Grau de satisfacció del sistema de BiciPalma

Macroscopically, the scaffold shape dictates the boundaries of tissue formation. Microscopically, the scaffold’s structural framework controls cell ingrowth, neo-tissue organization, and nutrient diffusion. Thus, a common strategy to increase cell infiltration is to provide interconnected void spaces through which cells can migrate, such as pores or channels. Polymer processing techniques utilized to create highly porous morphologies include high-pressure gas foaming (Silva et al. 2006), porogen leaching (De Groot et al. 1996, van Tienen et al. 2003, van Tienen et al. 2002), freeze-drying (De Groot et al. 1996, van Tienen et al. 2003, van Tienen et al. 2002), and phase separation (Liu et al. 2009). Porous polyester and polyurethane foams (De Groot et al. 1996, Maher et al. 2010, Silva et al. 2006, van Tienen et al. 2003, van Tienen et al. 2002) and type I collagen-based sponges (Rodkey et al. 1999, Stone et al. 1992) have long been explored as meniscal substitutes, with a pore diameter range of ~150–300 μm and porosity greater than 80% being the most efficacious for cell infiltration (Klompmaker et al. 1993, van Tienen et al. 2002). While these materials facilitate cellular ingrowth, matrix deposition, and integration with native meniscal tissue within 6 months (van Tienen et al. 2003, van Tienen et al. 2002), the internal microstructure does not permit the generation of circumferentially oriented collagen fibers, which are essential for load transfer. Indeed, a study investigating polyurethane foams for meniscal replacement found scaffold fragmentation and cartilage damage after 2 years in a canine total meniscectomy model (Welsing et al. 2008).

To engineer functional dense connective tissues, it may thus be necessary to provide an instructive microstructure that fosters organized collagen deposition and maturation. It is well known that cells sense and respond to micro- and nanoscale topography, such as

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ridges, grooves, and channels (Lim et al. 2007). Anisotropic features, such as those provided by aligned electrospun fibers, induce cell polarization and de novo collagen alignment along the primary fiber direction (Figure 3-10), leading to greater increases in tensile modulus compared to non-aligned scaffolds (Baker et al. 2007).

Figure 3-10: Nanofibrous topography organizes cells and matrix. (A) Scanning electron micrographs of aligned (top) and non-aligned (bottom) PCL scaffolds. Scale = 50 µm. (B) Mesenchymal stem cells cultured on nanofibrous scaffolds for 7 days (green = actin, blue = nuclei). Cells are oriented with the primary fiber axis on the aligned scaffold. Scale = 50 µm. (C) En face section of scaffold stained after 10 weeks of in vitro culture with Picrosirius Red (collagen) and Alcian Blue (proteoglycans). Deposited collagen is aligned with the fiber direction. Scale = 100 µm. Adapted from (Baker et al. 2007, Mauck et al. 2009).

Although these organized scaffolds promote ordered matrix deposition, dense aligned nanofibers pose a challenge to cell infiltration. While scaffold alignment may increase migration speed and persistence along the fiber length (Fraley et al. 2015, Riching et al. 2014), contact guidance also hinders migration normal to the fiber axis (Dickinson et al. 1994). Introducing porogens such as salt particles (Nam et al. 2007), ice crystals (Simonet et al. 2007), or fibers with faster degradative rates (Baker et al. 2008, Lee et al. 2013a, Phipps et al. 2012a), have been shown to increase porosity and cell infiltration.

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Figure 3-11: Fast-degrading nanofibers in a composite scaffold increase scaffold porosity, cell infiltration, and integration with native tissue. (A) Schematic illustrating composite PCL/PEO scaffold with fast-degrading PEO fibers (green) and slow-degrading PCL fibers (red) before and after removal of PEO. (B) Scanning electron micrographs of scaffold with 0% or 60% PEO content removed. Scale = 50 µm. (C) DAPI staining showing nuclei of cell- seeded scaffolds after 12 weeks of in vitro culture indicates enhanced cell colonization of 60% PEO scaffold. Scale = 0.5 mm. (D) Percent of nuclei in the center of the scaffold after 12 weeks of in vitroculture. (A–D) Adapted from (Baker et al. 2012). (E) Integration strength of PCL/PEO scaffold with native meniscus after 4 and 8 weeks of in vitro culture shows enhanced scaffold-tissue integration with 60% PEO scaffold at 8 weeks. Adapted from (Ionescu et al. 2013).

