ECM, and also for maintaining tissues with the physical characteristics they need to fulfill specific functions [88]. Collagens are key components of the ECM because of their high biocompatibility, low immunogenicity, interaction with growth factors and cell adhesion molecules, biodegradability, and ability to support cell ingrowth and remodeling [90]. Because of all these merits, collagen scaffold has been employed as a biomaterial in a variety of TE applications, including tissue repair and regeneration [91]. The collagen molecule is composed of three polypeptide chains [92]. A triple helix of collagen has a thousand or more amino-acid residues in each polypeptide chain. A glycine (Gly) is found in every third residue, resulting in an Aaa-Baa-Gly-Caa-Daa-Gly-repeat unit, where Aaa, Baa, Caa, and Daa can be any amino acid and will depend on the nature of the collagen (Fig. 7). The presence of Gly-on a triple helix of tropocollagen guarantees that the three strands are tightly packed together [92]. These parallel chains coil and wrap around each other to form a left-handed polyproline II-type (PPII) helix to create a right-handed triple helix maintained by interstrand hydrogen bonds and intrastrand n→ π * interactions [92]. In an n→ π * interaction, a nucleophile donates lone-pair (n) electron density to the empty π * orbital of a nearby carbonyl group. When these orbitals are mixed, energy is released, resulting in an attractive interaction. Only a few forms of collagen are employed in the manufacture of collagen-based biomaterials. The triple helices of tropocollagen aggregate into fibrils in fibrillar collagen, which agglomerate to form fibril-forming collagens. The types of collagen-based biomaterials most often employed for wound healing and tissue engineering are fibril-forming collagens, which include considerable sections of homologous sequences [92]. Pure collagen scaffolds degrade quickly, have poor mechanical characteristics, and lack structural stability, hence collagen-based composite biomaterials are preferred for cartilage tissue synthesis (CTE) [88,93]. Physical treatments or chemical agents could be used to increase the mechanical properties of collagen scaffolds while decreasing their solubility and absorption, while avoiding toxicity. This includes UV irradiation, gamma-ray, and dehydrothermal treatment (DHT) [94,95]. Without generating any toxicity, this kind of treatment could reduce the solubility of collagen while improving mechanical properties as well as triggering MSCs elongation and differentiation [96]. Commonly used in chemical modification is the cross-linking agent glutaraldehyde (GA). It may, however, produce collagen with a high degree of cross-linking but with toxic effect due to probable residual in the scaffold [97]. Because of their great biodegradability and biocompatibility, low toxicity, high effectiveness, and significant effectiveness, collagens are an excellent alternative for drug delivery [98]. However, collagen has weaker interactions with enzymes and other bioactive substances than hydrophobic polymers. In AC tissue engineering, Collagen-based scaffolds greatly increased the expression of certain osteogenic gene markers for human adipose stem cell (hASC) [99]. The chondrocytes embedded in hydrogels made of collagen also preserved their natural morphology and produced ECM unique to cartilage [100]. a) Silk Fibroin Silk fibroin (SF) is one of the oldest polymers found in nature, derived from Bombyx mori silk cocoon (mulberry sources), has a 380- million-year historical evolution and is considered an interesting polymer for a variety of biomedical purposes [101]. SF is a biocompatible substance located in the center of a protein fiber that is covered with sericin and has been employed in various TE applications [102]. The mechanical properties of SF derived from non-mulberry sources, such as the tasar silkworm (Antheraea mylitta), are superior to SF derived from mulberry sources [103]. Biomaterials manufactured using SF offer a number of benefits over natural polymers derived from allogeneic or xenogeneic tissue sources, including ease of extraction, low cost and risk of infection [101]. SF scaffolds are strong enough to be easily handled Fig. 7. Chemical structure of collagen. A.A. Elhadad et al. Applied Materials Today 29 (2022) 101603 8 during implantation and have mechanical qualities similar to those of the native tissue being repaired, providing optimal healing conditions [104]. SF-based biomaterials have a decreased risk of infection and lower costs due to less complicated manufacturing procedures [101]. The SF skeleton is made up of two proteins: sericin, which is the outer-layer