**4. The foreign body response to biomaterials**

The implantation of a biomaterial can intensify the inflammatory response by inducing a foreign body response (FBR). The FBR at the tissue-material interface is composed of macrophages and foreign body giant cells (FBGCs) and forms the end-stage response of the inflammatory and wound healing responses following implantation of a medical device, prosthesis, or biomaterial [10,29]. Typically, within 2-4 weeks the foreign material is encapsulated within an almost avascular, fibrous connective tissue, depending on the porosity of the biomaterial [52]. The FBR is characterized by the presence of macrophages and FBGCs together with the components of granulation tissue. Macrophages and FBGCs are believed to exert critical effects on both tissue and implanted material, e.g. degradation, and chemokine and cytokine production [31].

#### **4.1. Macrophage fusion**

Macrophages develop integrins, which play a major role in the adhesion of macrophages to a biomaterial and in the IL-4-induced macrophage fusion to form FBGCs [10]. Macrophageintegrin binding to the protein layer on the biomaterial surface provides intracellular signals that can modulate macrophage behavior, such as cytoskeletal rearrangements and the formation of adhesion structures, called podosomes. There is extensive interplay between intracellular signaling molecules activated by integrin binding and cytoskeletal proteins. Disruption of the adhesion signals promotes anoikis, i.e. apoptosis induced by cell detachment from its supportive matrix. A hypothesis is that macrophage fusion to form FBGCs is an escape mechanism to avoid apoptosis [10].

Macrophages adhere to the surface of an implanted biomaterial when they are unable to phagocytose the material due to a large material-to-cell size ratio. Phagocytosis of large, nondegradable implanted materials usually does not occur due to the size disparity. When the particle size in phagocytosis >5 μm, frustrated phagocytosis may occur instead, a process in which ROIs are secreted aimed to degrade the biomaterial [10,29]. Macrophages fuse with other macrophages to form multinucleated FBGCs, associated with chronic inflammation arising from the persistent presence of a foreign body. These multinucleated cells are characteristic of granulomatous inflammation and show abundant chromatin with scattered nuclei in an irregular pattern [31]. The fusion of macrophages to form FBGCs serves to prolong the life span of these frustrated macrophages, allowing continued release of cytokines and growth factors [27].

Lymphocytes also seem to play a critical role in the FBR. They have been observed to associate with adherent macrophages and FBGCs, and enhance macrophage adhesion and fusion, while the presence of macrophages stimulates lymphocytes to proliferate [31]. Dependent on the biomaterial, next to macrophages, lymphocytes themselves also produce inflammatory mediators [10,31]. Lymphocytes enhance adherent macrophage and FBGC activation in terms of inflammatory cytokine production via paracrine (indirect) and juxtacrine (direct) means [10]. T cells have been demonstrated to promote macrophage adhesion and fusion via paracrine effects, however, close association of lymphocytes and macrophages also suggests direct signaling which has been shown to dominate at later time points of their interaction [11].

#### **4.2. Macrophage phenotype in fusion**

macrophages within the granulation tissue ensures constant remodeling of the tissue matrix

The outcome of tissue regeneration or scar formation, i.e. fibrosis, is dependent on the duration of the chronic response that contributes to cytokine production and formation of granulation tissue [8]. Fibrosis is the excessive deposition of matrix components that results in destruction of normal tissue architecture and compromised tissue function and arises from a continuous injuring stimulus, excessive synthesis or decreased degradation [33]. Synthetic and degradative functions of fibroblasts are controlled and regulated by signals from the matrix, as well as leukocyte cytokines and growth factors, wherein macrophages and their phenotype

The implantation of a biomaterial can intensify the inflammatory response by inducing a foreign body response (FBR). The FBR at the tissue-material interface is composed of macrophages and foreign body giant cells (FBGCs) and forms the end-stage response of the inflammatory and wound healing responses following implantation of a medical device, prosthesis, or biomaterial [10,29]. Typically, within 2-4 weeks the foreign material is encapsulated within an almost avascular, fibrous connective tissue, depending on the porosity of the biomaterial [52]. The FBR is characterized by the presence of macrophages and FBGCs together with the components of granulation tissue. Macrophages and FBGCs are believed to exert critical effects on both tissue and implanted material, e.g. degradation, and chemokine

Macrophages develop integrins, which play a major role in the adhesion of macrophages to a biomaterial and in the IL-4-induced macrophage fusion to form FBGCs [10]. Macrophageintegrin binding to the protein layer on the biomaterial surface provides intracellular signals that can modulate macrophage behavior, such as cytoskeletal rearrangements and the formation of adhesion structures, called podosomes. There is extensive interplay between intracellular signaling molecules activated by integrin binding and cytoskeletal proteins. Disruption of the adhesion signals promotes anoikis, i.e. apoptosis induced by cell detachment from its supportive matrix. A hypothesis is that macrophage fusion to form FBGCs is an escape

Macrophages adhere to the surface of an implanted biomaterial when they are unable to phagocytose the material due to a large material-to-cell size ratio. Phagocytosis of large, nondegradable implanted materials usually does not occur due to the size disparity. When the particle size in phagocytosis >5 μm, frustrated phagocytosis may occur instead, a process in which ROIs are secreted aimed to degrade the biomaterial [10,29]. Macrophages fuse with other macrophages to form multinucleated FBGCs, associated with chronic inflammation arising from the persistent presence of a foreign body. These multinucleated cells are

and constant recruitment of fibroblasts and endothelial cells [27].

**4. The foreign body response to biomaterials**

play an important role [27,52].

226 Calcific Aortic Valve Disease

and cytokine production [31].

mechanism to avoid apoptosis [10].