For example, a composite scaffold of aligned, slow-degrading PCL fibers can be interspersed with a second population of water-soluble poly(ethylene oxide) (PEO) fibers (Baker et al. 2008) (Figure 3-11A). Rapid removal of the ‘sacrificial’ PEO fraction via

aqueous hydration results in a highly porous scaffold while preserving fiber directionality and mechanical anisotropy (Figure 3-11B). Although high porosity scaffolds (60% PEO) were initially weaker than low porosity scaffolds (0% PEO), collagen content within high porosity scaffolds (60% PEO) was significantly higher after 12 weeks of in vitro culture,

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resulting in superior tensile properties in the long-term (Baker et al. 2012). Furthermore, cellular infiltration of high porosity scaffolds (Figure 3-11C and Figure 3-11D) improves integration with tissue explants compared to scaffolds without PEO (Ionescu et al. 2013) (Figure 3-11E). Alternatively, instead of removing scaffold components, adding an interpenetrating fiber population (Cai et al. 2013, Lee et al. 2013b, Moutos et al. 2007) can provide vertical avenues for cellular ingress into the scaffold.

3.6.3 Biochemical Signals

Cellular perception of the biophysical microenvironment depends on the ability to adhere and exert forces on the surroundings. Thus, tunable systems that offer combined control over scaffold micromechanics, microstructure, and adhesivity have gained considerable interest. An attractive strategy is to use crosslinkable hydrogels, where synthetic polymers, such as poly(ethylene glycol) (PEG), and natural polysaccharides, such as hyaluronic acid (HA), are modified with functional groups (e.g., vinyl or thiol groups) to form hydrophilic crosslinked networks (Geckil et al. 2010). Mechanical properties rely on hydrogel crosslinking density, which is largely controlled by macromer modification and macromer density. Microstructure can be engineered via micromolding (Kim et al. 2012a), photolithography (Singh et al. 2014, Wade et al. 2015a), electrospinning (Kim et al. 2013, Wade et al. 2015a), and 3D printing (Billiet et al. 2012). Importantly, small oligopeptide sequences found within ECM proteins (e.g., RGD) can be conjugated to core molecules to provide adhesive ligands within the 3D matrix. For instance, Kim et al. fabricated nanofibrous scaffolds using methacrylated hyaluronic acid (MeHA) macromers with either 35% or 100% modification, resulting in soft and stiff fibers after final crosslinking (Kim et al. 2013). Additionally, RGD was coupled to the fibers at low, medium, or high concentrations. Although there was almost an order of magnitude difference between the bending moduli of soft and stiff fibers, cellular deformation of soft fibers were similar to

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that of stiff fibers at a low RGD density, and only became significantly different at a high RGD density, illustrating that proper adhesive cues are necessary for cell mechanosensing.

Furthermore, a depth-wise haptotactic RGD gradient can be formed within nanofibrous scaffolds by altering the flow rate of two polymer solutions with different conjugated-RGD concentrations to affect cell infiltration. Sundararaghavan and Burdick found that cell migration from aortic arch explants was more efficient in the presence of a low-to-high RGD gradient versus a high-to-low gradient (Sundararaghavan et al. 2011). Similarly, RGD-gradient channels formed by UV laser micropatterning guided neurite outgrowth towards a higher RGD concentration within a 3D hyaluronan gel (Musoke-Zawedde et al. 2006). Photo-immobilization of a laminin-1 gradient, an ECM glycoprotein found basement membranes, also increased neurite outgrowth in an agarose gel (Dodla et al. 2006). These results suggest that targeting adhesion via different ligand concentrations (Kim et al. 2013, Peyton et al. 2011, Singh et al. 2014, Sundararaghavan et al. 2011), ligand peptide sequences (Mhanna et al. 2014), and/or integrin antibodies (Wolf et al. 2013), can modulate migration without changing underlying material properties.

Cells isolated from the meniscus (Bhargava et al. 1999), tendon (Caliari et al. 2011), ligament (Hannafin et al. 1999), and shoulder capsule (Suzuki et al. 2001) can migrate towards a wide variety of soluble chemical gradients, including platelet-derived growth factor (PDGF) AB and BB (Bhargava et al. 1999, Bhargava et al. 2005, Hannafin et al. 1999, Suzuki et al. 2001), epidermal growth factor (EGF) (Kong et al. 2011), hepatocyte growth factor (Bhargava et al. 1999, Bhargava et al. 2005, Hannafin et al. 1999, Suzuki et al. 2001), bone morphogenetic protein-2 (BMP-2) (Bhargava et al. 1999, Hannafin et al. 1999), and interleukin-1 (Bhargava et al. 1999). Additionally, meniscal progenitors (Shen et al. 2014) and mesenchymal stem cells (Kucia et al. 2004) home towards stromal-cell

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derived factor-1α (SDF-1α), although this effect may be reduced in the absence of an injury (Abbott et al. 2004).