covering protein, and fibroin, which is the inner-layer protein [105]. Heavy chain (H-chain; MW 360 kDa), light chain (L-chain; MW 27 kDa), and P25 (MW 25 kDa) proteins make up SF. The main component of SF is the H-chain, which is in charge of the mechanical characteristics of the SF fibres [106]. Disulfide binding binds the L-chain to the H-chain, forming an H-chain–L-chain heterodimer that promotes secretion and intercellular transfer (Fig. 8). Six H-chain–L-chain heterodimers are assembled by one P25 [106]. H-chain, L-chain, and P25, which are present in a molar ratio of 6:6:1, are the three polypeptides that constitute Bombyx mori cocoon [107]. The amino acid sequence Gly-Ser-Gly-Ala-Gly-Ala-constitutes the primary high molar-mass structure of SF. This primary structure might crystallize as stable anti-parallel β-sheets, increasing SF stiffness and tensile strength [108]. On the other hand, the protein structure constitutes 70% of the -sheet crystalline structure of SF, which is bordered by hydrophilic and non-repetitive amorphous regions [109]. At room temperature, SF may be generated from a water-based solution with a neutral pH, and it can be formed into α -helix, however, it has poor mechanical qualities. Various treatment strategies have been tested in recent years to enhance the strength of SF, including enzymatic, chemical, and physical crosslinking of SF hydrogels [110]. Enzymatic crosslinkers such as genipin, chemical crosslinkers such as glutaraldehyde and carbodiimide hydrochloride, and physical crosslinking methods like heating, vortex shearing, ultrasonication, or electric current speed up the hydrogelation of SF by supplying enough energy to overcome the energy barrier of β -sheet formation [111]. SF fibres have demonstrated outstanding mechanical properties [112]. Large break strain (4–26%), ultimate strength (300–740 MPa), and toughness (70–78 MJ⋅m –3 ) are among them [112]. Furthermore, SF fibres are believed to be tougher than several synthetic fibres, including Kevlar (50 MJ⋅m –3 ), carbon fiber (25 MJ⋅m –3 ), and various collagens, such as tendon collagen (7.5 MJ⋅m –3 ) [113]. Additionally, SF fibres have the highest strength among common natural materials including wool, resilin, elastin, cotton, and byssus, and even some synthetic fibres like synthetic rubber and viscose rayon [107]. The most common dominant crystalline structures of SF are silk I and silk II. Silk I is a metastable crystalline structure with attached water molecules, while silk II is the most stable thanks to massive hydrogen bonding between neighbouring peptide blocks, leading to improved mechanical characteristics such as rigidity and tensile strength [114]. On the other hand, successful scaffold materials should be biodegradable and bioresorbable, since scaffolds should be gradually replaced by patients’ own cells and ECM during healing [107]. SF is an enzymatically biodegradable polymer that has not yet been proven to elicit any immune response [107,115]. The quantity of β-sheet secondary structures that exist affects how fast SF degrades. Compared to RSF silk scaffolds, native silk fibres degrade much more slowly. This is due to the larger amount of β-sheet secondary structure in native silk fibres than in RSF structures [115]. γ-radiation has also been proven to promote the degradation of the SF fiber, due to the conversion of silk II to silk I [116]. In comparison to synthetic polymers, the SF polymer is better suited for load-bearing TE applications, particularly in musculoskeletal TE, due to their ability to retain strength in vivo for long periods of time. This is because synthetic materials degrade faster and have less desirable mechanical properties, which is important in TE owing to the need for gradual degradation and load-bearing capabilities [101]. a) Alginate (ALG) ALG is a naturally occurring polymer obtained from the walls of brown algal cells that is widely used in TE due to its low cost, low immunogenicity, good biocompatibility, availability, and ability to physically crosslink [117]. It is a natural polysaccharide consisting of residues of D-mannuronic acid (M-block) and L-guluronic acid (G-block) that are 1→4 linked [75]. Alginate is a biocompatible material with high water content, superior porosity, and varied viscosity that is widely known for producing hydrogels that may be used to load cells and Fig. 8. Schematic diagram of silk fibroin molecular structure showing proteinic chains interaction to form hydrogen bonds to stabilize SF secondary structure (random coil, α -helix or β-sheets). A.A. Elhadad et al. Applied Materials Today 29 (2022) 101603 9 pharmaceuticals [118]. Alginate has indeed been widely