Depending on the type of the biomaterial used, scaffolds are divided in synthetic, natural, or composite. There are several requirements, which should be met when designing scaffolds for tendon TE. The material should be biocompatible, promoting new tissue formation without immunological response. Sufficient void space is an important requirement in order to allow adequate nutrient delivery to the seeded or infiltrated cell population (Hollister 2005; Petrigliano et al. 2006). In fact this void space quantified as scaffold porosity is determinant for cellular proliferation and ECM formation, which enables the scaffold integration with host tissue (Hollister 2005; Laurencin and Freeman 2005; Vunjak-Novakovic et al. 2004). The goal is to achieve the same mechanical properties (e.g., elastic modulus, toughness, and ultimate strength) as the host tissue to restore normal physiologic function and prevent stress shielding of the newly formed tissue (Laurencin and Freeman 2005; Vunjak-Novakovic et al. 2004; Woo et al. 1999, 2006).

Moreover, biodegradability is an important feature. When implanting a scaffold, it is expected that this will provide the conditions for new tissue formation at the same rate of its own degradation that occurs (Doroski et al. 2007; Laurencin and Freeman 2005; Vunjak-Novakovic et al. 2004). Synthetic scaffolds offer the advantage of being able to be tuned to have a wider variety of mechanical properties and degradation characteristics through different fabrication and processing techniques. Despite these advantages, their degradation and wear products may cause an unwanted host immune response or toxicity. Natural scaffolds do not carry these risks and are readily degraded. The most popular current technique for production of nanofibrous scaffolds is electrospinning because it is a continuous, scalable process, which allows production of fibers ranging from nanometers to microns in either a random or aligned fashion (Fig. 13.3).

Synthetic polymers allow precise control over molecular weight, degradation time, hydrophobicity, and other attributes, yet they may not interact with cells in a desired manner (Langer and Vacanti 1993).

First-generation synthetic scaffolds were inert materials, which failed to simulate the biomechanical properties of native tendon. They also did not behave as expected concerning creep, permanent deformation, stress shielding, abrasive wear and degradation, axial splitting, and extensibility (Petrigliano et al. 2006; Vunjak-Novakovic et al. 2004). Unlike the previous generation, new biomaterials have high initial strength and are designed to gradually transfer the load to the newly formed tissue, while the degradation process is progressing and the new tissue matures.

Poly(urethane) is a very heterogeneous and versatile group of polymers in what concerns their physical and mechanical properties. Depending on the fabrication process, these scaffolds can have different porosities and tunable degradation characteristics (Webb et al. 2006). These scaffolds seeded with fibroblasts have shown increased cell proliferation, matrix accumulation, and elastic modulus after a cyclic strain regimen (Webb et al. 2006). However, poly(urethanes) may release toxic degradation products and are unable to maintain similar mechanical properties to the native tendon upon degradation (Guelcher 2008).

Poly(esters) when used for tendon regeneration as scaffolds are produced in a fibrous form that is posteriorly processed to form higher ordered structures such as sheets (Deng et al. 2009; Wang et al. 2008), woven or knitted meshes (Heckmann et al. 2008), and braids (Freeman et al. 2007, 2009). PLLA, PGA, and PLGA are the most commonly used poly(esters) for tendon tissue engineering. The large interfiber pores have been shown to allow significant collagen deposition both in vitro and in vivo in a rabbit model for ACL reconstruction (Freeman et al. 2007, 2009; Van Eijk et al. 2008).

These materials have a wide range of degradation profiles and mechanical properties. However, a material that retains these properties during the replacement process by new tissue remains to be found (Vieira et al. 2009).