**4.1. Macrophage fusion**

The phenotype of the macrophages involved has been shown to play an important role in biomaterial scaffold remodeling [10,11,52]. The fusion of adherent macrophages to FBGCs is typically associated with a phenotype switch of the macrophages over time, going from a more pro-inflammatory activation state (M1) to a more anti-inflammatory activation state (M2). M1 versus M2 macrophage activation has led to morphological variants of multinucleated giant cells *in vitro* [10]. The M2 activation cytokines IL-4 and IL-13 promote macrophage fusion and the formation of large FBGCs with randomly arranged nuclei and high degrees of cytoplasmic spreading, while the M1 activation cytokine IFN-γ induces more limited degrees of macrophage fusion with resultant Langerhans-type giant cells. However, the activation state of fusing macrophages is neither M1-like, nor M2-like but rather an in-between state in the continuous spectrum of macrophage polarization. This suggests that biomaterial activation is unique in the process of inflammation [10,11].

The fusion of M2-activated macrophages into FBGCs is stimulated by IL-4 and IL-13, assumed to be secreted by activated T cells [11]. The precise origins of FBGC-inducing cytokines at the implant site remain unclear, with TH2 cells, NK cells, eosinophils, basophils, and mast cells as possible candidates [31]. Both IL-4 and IL-13 were found to up-regulate mannose receptors on fusing macrophages, which mediate endo- and phagocytosis, with localization of the receptor at the fusion interface [10]. MCP-1 is also involved in FBGC formation though not by recruiting cells but rather by guiding macrophage chemotaxis toward each other [10,11].

Biomaterial-adherent macrophages and FBGCs seem to show combined action of biomaterial degradation and down-modulation of pro-inflammatory mediators. Perhaps the presence of macrophage fusion and FBGC formation on biomaterial surfaces represents host downmodulation of pro-inflammatory cytokine production, possibly via phagocytic removal of macrophages actively releasing these cytokines [31]. Next to promoting M2 phenotype and macrophage fusion, IL-4 prevents apoptosis of biomaterial-adherent macrophages by inducing shedding of TNF-α receptor I, preventing this TNF-α-mediated process [11].

#### **4.3. Biomaterial degradation and fibrosis**

Macrophages and FBGCs mediate biomaterial degradation by concentrating phagocytic and oxidative activities at the interface between the cell and the biomaterial. During frustrated phagocytosis, macrophages and FBGCs release degradative mediators such as ROIs, degradative enzymes, and acid into the privileged zone between the cell membrane and the biomaterial surface such that immediate buffering or inhibition of these mediators is delayed or reduced [10]. In this process the phagocytic activity of macrophages decreases, while their degradative capacity increases [11].

biomaterial surface chemistry, scaffold architecture and mechanical properties, to the incorporation of bioactives and targeting of specific cell types. Apart from the scaffold itself, the implantation procedure contributes to the immune response. The method of implantation affects the degree of inflammation [27], and as such should be taken into account in scaffold

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Biocompatibility and thrombogeneity are particularly important during the onset of the immune response. Immediately after implantation, blood proteins precipitate onto the biomaterial surface, creating an inflammatory milieu that determines the activation of the complement and coagulation cascades. Biomaterial surface chemistry influences the proteins that adsorb, which mediates subsequent interactions with immune cells and may lead to their activation [8]. Factors that affect the amount, composition, and conformation of proteins within the initial layer include the hydrophobicity/hydrophilicity of a surface, as well as its charge and the distribution of charged groups [32]. Incorporation of anti-fouling properties into the biomaterial surface has proven an efficient method to block non-specific protein binding and promote specific biomolecule-binding. This is typically achieved by modifications with hydrophilic polymers, such as poly(ethylene glycol) (PEG), that act as molecular spacers and create a hydrophilic microenvironment that can resist non-specific protein adsorption and cell

Biomaterial surface topography and micron-scale architecture can modulate the cell-scaffold interactions that influence immune cell activation, alignment, infiltration, and fusion. Variations in surface roughness and topography affect cell adhesion, morphology, and cytokine secretion. The cell-surface interaction can change quantity and identity of secreted pro-inflammatory cytokines and chemokines, the gene expression pattern, and downstream remodeling events [11,47]. For example, one of the key cellular immune response mechanisms which can be targeted for control of biocompatibility is the mechanism for macrophage adhesion [31,54]. Macrophage fusion on biomaterial surfaces is material dependent, indicating that the surface must have an appropriate array of adsorbed proteins in order for adherent cells to adopt the necessary phenotype to fuse into FBGCs [10]. Furthermore, surface roughness of electrospun fibers has been shown to affect blood activation [55], illustrating the importance

of appropriate surface engineering, in particular in the early phases of inflammation.

Cell infiltration into the scaffold is one of the prerequisites for succesful tissue regeneration. It was shown that early infiltration of immune cells determines the degree of downstream ECM production and remodeling [56]. Cell infiltration is primarily determined by the scaffold architecture, or microstructure. Decellularized homograft/xenograft valves have shown limited cell infiltration, resulting in poor tissue remodeling and even degeneration. In contrast, decellularized *in vitro* tissue-engineered valves have shown fast repopulation with host cells and tissue remodeling following a distinct demarcation line. It has been suggested that this

design.

binding [53].

**5.2. Scaffold architecture**

**5.1. Biomaterial surface engineering**

FBGCs have the potential to be responsive to cellular signals via cell surface receptor expression as well as actively participate in the inflammatory response through the production of cytokines [10]. They produce anti-inflammatory cytokines, e.g. IL-10, which may be counterregulated by the proteolytic and pro-oxidant microenvironment. Additionally, FBGCs are thought to release pro-fibrotic factors, e.g. TGF-β and PDGF, which trigger the action of fibroblasts and endothelial cells. Continuous action of FBGCs is assumed to result in prolonged fibroblast activation and excessive biomaterial-associated matrix deposition, leading to impaired wound healing and excessive fibrosis [11]. Therefore, FBGC formation has appeared to be an undesirable phenomenon with a negative impact on biocompatibility, producing cytokines that bias wound healing cells toward a fibrogenic phenotype [31]. Efforts in the design of the biomaterial for *in situ* tissue engineering of heart valves should enhance the biocompatibility, limiting macrophage fusion into FBGCs. Surface chemistry-dependent modulation of the protein layer may enable different receptor binding and signaling in the immune cells leading to altered cellular responses, promoting wound healing while sustaining implant function [11].

