Articular cartilage is a highly organized tissue that provides remarkable load-bearing and low friction properties, allowing for smooth movement of diarthrodial joints [1]. Joint cartilage contains sparsely distributed chondrocytes embedded within the extracellular matrix (ECM). The ECM is mainly comprised of water, type II collagen, and glycosaminoglycans that provide the tissue with sufficient mechanical properties for several biological functions, such as load-bearing and low friction ^[2][3][2,3]. Due to the avascular, aneural, and non-lymphatic characteristics of cartilage, joint cartilage has self-regeneration and repair limitations [2]. When the cartilage gets damaged, if the diameter of the injury is greater than 4 mm, spontaneous self-repair capacity becomes limited [4]. Moreover, focal cartilage lesions predispose to developing early-onset osteoarthritis, which may lead to long rehabilitation periods and loss of function.

Treatment for cartilage injuries focuses on relieving pain [5]. It may include lifestyle changes, oral anti-inflammatories [6], physical therapy [7], intra-articular injections of hyaluronic acid or steroids [8], bisphosphonates [9], and even surgical interventions such as joint replacement [7]. Cell therapy is currently being proposed as an alternative strategy. In focal cartilage lesions, autologous chondrocyte implantation and mesenchymal stem cells (MSC) seeded onto scaffolds have been used to restore these defects with good results [10]. Other cell lines that have been used to repair focal cartilage lesions by implantation into the lesion site are embryonic stem cells and induced pluripotent stem cells (iPSCs) [11]; however, even though cell therapy has limited therapeutic activity, one of its main advantages is the short-term reduction in clinical symptoms. Thus, medicine has turned to a tissue engineering approach to prolong the therapeutic effect of cell therapy.

The successful use of tissue engineering techniques to form engineered cartilage is based on a combination of three critical elements: a cellular component, a bioactive component (such as growth factors), and a biocompatible and mechanically stable carrier vehicle/matrix scaffold [12]. The cellular component consists of healthy, viable cells that are accessible, manipulable, and nonimmunogenic. The bioactive component should promote the differentiation and maturation of the cellular component. The carrier has a dual function, acting as both a delivery vehicle and a scaffold [13]. In addition, the carrier should provide sufficient mechanical support to withstand in vivo forces [14] and must be degraded by the cells, giving way to its replacement or contributing to the formation of new tissue ^[15][16][17][15,16,17].

The basic types of biomaterials used In tissue engineering can be classified in relation to their origin as synthetic (usually chemical-nature materials) and natural (derived from biomolecules, tissues, or living organisms) [18] depending on its structural patterns as a polymer (composed of many repeated subunits) and a composite (a combination of two different biomaterials, a polymer, and a filler) [19]. Natural polymers such as collagen, silk fibroin, and fibrin are some of the most common, used as scaffolds for cartilage engineering. Fibrin is one of the most promising natural biomaterials for articular cartilage repair [20]. Fibrin polymers and composites have been used to induce regeneration as a vehicle for bioactive molecules to promote injury healing and delivery carriers for multiple cell lines [21]. Moreover, they are not expensive and are easy to obtain from whole blood.

2. Fibrin: Structure and Molecular Interactions

Fibrin is a native biopolymer derived from fibrinogen ^[22][23][22,23]. Fibrinogen is a blood component that plays an important role in hemostatic function. It is also related to cellular processes such as proliferation, differentiation, adhesion, migration, healing, inflammation, and angiogenesis [24]. Fibrinogen is described as a long glycoprotein (340 Kda) made up of a dimer of three disulfide-linked polypeptide chains called Aα (66,500 Da), Bβ (52,000 Da), and γ (46,500 Da) ^[22][25][22,25]. Fibrinogen consists of two globular D regions and one central globular E region, each with a part of α-helical coiled-coils. Fibrinogen is transformed into fibrin monomers during blood clotting due to thrombin [23], which cleaves fibrinopeptide A (FpA) and fibrinopeptide B (FpB) from the N-terminal sites of the Aα and Bβ chains of fibrinogen, respectively Figure 1A. At this point the fibrin fibers made of two fibrin nano peptides half-staggered with a crystalline-like structure, can reach a size of 100 nm, After FpAs cleavage each α chains have a new sequence on the N-terminal (Gly-Pro-Arg) called knobs “A” [26], then these fibers can come together and form a crosslinked mesh. This initiates fibrin assembly by exposing a polymerization site called EA. Each EA site combines with a constitutive complementary-binding pocket in the D domain (Da) to form the initial EA:Da association by forming intermolecular ε-((γ-glutamyl) lysine bonds, causing double-stranded twisting fibrils by aligning in a staggered overlapping end-to-middle domain arrangement [27]. The g-g crosslinks form reciprocally between glutamine 398 or 399 and lysine 406, and other flexible bonds form such as a–a crosslinks Gln-221, -237, -328, -366 and Lys-418, -448, -508, -539, -556, -580, and -601, all conferring particular mechanical and elastic properties [28] (Figure 1B). The C-terminal region of each fibrinogen or fibrin c-chain contains one crosslinking site at factor XIII or XIIIa. These give fibrin structural integrity and stabilize the clot against proteolytic and mechanical insults because of isopeptide bond formation, passing from a soluble state to an insoluble one crosslinked by ε-(γ-glutamyl)-lysine stable covalent bonds [25] (Figure 1C). All the molecular and structural properties mentioned above allow crosslinking of the fibrin with different biomaterials, enhancing fibrin mechanical and elastic properties, and generating new biomaterials and scaffolds that resemble the physical properties of articular cartilage.