Also natural scaffolds have evolved during time. Initially, porous scaffolds composed of randomly oriented collagen type I fibers were used. These scaffolds without cells had suboptimal mechanical properties (Torres et al. 2000). In order to overcome this limitation, other approaches were tried. Gels composed of aligned collagen fibers seeded with MSCs and submitted to mechanical stimulation had showed significantly better mechanical properties (Altman et al. 2002; Shearn et al. 2007). MSC-seeded porous collagen sponges which generate aligned collagen fibers under cyclic strain have demonstrated increased ECM gene expression and protein deposition (Juncosa-Melvin et al. 2007). The main drawbacks of collagen matrices are the need for processing the collagen derived from animal tissue in order to remove foreign antigens and potential donor pathogens to reduce immunogenicity. These processes reduce the mechanical strength and collagen also undergoes fast in vivo degradation (Liao et al. 2008; Vieira et al. 2009).

Fibrin matrix cross-linked with thrombin and containing bone marrow stromal cells showed initially promising results, but as with collagen matrices, they are still unable to achieve the ideal mechanical properties (Liao et al. 2008; Vieira et al. 2009).

From the natural scaffolds discussed here, silk represents the strongest, being also biocompatible, promoting cell adhesion, proliferation, and ECM deposition (Vunjak-Novakovic et al. 2004). Its degradation occurs at a relative slow rate in vivo, and when its fibers are arranged in a wire rope, it achieves similar mechanical properties to those of native ligament or tendon (Chen et al. 2003, 2006; Moreau et al. 2008). However, silk presents some concerns about the adequate removal of contaminating proteins and potential immunogenicity since it is not a human native protein (Vieira et al. 2009).

Composite scaffolds are the result of the association of natural and synthetic materials in an attempt to use the positive aspects of each material in synergy. There are several studies in which fibrous scaffolds were combined with hydrogels to allow homogeneous cell seeding within the fibrous scaffold, promote the alignment of cells and ECM in the gel, reinforce the gel scaffold, and allow delivery of nutrients and growth factors through the gel (Fan et al. 2008; Hayami et al. 2010).

Some groups have been developing composite materials in order to replicate tendon interfaces. A stratified scaffold using three distinct continuous phases was produced using knitted meshes and sintered microspheres. Each phase was designed to reproduce a specific region of the tissue. This scaffold supported cell proliferation, migration, and production of ECM, maintaining distinct cellular regions and phase-specific ECM in vitro (Spalazzi et al. 2008).

Other approaches have been exploring the incorporation of ceramics with different fabrication technics to achieve graded mineral deposition (Paxton et al. 2008).

Human hair keratin offers great possibilities for tissue regeneration and repair applications because of its abundance, bioactive properties, and potential as a reliable autologous material (Rouse and Van Dyke 2010). Among existing biomaterials, fibrous matrices produced by electrospinning technique are widely used because they can mimic the nanoscale structure and complexity of the ECM and are controllable in fiber diameter and alignment (Khorshidi et al. 2016). It has been previously reported success in electrospinning randomly oriented keratin fibers by blending human hair keratin with a small amount of poly(ethylene oxide) (PEO) and showed that these are compliant for fibroblast attachment and proliferation (Sow et al. 2013).

Despite these scaffolds have an enormous potential, they are still in a premature phase of development. Their use in large scale for tendon regeneration in humans is still far from being a reality.

Once tenocytes are the predominant cell type found in native tendon, it is logical to integrate them in biomaterials matrices in order to try replicating native tendon. In fact, these cells are responsible for producing and maintaining the ECM which provides mechanical strength (Benjamin and Ralphs 2000). Tenocytes not only are readily incorporated into most scaffolds where they adhere and proliferate (Guelcher et al. 2008; Wang et al. 2008) but also native cells can undergo migration and infiltration into implanted scaffolds (Freeman et al. 2009). A considerable flaw in this process is the need for autogenous cells in order to avoid host immune response. These cells need to be harvested from the patient’s tendons, which may lead to significant donor site morbidity.

Dermal fibroblasts represent a good alternative in order to overcome the difficulties and complications associated to tendon fibroblast harvesting. These cells can be obtained through a simple skin biopsy (Cornwell et al. 2004, 2007). Concerning basic behavioral characteristics, these cells adhere and proliferate in a similar way and appear to synthesize many of the same ECM components (Deng et al. 2009). Not being derived from the same tissue for which they are being applied remains the question if at long term the behavior will be similar to that of native tendon.