### **5. Modulating the immune response**

The implantation of any biomaterial initiates an immune response. However, the extent and severity of this response can be modulated by adapting scaffold properties. As described in the previous sections, the immune response is a multi-phased cascade involving many different components. The combined effect of these components will determine the end-stage outcome of the immune response, ranging from pathological fibrotic repair to fully functional regeneration of the original tissue. By interfering with specific elements within this inflammatory cascade, the downstream outcome can be drastically affected, for better or for worse. In this plane of intersection, immuno-modulating scaffolds for *in situ* tissue engineered heart valves are being developed. The development of such a 'smart' scaffold bridges multiple length-scales and is dependent on a multitude of scaffold features. Biological scaffolds inherently come with a natural architecture and a cocktail of signaling components, which would be difficult to replicate with a synthetic counterpart. Synthetic scaffolds, on the other hand, offer a more dedicated control of individual elements in comparison to biological scaffolds. Either scaffold type can be modified within its own framework. There is a legion of possibilities to modify scaffolds over various interdependent scales, ranging from tuning biomaterial surface chemistry, scaffold architecture and mechanical properties, to the incorporation of bioactives and targeting of specific cell types. Apart from the scaffold itself, the implantation procedure contributes to the immune response. The method of implantation affects the degree of inflammation [27], and as such should be taken into account in scaffold design.

#### **5.1. Biomaterial surface engineering**

**4.3. Biomaterial degradation and fibrosis**

228 Calcific Aortic Valve Disease

degradative capacity increases [11].

implant function [11].

**5. Modulating the immune response**

Macrophages and FBGCs mediate biomaterial degradation by concentrating phagocytic and oxidative activities at the interface between the cell and the biomaterial. During frustrated phagocytosis, macrophages and FBGCs release degradative mediators such as ROIs, degradative enzymes, and acid into the privileged zone between the cell membrane and the biomaterial surface such that immediate buffering or inhibition of these mediators is delayed or reduced [10]. In this process the phagocytic activity of macrophages decreases, while their

FBGCs have the potential to be responsive to cellular signals via cell surface receptor expression as well as actively participate in the inflammatory response through the production of cytokines [10]. They produce anti-inflammatory cytokines, e.g. IL-10, which may be counterregulated by the proteolytic and pro-oxidant microenvironment. Additionally, FBGCs are thought to release pro-fibrotic factors, e.g. TGF-β and PDGF, which trigger the action of fibroblasts and endothelial cells. Continuous action of FBGCs is assumed to result in prolonged fibroblast activation and excessive biomaterial-associated matrix deposition, leading to impaired wound healing and excessive fibrosis [11]. Therefore, FBGC formation has appeared to be an undesirable phenomenon with a negative impact on biocompatibility, producing cytokines that bias wound healing cells toward a fibrogenic phenotype [31]. Efforts in the design of the biomaterial for *in situ* tissue engineering of heart valves should enhance the biocompatibility, limiting macrophage fusion into FBGCs. Surface chemistry-dependent modulation of the protein layer may enable different receptor binding and signaling in the immune cells leading to altered cellular responses, promoting wound healing while sustaining

The implantation of any biomaterial initiates an immune response. However, the extent and severity of this response can be modulated by adapting scaffold properties. As described in the previous sections, the immune response is a multi-phased cascade involving many different components. The combined effect of these components will determine the end-stage outcome of the immune response, ranging from pathological fibrotic repair to fully functional regeneration of the original tissue. By interfering with specific elements within this inflammatory cascade, the downstream outcome can be drastically affected, for better or for worse. In this plane of intersection, immuno-modulating scaffolds for *in situ* tissue engineered heart valves are being developed. The development of such a 'smart' scaffold bridges multiple length-scales and is dependent on a multitude of scaffold features. Biological scaffolds inherently come with a natural architecture and a cocktail of signaling components, which would be difficult to replicate with a synthetic counterpart. Synthetic scaffolds, on the other hand, offer a more dedicated control of individual elements in comparison to biological scaffolds. Either scaffold type can be modified within its own framework. There is a legion of possibilities to modify scaffolds over various interdependent scales, ranging from tuning

Biocompatibility and thrombogeneity are particularly important during the onset of the immune response. Immediately after implantation, blood proteins precipitate onto the biomaterial surface, creating an inflammatory milieu that determines the activation of the complement and coagulation cascades. Biomaterial surface chemistry influences the proteins that adsorb, which mediates subsequent interactions with immune cells and may lead to their activation [8]. Factors that affect the amount, composition, and conformation of proteins within the initial layer include the hydrophobicity/hydrophilicity of a surface, as well as its charge and the distribution of charged groups [32]. Incorporation of anti-fouling properties into the biomaterial surface has proven an efficient method to block non-specific protein binding and promote specific biomolecule-binding. This is typically achieved by modifications with hydrophilic polymers, such as poly(ethylene glycol) (PEG), that act as molecular spacers and create a hydrophilic microenvironment that can resist non-specific protein adsorption and cell binding [53].