Therefore, biomaterials functionalized with chemoattractants also have potential to affect 3D cell infiltration. For example, Moore et al. immobilized neurotrophic gradients to advance neurite penetration into hydrogels (Moore et al. 2006). However, immobilization limits gradient sensing to cells that are already in contact with the scaffold, whereas delivery of soluble chemoattractants may also induce cell migration towards the wound site, thereby enhancing the intrinsic repair response of endogenous cells. In particular, the large surface area of nanofibers is well suited for an initial burst delivery, followed by sustained release as the fiber eventually degrades. Phipps et al. found that eluant from nanofibrous PCL/collagen/hydroxyapatite composites adsorbed with PDGF-BB enhanced directional migration of MSCs in vitro (Phipps et al. 2012b). Jin et al. used a PLA nanofibrous scaffold, embedded with PLGA microspheres releasing PDGF-BB, to enhance scaffold tissue ingrowth and neovascularization in a subcutaneous rodent model (Jin et al. 2008). Notably, sustained PDGF release from microspheres was more effective than a surface coating of PDGF, demonstrating that temporal control is necessary to establish a soluble gradient in vivo. Likewise, localized release of SDF-1α increased progenitor cell engraftment and colonization of scaffolds in subcutaneous experiments (Thevenot et al. 2010), reducing inflammatory cell infiltration (Sarkar et al. 2011, Thevenot et al. 2010) and improving the repair of various mesenchymal tissues, including bone (Holloway et al. 2015), myocardium (Purcell et al. 2012), skeletal muscle (Rybalko et al. 2015), and skin (Sarkar et al. 2011). Collectively, these findings indicate that scaffolds delivering soluble chemoattractants may also enhance dense connective tissue repair. However, it is important to note that the steric hindrance posed by dense, stiff, and/or non- degradable scaffolds may limit cellular ingress regardless of haptotactic or chemotactic cues.

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3.7 Summary and Future Directions

In conclusion, cell migration in confined 3D environments is governed by a complex set of cellular and extracellular variables, including cell mechanics and force transduction, matrix micromechanics and microstructure, and biochemical signals in the environment. While blocking metastatic invasion remains the primary impetus behind basic research in this field, enhancing interstitial migration can potentially benefit fibrous tissue repair and regenerative medicine. Decreasing matrix density and stiffness at the wound interface may facilitate local cell migration and intrinsic repair (Qu et al. 2015). Further modulating cell mechanosensing after cell movement into tissue and scaffolds may direct cells toward a proper lineage. For instance, while RGD ligands promote MSC adhesion and migration, prolonged adhesion may inhibit chondrogenic differentiation in the long-term. By combining a MMP-degradable sequence with the RGD peptide, Salinas and co-workers demonstrated that cells within a hydrogel could self-regulate the amount of adhesive ligands in the environment to support a chondrogenic phenotype (Salinas et al. 2008). Recent advances in biosynthetic polymers (Krishna et al. 2010), drug encapsulation and delivery (Garg et al. 2012), micro- and nanopatterning (Hahn et al. 2006, Kim et al. 2012a, Lim et al. 2007), and rapid prototyping (Billiet et al. 2012) gives us more control over the 3D microenvironment now than ever before. However, most of the aforementioned studies were performed in vitro with little or no mechanical loading. In an in vivo environment, tissues and scaffolds experience a variety of tensile, compressive, and/or shear forces. In dense connective tissues, strain transfer often depends on the relative sliding of collagen fascicles, fibers, and/or fibrils, resulting in spatiotemporal changes in the tissue microstructure (Han et al. 2013, Nerurkar et al. 2011, Szczesny et al. 2014). Dynamic tensile loading also activates mechanoactive signaling pathways that influence downstream cell force sensing and differentiation (Driscoll et al. 2015, Heo et al. 2011).

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Lastly, cell migration may be influenced by intercellular interactions, such as cell-cell adhesions and paracrine signaling (Friedl et al. 2010, Raeber et al. 2007), inflammation (Frenkel et al. 1996), and changes with cell age (Pajerowski et al. 2007). Although it is unknown how these physiological events collectively affect interstitial migration in vivo, it may be possible to study this phenomenon in vitro using bioreactors and live cell tracking systems, coupled with real-time second harmonic generation imaging to visualize ECM collagen. By designing smart, dynamic scaffolds that reflect the optimal microenvironmental niche for tissue growth and maintenance over time, ultimately we may recapitulate, and perhaps even augment, the natural biological cues that direct repair and regeneration.

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CHAPTER 4: BIOMATERIAL-MEDIATED DELIVERY OF DEGRADATIVE ENZYMES