used to encapsulate many types of cells in hydrogels due to these attributes, whilst being approved by the FDA for human use as a food ingredient and wound dressing material [117]. It also preserves long-term stability in vivo because mammals lack alginate lyases or other related enzymes, making it impossible for alginate to break down [119]. The biocompatibility, biodegradability, and inexpensive cost of ALG make them one of the top alternatives for cartilage regeneration [120]. Because pores operate as a good environment for cell migration, ALG has been employed by researchers to build a porous 3D spongy scaffold that provides support for chondrocytes [121]. The linear chains of ALG contain repeated monomeric units of α -L-guluronic acid blocks and 1→4-linked β-D-mannuronic acid epimers with free functional hydroxyl (OH) and carboxyl (COOH) groups (Fig. 9). Alginate source and extraction procedures have a significant impact on the M/G ratio of these monomers, their organization as homopolymeric (GG or MM) or heteropolymeric (GM) blocks, and the molecular weight of the polymer chains [122]. The gel strength is also affected by the proportion of heteropolymeric (GM) or homopolymeric (GG or MM) regions in alginate. Alginates with a high percentage of heteropolymeric areas create weaker gels than those with a high percentage of homopolymeric regions, which strengthen intermolecular linkages [123]. The viscosity and gelation kinetics of alginate solutions are also affected by the molecular weight of the alginate, as well as the type, quantity, and length of the G/M block [124]. Alginate has a low mechanical strength however its mechanical characteristics could be improved crosslinking strategy. During crosslinking, the ALG can interact with divalent cations (such as Ca 2+ ), resulting in hydrogels (polymeric 3D meshes capable of storing enormous quantities of water) under moderate and biocompatible circumstances. Ionic gelation of alginate occurs because of interactions between divalent cations and alginate free carboxylate (COO ‾ ) groups, mostly but not only those present in G blocks [123]. Other divalent cations that can cause alginate crosslinking include Ba 2+ , Mn 2+ , Sr 2+ , Zn 2+ , and Mg 2+ , however, their affinities are distinct [125]. Ca 2+ has a lower affinity for alginate than Ba 2+ and Sr 2+ . various forms of alginate hydrogel scaffolds have been produced and employed as a potential candidate for AV repair. The chondrocytes embedded within alginate hydrogels maintained their natural morphology and preserved their ability to secrete cartilage markers like GAGs and COL II [126]. a) Cellulose Cellulose (CS) is an unbranched and most abundant polymeric material on earth, which is the main structural component that accounts for around 40% of the carbon in plants [127]. In nature, CS may be found as microfibrils in wood and plant cell walls, algal tissues, and tunicate epidermal cell membranes, as well as in the form of nanofiber networks produced by bacteria. The vast number of hydroxyl groups in a cellulose chain allows the production of hydrogels with a variety of biodegradable structures with suitable physical and mechanical qualities that may be used as a platform for enhanced TE [127]. CS (C 6 H 10 O 5 ) n is a high molecular weight, linear polysaccharide made up of repeated connected D-glucose units connected by β(1→4) glycosidic bonds [128]. Each glucose residue has three hydroxyl groups, with degrees of polymerization ranging from 1000 to 15,000 units, depending on the origin and treatment of cellulose Intra-chain hydrogen bonding between hydroxyl groups and oxygen of the next ring of cellulose molecules helps to balance linkages, leading to unbranched cellulose chains [129]. Cellulose is a semi-crystalline polymer with varying degrees of crystallinity depending on its source, processing circumstances, and pretreatment. The degree of crystallinity of wood-based and plant-based cellulose typically varies from 40 to 60%, but CS derived from other sources, such as bacteria and tunicin, has a greater degree of crystallinity, ranging from 80 to 100% [130]. Four different allomorphs of crystalline CS have been identified: cellulose I, II, III, and IV (Fig. 10) [130]. Cellulose I, which is thermodynamically metastable, can be transformed into cellulose II or III (Fig. 10b). Until recently, cellulose II has been thought to have the most stable structure, notably a monoclinic structure, which makes it technically useful. John Mercer discovered the conversion of cellulose I to cellulose II in 1844 when treating cotton with an alkaline solution [131]. Cellulose II is generally generated via mercerization (aqueous sodium