3. Mechanical and Physical Properties of the Fibrin Scaffolds

Due to the mechanical resistance, elastic, and mesh-like nature, fibrin has been used as a sealant for surgical procedures and recently as hydrogel scaffold for cartilage engineering. Generally, fibrin scaffolds can be manufactured in three forms: fibrin glues, fibrin hydrogels, and fibrin microbeads [13]. The fibrin glues are obtained from plasma cryoprecipitate (which contains fibrinogen, fibronectin, and XIIIa factor). The cryoprecipitate is mixed with thrombin and calcium to obtain a fibrin polymer, which can be used as a patch (2D) or as a 3D scaffold [29]. Fibrin hydrogels are made from purified fibrinogen, thrombin, and calcium salt. The main difference between fibrin glues and hydrogels is the presence of coagulable proteins in the fibrin glue [30]. On the other hand, fibrin microbeads are obtained from fragmented plasma and thrombin; however, polymerization takes place in an emulsifier at 75 °C where the fibrinogen gets denaturalized and the XVIIIa factor crosslinks the fibrin into a more stable and dense form [13]. In all the fibrin scaffolds, the mechanical strength will depend on the amount of fibrin and thrombin [31]; however, when it comes to fibrin composites, the fibrin becomes a filler and a functional part, and the mechanical strength will depend on the other component or phase of the biomaterial. To increase their mechanical strength, the fibrin scaffolds have been combined with different biomaterials, such as poly lactic-co-glycolic acid (PLGA) ^[32][33][34][32,33,34], hyaluronic acid (HA) [35], chitosan–alginate [36], polycaprolactone (PCL) [37], and although the results have been promising, increasing approximately 60 times the fibrin mechanical strength, in some cases, they match articular cartilage (0.24–0.85 Mpa [38]); however, it has not yet been possible to develop a scaffold that can match all the properties of cartilage. For example, an increase in the amount of fibrin leads to greater mechanical strength, but the pore size decreases (Table 1). Pore size must be appropriate for some cell types; for example, 150–250 µm are desirable for articular chondrocytes [39] and 200–300 µm for MSC [40] to promote cell proliferation and the preservation of chondrogenic differentiation into the scaffold, ensuring the diffusion of oxygen, nutrients, and products of metabolism. Fibrin combinations with some biomaterials achieve a suitable pore size, but due to the chemical nature of the biomaterial [41], there is a lack of growth factors that can improve or maintain chondrogenesis, so these must be provided to the scaffold formulation [42]. Other combinations have achieved an accurate pore size and mechanical strength, but with a decrease in scaffold elasticity, as shown in Table 1. Creating a three-dimensional scaffold does not guarantee the creation of high-quality new cartilage tissue on its own. Instead, some help is needed to improve or induce the generation of new tissue. There are also numerous molecular interactions and conditions that contribute to the mechanical resistance of the scaffold. Usually, the scaffold is seeded with a cellular component, and this can improve the mechanical properties through ECM production [43]. Depending on the approach, the cell component in the scaffold should replace part of the scaffold for new tissue. At this point, the ECM can add mechanical support, elasticity, and stiffness [44]. That is why all the properties should be tested with and without cells to know the limitations of scaffolds. With the creation of different scaffolds for cartilage repair, it is critical to compare their properties to choose a therapeutic approach. However, it has been difficult to compare the properties of the scaffolds due to inconsistency in the number and type of tests performed to prove their functionality. Ideally, a minimum number of tests should be set to evaluate their functionality and compare with others.