Biomaterial surface topography and micron-scale architecture can modulate the cell-scaffold interactions that influence immune cell activation, alignment, infiltration, and fusion. Variations in surface roughness and topography affect cell adhesion, morphology, and cytokine secretion. The cell-surface interaction can change quantity and identity of secreted pro-inflammatory cytokines and chemokines, the gene expression pattern, and downstream remodeling events [11,47]. For example, one of the key cellular immune response mechanisms which can be targeted for control of biocompatibility is the mechanism for macrophage adhesion [31,54]. Macrophage fusion on biomaterial surfaces is material dependent, indicating that the surface must have an appropriate array of adsorbed proteins in order for adherent cells to adopt the necessary phenotype to fuse into FBGCs [10]. Furthermore, surface roughness of electrospun fibers has been shown to affect blood activation [55], illustrating the importance of appropriate surface engineering, in particular in the early phases of inflammation.

#### **5.2. Scaffold architecture**

Cell infiltration into the scaffold is one of the prerequisites for succesful tissue regeneration. It was shown that early infiltration of immune cells determines the degree of downstream ECM production and remodeling [56]. Cell infiltration is primarily determined by the scaffold architecture, or microstructure. Decellularized homograft/xenograft valves have shown limited cell infiltration, resulting in poor tissue remodeling and even degeneration. In contrast, decellularized *in vitro* tissue-engineered valves have shown fast repopulation with host cells and tissue remodeling following a distinct demarcation line. It has been suggested that this critical difference in cell infiltration is due to a lack of the dense, native-like microstructural arrangement in tissue-engineered valves, as opposed to native valves [57].

properties, the scaffold architecture and the degradation rate. The intrinsic material properties (e.g. stiffness) and the scaffold architecture (e.g. anisotropy) determine the local stresses and strains experienced by the cell. It is well recognized that mechanical conditioning is an important stimulus for ECM production and remodeling. It has been hypothesized that polymeric scaffolds can divert loads from the cells, so-called 'cell shielding', resulting in hampered ECM production. Furthermore, the scaffold or matrix stiffness can modulate the differentiation of cells into pathological phenotypes, e.g. osteoblastic or myofibroblastic, in response to mechanical and biochemical cues [5]. Therefore, efficient transduction of loads from the biomaterial to the cells is crucial. Elastomers typically exhibit adequate mechano‐ transduction properties, making them a favorable class of materials for application as synthetic

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Mechanical integrity of the scaffold is dependent on the degradation rate of the material, or rather on the balance between material degradation and ECM production. Accelerated material degradation can result in mechanical instability and valve failure. On the other hand, long-term presence of the biomaterial results in prolonged macrophage activity. Macrophages typically persist at the implantation site until the biomaterial is completely resorbed. When uncontrolled, this may lead to excessive chronic inflammation resulting in fibrosis, calcification, and/or degeneration. Mineralization of synthetic or biologic scaffolds is end-stage pathology, generally irreversible and untreatable. This underlines the importance of timely degradation of the biomaterial. Apart from proper material selection, degradation rate of polymers is tunable by varying copolymer ratios [3]. Variations in degradation kinetics of materials are also employed for controlled delivery of bioactives, for example by introducing fast-degrading fibrin gel [65] or synthetic or biological microspheres [23,66] into the scaffold.

Throughout the course of the immune response, signaling factors orchestrate the actions of the immune cells. By incorporating bioactive factors into the biomaterial scaffold, the cellular niche can be modulated locally. These biochemical factors can direct local cellular function, or promote recruitment of specific cell types via chemotaxis. Additionally, the crosstalk between immune cells and tissue cells can be enhanced, regulating the healing process [11]. Since these signaling factors play a role in a specific phase of the immune response, spatio-temporal control of growth factor or cytokine release has been the aim for many tissue engineering scaffolds. For example, long-term release of stromal cell-derived factor 1α (SDF-1α) from porous PLGA scaffolds has demonstrated to result in reduced numbers and degranulation of mast cells at the scaffold in a subcutaneous mouse model. This led to downstream alterations in the inflammatory cascade, jumpstarting regeneration with enhanced participation of progenitor cells, increased angiogenesis and decreased fibrosis [67]. De Visscher *et al.* developed heart valves constructed from photo-oxidized bovine pericardium, which were impregnated with SDF-1α in combination with fibronectin to improve the SDF-1α presentation to the cells. Implanted in the pulmonary position in sheep, these valves demonstrated improved homing of primitive cells and normal functioning at 5 months follow-up [68]. Other pro-angiogenic factors, such as VEGF, have shown to play a similar role in vascularization and endotheliali‐

scaffolds [25].

**5.4. Incorporation of bioactives**

For synthetic scaffolds, the importance of scaffold architecture is even more evident. In contrast to natural ECM, cells are typically unable to rapidly break down synthetic biomaterials in order to migrate. Therefore, cell infiltration into a synthetic scaffold is generally dependent on the available pore size, or void space [58,59]. Apart from overall cell infiltration, the pore size can also affect cell phenotype. For example, pore size has shown to be an important factor in the degree of macrophage fusion and material encapsulation. Porous implants with uniform spherical pores of 30-40 μm were shown to elicit healing with minimal fibrosis, high vascularity, and a higher M2/M1 macrophage ratio [52]. It has to be noted however, that the optimal pore size is not generic and has to be tailored to the application.

Synthetic scaffolds for heart valve tissue engineering typically consist of nano- or microfibers, with a high surface area-to-volume ratio. This fibrous architecture dictates the behavior of infiltrating cells. In addition to the void space, the fiber diameter and inter-fiber distance determine cell adhesion, spreading and proliferation [60,61]. Fiber diameter has also shown to affect platelet adhesion and coagulation activation [55]. Fiber alignment guides cell orientation and migration via contact guidance. Furthermore, it was shown that fiber alignment enhanced cell infiltration into a nanofibrous PLLA scaffold [62]. Novel processing techniques to produce fibrous 3D scaffolds with adjustable void space and/or aligned fibers (e.g. low-temperature electrospinning [63]), enhance the degrees of freedom in scaffold modification via 3D architecture (figure 4). For complex structures, such as the aortic valve, multi-layered scaffolds might be required to achieve suitable local cues [59,64].