hydroxide treatment) or regeneration (dissolution and recrystallization) of native cellulose [132]. During this process, cellulose I’s parallel chain arrangement transforms into cellulose II’s more stable antiparallel chain arrangement. Cellulose III is created by treating cellulose I or II with liquid ammonia and is referred to as cellulose IIII or IIIII, respectively [131]. Following thermal treatments, both cellulose IIII and IIIII may be converted to cellulose IV [132]. CS scaffolds have attracted a lot of attention in the biomedical industry because of their low cost, low density, biocompatibility, biodegradability, low cytotoxicity, and mechanical properties. Cellular adhesion, growth, and proliferation, as well as wound healing, have all been reported to be stimulated by CS [127]. Nevertheless, its reduced solubility in water and common organic solvents, and low resistance to microbial attacks have limited its applications [133]. This is due to intramolecular and intermolecular hydrogen bonding stabilizing cellulose molecules, as well as electrostatic and hydrophobic interactions within the integrated fibrils, generating stiff bundles with many hydroxyl groups dispersed throughout the backbone [134]. Cellulose can be chemically modified to address less favourable features or to acquire new desirable properties by substituting its natural hydroxyl groups with functional groups such as certain acids, chlorides, and oxides. a) Gelatin Gelatin is a naturally occurring biodegradable and biocompatible polymer composed primarily of proteins, mineral salts, and water [135]. It may be found in several sources, including cow bones, fish, pig skins, and insects [135]. It is a molecular derivative made from the irreversible hydrolysis of collagen’s triple helix structure, which can be achieved by heat or enzymatic denaturation [136]. Gelatin is less organised than collagen (Fig. 11), but has a molecular structure that is quite comparable to collagen, suggesting that it might be used as a biomaterial for cellular growth in vitro [136]. Because gelatin is less expensive and more soluble than other ECM proteins, it has attracted a great deal of attention. Fig. 9. Chemical structure of alginate. α -L-guluronic acid and β-D-mannur- onic units. A.A. Elhadad et al. Applied Materials Today 29 (2022) 101603 16 variety of crosslinking mechanisms, such as physical or chemical crosslinking [225]. To date, alginate is by far the most frequently applied material in 3D-bioprinting techniques; however, synthetic hydrogels have received much interest due to their better control over parameters like degradation and mechanical stability [226]. Comparing natural and artificial hydrogels as ingredients for bio-inks for 3D bioprinting is shown in Table 2. Hybrid bio-inks are hydrogels that are produced using both natural and artificial materials. Researchers can use the advantages of this combination to produce a bio-ink that is exactly tailored to the desired application. 5.3. Cartilage bioprinting approaches Technically, cartilage structures are printed employing either direct bio-ink (called self-supporting hydrogel bioprinting) and printing or hybrid bioprinting, which alternates the printing of bio-inks with thermoplastic-polymer networks (called hybrid bioprinting) [229]. 5.3.1. Self-supporting hydrogel bioprinting CTE cartilage structures are created using self-supporting hydrogel bioprinting processes by printing natural and synthetic hydrogels containing stem cells or chondrocytes [209]. Chondrocytes and stem cells placed in alginate hydrogels have been shown to be viable and metabolically active [230]. Alginate is a frequent component in bio-inks for printing cartilage structures because of its rapid crosslinking. Also, alginate and nanocellulose were combined to create a highly printable bio-ink [126]. The printed structures allowed human nasoseptal chondrocytes to be cultured and could be printed into more complex forms [126]. Sulfated alginate can bind growth factors, including fibroblast growth factor (FGF) and transforming growth factor (TGF), without compromising its printability [231]. The use of alginate in CTE is currently limited because of the lack of adequate cell adhesion sites. The bioactivity and printability of the bio-ink were greatly enhanced by integrating BioCartilage (extracellular matrix particles of cartilage) and gellan in alginate, and the resulting patient-specific cartilage grafts demonstrated good mechanical and biological characteristics [232]. Due to its structural and biological features, hyaluronic acid (HA), a key component of cartilage ECM, can regulate cellular signaling, wound healing, and ECM organization [144]. Owing to its viscoelastic and bioactive qualities, HA has recently been investigated as a "building block" in several bio-ink formulations for cartilage bioprinting in CTE [233]. However, due to its water solubility, one of the main disadvantages of unmodified HA for cartilage bioprinting is its low stability. To address the issue of the low stability of printed HA, a photo-crosslinkable dextran derivate or acrylate Pluronic was added to increase the material’s mechanical characteristics and printability. Furthermore, inserted chondrocytes showed that this bio-inks formulation was compatible with them [234]. Although gelatin gel has been shown to enhance chondrocyte survival and differentiation, it is difficult to print due to its low viscosity and de-crosslinking at 37 ◦C [235]. As a result, gelatin is frequently photo-crosslinkable by a simple reaction with an acrylate or methacrylate agent [236]. 5.3.2. Hybrid bioprinting A hybrid construct that combines the benefits of hydrogels and thermoplastics has been proposed, with potential applications in CTE [62]. Scaffolds comprised of thermoplastic polymers have better structural qualities, whereas hydrogels offer chondrocytes with a physiologically advantageous, highly hydrated microstructure, similar to native cartilage ECM. Hybrid cartilage structures are created by printing thermoplastic polymer and cell-laden hydrogels alternately [62]. 5.4. Zonal cartilage bioprinting The zonal cartilage construct (Fig. 17) that represents the inherent structural depth-dependent properties of AC may have an advantage over homogeneous construction [237]. The following techniques can be used to create zonal cartilage: (i) chondrocyte subpopulations from diverse zones of cartilage; (ii) a single cell source paired with the appropriate biochemical and/or biomechanical stimuli; and (iii) a variety of biomaterials and smart scaffold designs [238]. In strategy (i), zonal chondrocytes are extracted from distinct layers of cartilage; however, this technique has limitations such as donor site morbidity and limited availability [239]. In comparison to this one, the strategy (ii), which uses a single cell source paired with the appropriate biochemical and/or biomechanical cues, may be a simpler and more practical strategy [240]. BMSCs have been co-cultured with several substances to stimulate differentiation into zonal chondrogenic cells. This technology has great potential because it would be easier to use and might perhaps overcome the issues that come with directly isolating zonal chondrocytes from cartilage [240]. Researchers provided a good illustration of Strategy (iii), by cultivating BMSCs on electrospun and oriented PCL scaffolds, they were able to successfully simulate the cells and ECM architecture present in the superficial zone [241]. These strategies can be used to create bio-inks for the creation of complicated zonal structures. 5.5. 3D-bioprinting vs traditional manufacturing Conventional fabrication methods involve milling and cutting of pieces or even hand-made moulds to build the structures [242]. In contrast, additive manufacturing (AM) refers to a set of techniques, including 3D-bioprinting, that allow for the quicker and more precise production of diverse structures, such as organs and blood vessels [242]. In addition, 3D-bioprinting technique allows the development of tissue-like 3D structures with mechanical and biological features suitable for the deposition of living cells and the restoration of tissue and organ function as opposed to conventional manufacturing methods for producing cell-free scaffolds [243]. Other benefits include the ability to maintain shape and chemistry in a controlled manner due to linked porosity as well as no risk of tissue rejection or disease transmission [244]. Over traditional manufacturing techniques, 3D- bioprinting has various advantages, including precise cell dispersion, high-resolution cell deposition, scalability, and pricing [244]. Table 2 Comparison between natural and synthetic hydrogels as bio-inks ingredients for 3d-bioprinting. Bioink Hydrogel Synthetic Natural Description Made up of a hydrophilic polymer network linked via a covalent bond or physical attraction [220] Derived from synthetic and natural polymers, e.g. polyethylene glycol (PEG)- based materials such as PEG diacrylate (PEGDA) and polyacrylamide (PAAm)-based gels Obtained from biological material, e.g., collagen, fibrin, hyaluronic acid Advantage Hydrophilic, allows for gas and nutrition delivery, compatible, nontoxic Easily modified, e.g., non-immune response Highly biocompatible Disadvantage Poor mechanical aspects and poor cell adhesion [220] lack of adequate cell