**Figure 4.** (A) Photograph of an electrospun poly(ε-caprolactone) heart valve demonstrating 3D valve architecture, and scanning electron micrographs of its microstructure showing either random (B) or aligned microfibers (C) (scalebar = 100 µm; images courtesy of M. Simonet and G. Argento).

#### **5.3. Mechanical properties and degradation rate**

Heart valve scaffolds require appropriate mechanical properties to endure the cyclic stresses and strains exerted by the hemodynamic environment. However, next to proper functioning in the hemodynamic environment, scaffold mechanical properties play an important role on the cellular level. The macromechanical properties are determined by the intrinsic material properties, the scaffold architecture and the degradation rate. The intrinsic material properties (e.g. stiffness) and the scaffold architecture (e.g. anisotropy) determine the local stresses and strains experienced by the cell. It is well recognized that mechanical conditioning is an important stimulus for ECM production and remodeling. It has been hypothesized that polymeric scaffolds can divert loads from the cells, so-called 'cell shielding', resulting in hampered ECM production. Furthermore, the scaffold or matrix stiffness can modulate the differentiation of cells into pathological phenotypes, e.g. osteoblastic or myofibroblastic, in response to mechanical and biochemical cues [5]. Therefore, efficient transduction of loads from the biomaterial to the cells is crucial. Elastomers typically exhibit adequate mechano‐ transduction properties, making them a favorable class of materials for application as synthetic scaffolds [25].

Mechanical integrity of the scaffold is dependent on the degradation rate of the material, or rather on the balance between material degradation and ECM production. Accelerated material degradation can result in mechanical instability and valve failure. On the other hand, long-term presence of the biomaterial results in prolonged macrophage activity. Macrophages typically persist at the implantation site until the biomaterial is completely resorbed. When uncontrolled, this may lead to excessive chronic inflammation resulting in fibrosis, calcification, and/or degeneration. Mineralization of synthetic or biologic scaffolds is end-stage pathology, generally irreversible and untreatable. This underlines the importance of timely degradation of the biomaterial. Apart from proper material selection, degradation rate of polymers is tunable by varying copolymer ratios [3]. Variations in degradation kinetics of materials are also employed for controlled delivery of bioactives, for example by introducing fast-degrading fibrin gel [65] or synthetic or biological microspheres [23,66] into the scaffold.

#### **5.4. Incorporation of bioactives**

critical difference in cell infiltration is due to a lack of the dense, native-like microstructural

For synthetic scaffolds, the importance of scaffold architecture is even more evident. In contrast to natural ECM, cells are typically unable to rapidly break down synthetic biomaterials in order to migrate. Therefore, cell infiltration into a synthetic scaffold is generally dependent on the available pore size, or void space [58,59]. Apart from overall cell infiltration, the pore size can also affect cell phenotype. For example, pore size has shown to be an important factor in the degree of macrophage fusion and material encapsulation. Porous implants with uniform spherical pores of 30-40 μm were shown to elicit healing with minimal fibrosis, high vascularity, and a higher M2/M1 macrophage ratio [52]. It has to be noted however, that the

Synthetic scaffolds for heart valve tissue engineering typically consist of nano- or microfibers, with a high surface area-to-volume ratio. This fibrous architecture dictates the behavior of infiltrating cells. In addition to the void space, the fiber diameter and inter-fiber distance determine cell adhesion, spreading and proliferation [60,61]. Fiber diameter has also shown to affect platelet adhesion and coagulation activation [55]. Fiber alignment guides cell orientation and migration via contact guidance. Furthermore, it was shown that fiber alignment enhanced cell infiltration into a nanofibrous PLLA scaffold [62]. Novel processing techniques to produce fibrous 3D scaffolds with adjustable void space and/or aligned fibers (e.g. low-temperature electrospinning [63]), enhance the degrees of freedom in scaffold modification via 3D architecture (figure 4). For complex structures, such as the aortic valve,

**Figure 4.** (A) Photograph of an electrospun poly(ε-caprolactone) heart valve demonstrating 3D valve architecture, and scanning electron micrographs of its microstructure showing either random (B) or aligned microfibers (C) (scalebar =

Heart valve scaffolds require appropriate mechanical properties to endure the cyclic stresses and strains exerted by the hemodynamic environment. However, next to proper functioning in the hemodynamic environment, scaffold mechanical properties play an important role on the cellular level. The macromechanical properties are determined by the intrinsic material

100 µm; images courtesy of M. Simonet and G. Argento).

**5.3. Mechanical properties and degradation rate**

arrangement in tissue-engineered valves, as opposed to native valves [57].

230 Calcific Aortic Valve Disease

optimal pore size is not generic and has to be tailored to the application.

multi-layered scaffolds might be required to achieve suitable local cues [59,64].

Throughout the course of the immune response, signaling factors orchestrate the actions of the immune cells. By incorporating bioactive factors into the biomaterial scaffold, the cellular niche can be modulated locally. These biochemical factors can direct local cellular function, or promote recruitment of specific cell types via chemotaxis. Additionally, the crosstalk between immune cells and tissue cells can be enhanced, regulating the healing process [11]. Since these signaling factors play a role in a specific phase of the immune response, spatio-temporal control of growth factor or cytokine release has been the aim for many tissue engineering scaffolds. For example, long-term release of stromal cell-derived factor 1α (SDF-1α) from porous PLGA scaffolds has demonstrated to result in reduced numbers and degranulation of mast cells at the scaffold in a subcutaneous mouse model. This led to downstream alterations in the inflammatory cascade, jumpstarting regeneration with enhanced participation of progenitor cells, increased angiogenesis and decreased fibrosis [67]. De Visscher *et al.* developed heart valves constructed from photo-oxidized bovine pericardium, which were impregnated with SDF-1α in combination with fibronectin to improve the SDF-1α presentation to the cells. Implanted in the pulmonary position in sheep, these valves demonstrated improved homing of primitive cells and normal functioning at 5 months follow-up [68]. Other pro-angiogenic factors, such as VEGF, have shown to play a similar role in vascularization and endotheliali‐ zation via recruitment of bone marrow-derived circulating cells, with an essential paracrine role for myeloid cells [69]. Injectable hydrogels featuring a sustained release of VEGF, either or not combined with PDGF, have shown to enhance angiogenesis [70,71]. Dual delivery of MCP-1 and VEGF was applied to promote early monocyte invasion as well as angiogenesis. This was shown to increase mature vessel formation via enhanced endothelial and smooth muscle cell recruitment and displayed a trend of macrophage polarization to the M2 type in a time- and dose-dependent manner [66]. MCP-1 has demonstrated to be a potent immunemodulatory factor, leading to successful remodeling and regeneration of a PCL/PLLA blood vessel graft in mice [23]. Decellularized porcine aortic valves coated with a fusion protein of fibronectin and hepatocyte growth factor (HGF) demonstrated modest acceleration of infiltration of tissue cells, particularly in the valve leaflet, after implantation in a dog model [72].

recognized as a natural binding site for many growth factors and cytokines, a feature which has been exploited by developing heparin-mimetic peptide nanofibers that are capable of

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Clearly, the use of bioactives or on-the-fly harvestable cells is a powerful tool to create immunemodulating scaffolds. Methods using covalent immobilization of factors [65], microspheres with controlled degradation profiles [66] or hollow-fiber electrospinning techniques [59] enable optimized spatio-temporal control of one or multiple factors. Furthermore, advanced hydrogels have been developed to offer on-demand, remote-controlled release using a magnetic field [75]. With state-of-the-art supramolecular polymers it is possible to engineer cell-responsive substrates [76], offering truly 'smart' scaffolds that can interact with their

The inflammatory response is mainly driven by colonization of the scaffold by blood-derived cells. The nature of the infiltrating cells and their differentiation were demonstrated of pivotal importance to control the delicate balance between fibrotic or functional regenerated ECM production. Of all the cells involved in the immune response, several cell populations can be

Macrophages are the predominant mediators throughout the entire immune response, making them an attractive therapeutic target [77]. Although, the precise nature of macrophage plasticity and polarization has yet to be illuminated, it has been shown that early macrophage phenotype determines the end-stage outcome in various biological matrices. In particular, an increased ratio of M2/M1 macrophages correlates to enhanced remodeling, which is likely mediated by differential attraction of secondary cells [78]. By promoting the M2 phenotype, either via specific recruitment or local polarization, the inflammatory response may instantly be directed towards healing instead of inflammation [48,52]. With the identification of multiple subtypes, it is likely that the various macrophage phenotypes play a critical role throughout the various stages of acute inflammation, the healing phase and the resolution of inflammation

ECM production and remodeling is governed by the attraction of secondary cells to the scaffold, consisting of mature (myo-)fibroblasts and endothelial cells, as well as various stem/progenitor cells, released into the circulation by the bone marrow. Furthermore, it has been suggested that adult valve interstitial cells are continuously replenished via circulat‐ ing endothelial or mesenchymal cell precursors derived from the bone marrow and subsequently undergo endothelial-to-mesenchymal-transition (EndoMT) [5]. Circulating progenitors, such as endothelial progenitor cells (EPC), can have a significant influence on the inflammatory response [4,12,79]. EPC are hypothesized to be an important target cell for endothelialization. Rapid formation of an endothelium over a scaffold is desirable as it acts as a dynamic and selective barrier by maintaining a nonthrombogenic surface, controlling the transfer of molecules across the layer, and regulating immune and inflamma‐ tory reactions. The endothelial layer also interacts with underlying cells to regulate their

binding growth factors such as VEGF and HGF [74].

environment to mediate the host response to the biomaterial.

identified as target cells for *in situ* heart valve tissue engineering.

**5.5. Cell recruitment and differentiation**

(figure 5).

growth and proliferation [12].

Single extracellular molecules can impact both pro-inflammatory and anti-inflammatory pathways in different cell types participating in the repair response [35]. The shift from proinflammatory to anti-inflammatory response is generally mediated by lipoxins, protectins, and resolvins, actively promoting resolution of infection and tissue repair. Lipoxins are arachidonic acid (AA) derivatives generated by lipoxygenases, and stop the influx of neutrophils, promote the uptake of apoptotic neutrophils by macrophages and recruit additional monocytes to help clear away dead cells and tissue debris [28,39]. Incorporation of such bioactive components may enhance the resolution of the inflammatory response, avoiding uncontrolled chronic inflammation. Resolution of inflammation is also mediated by glucocorticoids, which inhibit inflammatory cell activation by withdrawing the synthesis of inflammatory mediators, and promote resolution of inflammation by enhancing anti-inflammatory cytokine release [29,54]. Glucocorticoids have been shown to modulate the phenotype of infiltrating macrophages and lymphocytes and could thus be used locally to regulate the cellular response [54].

An alternative to incorporating specific signaling moieties into a scaffold is to preseed the scaffold with cells that act as natural signaling factories. Cells, typically bone marrow-derived mononuclear cells, harvested from the host are directly seeded into a scaffold, which is subsequently implanted in a single operation. Although the preseeded cells are cleared from the scaffold within several days after implantation, they mediate the immune response via paracrine signaling by secreting a natural cocktail of growth factors and cytokines. This approach has shown prosperous results in clinical trials using synthetic blood vessel grafts [24]. Furthermore, a similar approach using decellularized tissue-engineered heart valves has shown promising short-term results after 4 week implantation in the pulmonary position in non-human primates [73].