adhesion sites [227] Limited modification [227] Viscosity Adjustable [228] Variable (PEG: low, Pluronic-acid: High) [227] Gelatin and Fibrinogen: low Hyaluronic Acid: high Silk: high [227] A.A. Elhadad et al. Applied Materials Today 29 (2022) 101603 17 6. Conclusion and future directions The limited spontaneous repair that follows cartilage injury after high-impact sports or diseases, is demanding a growing interest for more appropriate tissue engineering approaches in cartilage regeneration applications. Polymers and polymeric composites are among the most used materials for this purpose due to their high biocompatibility, tunable properties [245–247], and the possibility of tailoring their mechanical properties by altering the degree of cross-linking [119,248]. In this sense, their biocompatibility has been extensively studied in various in vitro and in vivo experiments for a variety of tissues [249–252], including bone and cartilage [253–255]. Recent research has focused on the development of "biphasic implants", to overcome the feared osteochondral defects, and it is based on a flexible and porous bio polymeric composite that has been deposited and infiltrated on the surface of porous titanium substrates. On the one hand, gelatinous biopolymers appear as interesting entities, but they must satisfy tribo-mechanical parameters (elasticity and wear resistance) and should have sufficient porosity to allow for vascularization to be a viable candidate for replacing articular cartilage. On the other hand, the subchondral bone tissue could be replaced with a metal piece with a porosity that provides the bio-mechanical (stiffness and yield strength) and bio-functional (bone ingrowth) balance of the cortical bone that it intends to replace, avoiding bone resorption issues. Interestingly, the use of biphasic structures to support cartilage tissue cells as improved scaffolds conveniently infiltrated through a porous titanium substrate offers help for cells to proliferate, differentiate, and form cartilage tissue in vitro. Despite the promising advances previously described, unfortunately, there is still a significant challenge based on the need for proposing personalized "tailored-made" solutions (type of joint, size and severity of the injury, as well as the age and gender of the patient, etc.), which would require multiple efforts, more knowledge and synergy including multidisciplinary research teams, STEM (engineers, medical, chemical, pharmacists, microbiologists, etc.). New and adapted manufacturing routes (e.g., laser sintering, inject printing, spark plasma sintering, electric current activated/assisted sintering, etc.) should be explored along with the development of low Young’s modulus titanium alloys and the synthesis of novel biopolymers with anti-inflammatory and antibacterial properties to explore more creative and new possibilities. This review broadens our understanding and promotes the potential use of biomaterials to stimulate specific cells to help heal the body to repair its own tissues in a natural way, following physiological processes as convenient future investigations. Funding This work was supported by the Ministry of Science and Innovation of Spain under the grant PID2019–109371GB-I00, by the Junta de Andalucía (Spain) through the Project Ref. P20_00671 and I +D +i FEDER Andalucía (Spain) through the Project Ref. US-1,380,878. A.E. thanks Ministerio de Universidades for an “Ayuda Recualificaci´ on María Zambrano 2022–2024′′. CRediT authorship contribution statement Amir A. Elhadad: Writing – original draft, Writing – review & editing. Ana Alcudia: Conceptualization, Methodology, Project administration, Funding acquisition. Bel´ en Begines: Writing – original draft, Writing – review & editing. Eva M. P´ erez-Soriano: Writing – original draft, Writing – review & editing. Yadir Torres: Conceptualization, Methodology, Project administration, Funding acquisition. Declaration of Competing Interest The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper. Data Availability No data was used for the research described in the article. Acknowledgments The authors would like to dedicate this paper to the memory of Dr Juan J. Pav´ on Palacio (University of Antioquia, Colombia). The library of Universidad de Sevilla for providing access to scientific and technical Fig. 17. Schematic of hydrogel bioprinting for fabrication of zonal cartilage constructs. A.A. Elhadad et al. Applied Materials Today 29 (2022) 101603 18 information resources. Some figures have been designed using resources from Freepik.com (brfgx and macrovector). References [1] S.J. Gilbert, C.S. Bonnet, E.J. 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