Apart from boosting selected signaling molecules, biomaterials can be designed to tether endogenously released factors to promote a regenerative microenvironment. Natural occurring GAGs have been identified to bind and modify inflammatory factors like interleukins and chemokines, e.g. IL-10 [11]. Subtle differences in GAG structure and/or sequence might be sensed by signaling molecules, guiding their interaction with the ECM and mediating their presentation to leukocytes [11]. In this way, physiological cytokine concentrations are ensured, reducing the risk of adverse side-effects. Heparan sulfate is wellrecognized as a natural binding site for many growth factors and cytokines, a feature which has been exploited by developing heparin-mimetic peptide nanofibers that are capable of binding growth factors such as VEGF and HGF [74].

Clearly, the use of bioactives or on-the-fly harvestable cells is a powerful tool to create immunemodulating scaffolds. Methods using covalent immobilization of factors [65], microspheres with controlled degradation profiles [66] or hollow-fiber electrospinning techniques [59] enable optimized spatio-temporal control of one or multiple factors. Furthermore, advanced hydrogels have been developed to offer on-demand, remote-controlled release using a magnetic field [75]. With state-of-the-art supramolecular polymers it is possible to engineer cell-responsive substrates [76], offering truly 'smart' scaffolds that can interact with their environment to mediate the host response to the biomaterial.

#### **5.5. Cell recruitment and differentiation**

zation via recruitment of bone marrow-derived circulating cells, with an essential paracrine role for myeloid cells [69]. Injectable hydrogels featuring a sustained release of VEGF, either or not combined with PDGF, have shown to enhance angiogenesis [70,71]. Dual delivery of MCP-1 and VEGF was applied to promote early monocyte invasion as well as angiogenesis. This was shown to increase mature vessel formation via enhanced endothelial and smooth muscle cell recruitment and displayed a trend of macrophage polarization to the M2 type in a time- and dose-dependent manner [66]. MCP-1 has demonstrated to be a potent immunemodulatory factor, leading to successful remodeling and regeneration of a PCL/PLLA blood vessel graft in mice [23]. Decellularized porcine aortic valves coated with a fusion protein of fibronectin and hepatocyte growth factor (HGF) demonstrated modest acceleration of infiltration of tissue cells, particularly in the valve leaflet, after implantation in a dog model

Single extracellular molecules can impact both pro-inflammatory and anti-inflammatory pathways in different cell types participating in the repair response [35]. The shift from proinflammatory to anti-inflammatory response is generally mediated by lipoxins, protectins, and resolvins, actively promoting resolution of infection and tissue repair. Lipoxins are arachidonic acid (AA) derivatives generated by lipoxygenases, and stop the influx of neutrophils, promote the uptake of apoptotic neutrophils by macrophages and recruit additional monocytes to help clear away dead cells and tissue debris [28,39]. Incorporation of such bioactive components may enhance the resolution of the inflammatory response, avoiding uncontrolled chronic inflammation. Resolution of inflammation is also mediated by glucocorticoids, which inhibit inflammatory cell activation by withdrawing the synthesis of inflammatory mediators, and promote resolution of inflammation by enhancing anti-inflammatory cytokine release [29,54]. Glucocorticoids have been shown to modulate the phenotype of infiltrating macrophages and

lymphocytes and could thus be used locally to regulate the cellular response [54].

An alternative to incorporating specific signaling moieties into a scaffold is to preseed the scaffold with cells that act as natural signaling factories. Cells, typically bone marrow-derived mononuclear cells, harvested from the host are directly seeded into a scaffold, which is subsequently implanted in a single operation. Although the preseeded cells are cleared from the scaffold within several days after implantation, they mediate the immune response via paracrine signaling by secreting a natural cocktail of growth factors and cytokines. This approach has shown prosperous results in clinical trials using synthetic blood vessel grafts [24]. Furthermore, a similar approach using decellularized tissue-engineered heart valves has shown promising short-term results after 4 week implantation in the pulmonary position in

Apart from boosting selected signaling molecules, biomaterials can be designed to tether endogenously released factors to promote a regenerative microenvironment. Natural occurring GAGs have been identified to bind and modify inflammatory factors like interleukins and chemokines, e.g. IL-10 [11]. Subtle differences in GAG structure and/or sequence might be sensed by signaling molecules, guiding their interaction with the ECM and mediating their presentation to leukocytes [11]. In this way, physiological cytokine concentrations are ensured, reducing the risk of adverse side-effects. Heparan sulfate is well-

[72].

232 Calcific Aortic Valve Disease

non-human primates [73].

The inflammatory response is mainly driven by colonization of the scaffold by blood-derived cells. The nature of the infiltrating cells and their differentiation were demonstrated of pivotal importance to control the delicate balance between fibrotic or functional regenerated ECM production. Of all the cells involved in the immune response, several cell populations can be identified as target cells for *in situ* heart valve tissue engineering.

Macrophages are the predominant mediators throughout the entire immune response, making them an attractive therapeutic target [77]. Although, the precise nature of macrophage plasticity and polarization has yet to be illuminated, it has been shown that early macrophage phenotype determines the end-stage outcome in various biological matrices. In particular, an increased ratio of M2/M1 macrophages correlates to enhanced remodeling, which is likely mediated by differential attraction of secondary cells [78]. By promoting the M2 phenotype, either via specific recruitment or local polarization, the inflammatory response may instantly be directed towards healing instead of inflammation [48,52]. With the identification of multiple subtypes, it is likely that the various macrophage phenotypes play a critical role throughout the various stages of acute inflammation, the healing phase and the resolution of inflammation (figure 5).

ECM production and remodeling is governed by the attraction of secondary cells to the scaffold, consisting of mature (myo-)fibroblasts and endothelial cells, as well as various stem/progenitor cells, released into the circulation by the bone marrow. Furthermore, it has been suggested that adult valve interstitial cells are continuously replenished via circulat‐ ing endothelial or mesenchymal cell precursors derived from the bone marrow and subsequently undergo endothelial-to-mesenchymal-transition (EndoMT) [5]. Circulating progenitors, such as endothelial progenitor cells (EPC), can have a significant influence on the inflammatory response [4,12,79]. EPC are hypothesized to be an important target cell for endothelialization. Rapid formation of an endothelium over a scaffold is desirable as it acts as a dynamic and selective barrier by maintaining a nonthrombogenic surface, controlling the transfer of molecules across the layer, and regulating immune and inflamma‐ tory reactions. The endothelial layer also interacts with underlying cells to regulate their growth and proliferation [12].

replacement causes extensive tissue damage, giving rise to DAMPs, which prime the system for an enhanced immune response [29]. As an alternative, various transvascular, catheterbased techniques, as well as alternative minimally invasive surgical techniques, such as the transapical approach, have been developed [4,80,81]. This has implication for the scaffold design as the scaffold must be crimped and incorporated into a stent. Upon delivery at the valve annulus, the scaffold must also be able to expand properly, be held in place and instantly function within the hemodynamic environment. Transapical valve implantation of preseeded decellularized tissue engineered heart valves into both the aortic and pulmonary position has

The Immune Response in *In Situ* Tissue Engineering of Aortic Heart Valves

http://dx.doi.org/10.5772/54354

235

*In situ* tissue engineering of heart valves represents a quick, cheap, and on-demand approach. Immunomodulatory scaffolds hold great promise for future application and commercializa‐ tion. However, some priority challenges remain to be addressed in the translation from bench

One of the main challenges for *in situ* tissue engineering is to stimulate functional ECM formation without inducing fibrosis. To maintain functionality of the valve, rapid ECM formation is required in order to overtake the load-bearing role of the degrading scaffold. However, cells and molecules that are stimulatory for ECM production have been designated as pro-fibrotic mediators. This poses a paradoxal challenge. Macrophage plasticity is a striking example. M2 macrophages have been identified as pro-wound healing cells, promoting ECM production by secretion of IL-4 and TGF-β. On the other hand, both IL-4 and TGF-β are strong inducers of fibrosis if not tightly regulated. Chemokines, such as MCP-1, have been identified as pro-fibrotic mediators by attracting fibrocytes and stimulation of M2 polarization [84,85]. On the other hand, MCP-1 inhibition leads to delayed or inhibited wound healing [86]. Fibrocytes are blood-borne mesenchymal stem cell progenitors with a fibroblast/myofibro‐

/collagen type I+

ECM formation and fibrosis. The same holds for EndoMT-derived (myo-)fibroblasts. However, the local activation state of recruited myofibroblasts, rather than the source, determines their ECM remodeling activity. For example, TLR-signaling promotes fibroblasts to differentiate into collagen-producing myofibroblasts [84]. Valvular interstitial cells (VICs) in the adult valve have a quiescent myofibroblast-like phenotype. Regulating the activation state of colonizing myofibroblasts in the scaffold is pivotal in the prevention of fibrosis and obtaining a VIC-like population. The TGF-β pathway is one of the main players in this process. Furthermore, IL-10 has been shown to inhibit fibrosis in numerous animal models [84], underlining that timely resolution of inflammation is one of the main challenges for *in situ*

) that similarly have been related to both

already proven feasible in pre-clinical models [82,83].

**6. Challenges and pitfalls**

**6.1. ECM formation versus fibrosis**

blast-like phenotype (CD34+

tissue engineering.

/CD45+

to bed.

**Figure 5.** Hypothesized role of the various macrophage polarization states throughout the process of inflammation and tissue regeneration in response to scaffold implantation. Illustration by Anthal Smits.

Mesenchymal stem cells (MSCs) proliferate during the healing phase, directed by cytokines secreted by nearby cells, e.g. activated platelets and macrophages, and by ECM components such as collagen peptides and fibronectin [29]. MSCs produce an immunoprivileged environment by preventing the activation and proliferation of DCs, T cells, macrophages, and PMNs through direct cell-cell interactions and paracrine signaling [8]. Cells derived from immunoprivileged regions have been delivered to promote cell engraftment and protect grafts against autoimmune and allogeneic rejection. These cells secrete a range of factors, eg. TGFβ and IL-10, inducing regulatory T-cell differentiation/expansion, which enhances immuno‐ protection [8].

Recruitment and adhesion of target cell types can be achieved by offering binding domains on the scaffold, for example using supramolecular building blocks with cell-specific peptide sequences [76,79]. When combined with anti-fouling materials, such as PEG, this results in highly selective substrates.

#### **5.6. Minimally invasive implantation methods**

Independent of the biomaterial, the injury incurred during the implantation process will trigger an immune response, due to the disruption of host tissue and induction of cell damage. Besides substantial mortality and morbidity risks, invasive open heart surgery for heart valve replacement causes extensive tissue damage, giving rise to DAMPs, which prime the system for an enhanced immune response [29]. As an alternative, various transvascular, catheterbased techniques, as well as alternative minimally invasive surgical techniques, such as the transapical approach, have been developed [4,80,81]. This has implication for the scaffold design as the scaffold must be crimped and incorporated into a stent. Upon delivery at the valve annulus, the scaffold must also be able to expand properly, be held in place and instantly function within the hemodynamic environment. Transapical valve implantation of preseeded decellularized tissue engineered heart valves into both the aortic and pulmonary position has already proven feasible in pre-clinical models [82,83].
