**2. Tissue engineering**

The different drawbacks related to commercial replacement devices compromise their durability once in the patient and have shift the attention of cardiac surgeons and biomedical engineers towards a new therapeutic concept: heart valve tissue engineering. The first general definition of this approach has been proposed by Langer and Vacanti, as the *in vitro* creation

of a viable tissue by combining separate elements, i.e. cells plus an extracellular matrix (ECM), properly conditioned to attain the correct mature function [4]. This universal paradigm has to be applied also to the reconstruction of the valve tissue. The rationale is to achieve the ability to construct *in vitro* a valve with adequate biomechanical properties, good hemodynamic performance, vital competence, growth/remodelling permissiveness and lack of inflammato‐ ry/immunological reactions. Such researches have required the synergistic application of different scientific disciplines, from cell biology to engineering and surgery. *In vitro* valve creation is commonly pursued via two different methods diverging for the starter matrix.

the polymer segments has been carried out and is still under study to improve the durability of polyurethane-based devices *in vivo* [8, 10]. In particular, when modified with polyhe‐ dral oligomeric silsesquioxanes (POSS), polycarbonate urethane-composed polymeric substitutes offer improved biological and hydrodynamic functioning in respect to biopros‐

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Shinoka et al. firstly reported in 1995 the application of aliphatic polyesters in TEHV formu‐ lation. The constructs were composed of polyglactin, polyglycolic acid (PGA) or polylactic acid (PLA) [12-14]. The scarce pliability of these biomaterials did not allow, however, a perfect shape modelling [15]. Conversely, polyhydroxyalcanoates, as polyhydroxyoctanoate (PHA) mixed or not with poly-4-hydroxybutyrate (PH4B), have demonstrated better thermoplastic proficiencies: a polyester group combined with bacterial-derived hydroxyacids is the chemical

Polystyrene/polyisobutylene compounds were also developed for cardiovascular structure fabrication, showing superior resistance to the high environmental heart valve stresses [18].

Novel promising biopolymers are currently tested to better mimic chemo-physical properties of native heart valves: *inter alia*, polyvinyl alcohol-bacterial cellulose-based hydrogels can be

Contemporary procedures for polymeric aortic heart valve fabrication must rely on optimal tricuspid design, used as template for successive valve production. Multiple dip-coatings into poorly concentrated polymeric solutions have as major drawback inhomogeneous tissue

Another manufacturing technique combines the use of solvent and thermal treatment to properly shape polymeric films in a desired arrangement. Tri-dimensional models, addition‐ ally created through direct laser 'recording' in photo-sensitive polymers, act as blueprint for successive casting of hot biopolymers, which will assume the chosen conformation during chilling [20]. Similarly, injection systems assisted by hot/cold baths can be applied to the same

Despite the evident ability of these biopolymers to undergo cell remodelling, a proper mature, cell-operated tissue architectural reconstitution might require a chemo-mechanical stimulation for a long time. Often, to the best of their biostability and mechanical behaviour, biopolymeric heart valves do not develop a trilaminate structure and remarkably, the elastin network in

On this account, another research stream inside the heart valve tissue engineering approach prefers the usage of animal-derived decellularized scaffolds. A suitable decellularization procedure allows the removal of xenogeneic cells, maintaining all the fibre composition and distribution of the natural ECM. In addition, as further advantage in respect to commonly produced bioprostheses, the treatment with glutaraldehyde can be skipped following the absence of xenogeneic cells. The avoidance of this cytotoxic agent enables remodelling

leaflet layer *ventricularis* and wall *media* can be inconsistently achieved.

opportunely modelled for a broad range of tuneable mechanical properties [19].

thetic valves [11].

thickness.

aim [21].

composition of these last polymers [16,17].

*2.1.2. Decellularized extracellular matrices*

#### **2.1. Biomaterial scaffolds**

#### *2.1.1. Polymeric materials*

The choice of the ECM scaffold is not only a distinction parameter among diverse approaches, but also essential for the successful realization of tissue-engineered heart valves (TEHVs). ECM is able to establish the necessary 3D configuration and guides cell attachment and structural development of the new tissue.

Synthetic materials have often represented the privileged option in TEHV formulations. Biopolymers as aliphatic polyesters, polyhydroxyalcanoates or different polyurethane compounds have been preferentially employed, providing scaffolds with controllable chemophysical characteristics as reproducibility, porousness and biodegradability rate. The chosen biomaterials have to respond to important requirements, as good cell-affinity and adequate structural architecture able to sustain the organ mechanics. To enable cell adhesion and spreading in the selected biopolymeric mesh, the modulation of porousness until 90% achievement is recommended [5].

In particular, most of these materials offer a further regenerative advantage also thanks to the good immunotolerance induced in the host body. In fact, after an initial guiding effect, the non-natural material is progressively degraded by colonizing cells, which in turn operate a new matrix synthesis. The process results in a newly produced tissue of completely autologous origin.

First generation polymeric heart valves were designed to overcome the poor durability and excessive wear shown by Teflon fabric-composed substitutes. These more rigid valves, with caged-ball or low-profile design, were based either on metals, like titanium or stellite 21, or on silicone with fixed fabric sewing rings. While metallic devices mostly presented difficulties in the insertion phase, elastomeric ball valves demonstrated less stability in the mid/long-term evaluation [6]. Polyurethanes were lately proposed for their relatively good haemodynamic behaviour especially in contact with blood cells and indeed for their easy manipulable chemical structure [7,8].

Again, biostability represented the major drawback associated to these elastomers togeth‐ er to the high calcification potential: polyester, polyether and polycarbonate urethanes were sequentially suggested for valve fabrication with minor biodegradation, but still insuffi‐ cient stability [9]. The introduction of further chemical groups and other modifications in the polymer segments has been carried out and is still under study to improve the durability of polyurethane-based devices *in vivo* [8, 10]. In particular, when modified with polyhe‐ dral oligomeric silsesquioxanes (POSS), polycarbonate urethane-composed polymeric substitutes offer improved biological and hydrodynamic functioning in respect to biopros‐ thetic valves [11].

Shinoka et al. firstly reported in 1995 the application of aliphatic polyesters in TEHV formu‐ lation. The constructs were composed of polyglactin, polyglycolic acid (PGA) or polylactic acid (PLA) [12-14]. The scarce pliability of these biomaterials did not allow, however, a perfect shape modelling [15]. Conversely, polyhydroxyalcanoates, as polyhydroxyoctanoate (PHA) mixed or not with poly-4-hydroxybutyrate (PH4B), have demonstrated better thermoplastic proficiencies: a polyester group combined with bacterial-derived hydroxyacids is the chemical composition of these last polymers [16,17].

Polystyrene/polyisobutylene compounds were also developed for cardiovascular structure fabrication, showing superior resistance to the high environmental heart valve stresses [18].

Novel promising biopolymers are currently tested to better mimic chemo-physical properties of native heart valves: *inter alia*, polyvinyl alcohol-bacterial cellulose-based hydrogels can be opportunely modelled for a broad range of tuneable mechanical properties [19].

Contemporary procedures for polymeric aortic heart valve fabrication must rely on optimal tricuspid design, used as template for successive valve production. Multiple dip-coatings into poorly concentrated polymeric solutions have as major drawback inhomogeneous tissue thickness.

Another manufacturing technique combines the use of solvent and thermal treatment to properly shape polymeric films in a desired arrangement. Tri-dimensional models, addition‐ ally created through direct laser 'recording' in photo-sensitive polymers, act as blueprint for successive casting of hot biopolymers, which will assume the chosen conformation during chilling [20]. Similarly, injection systems assisted by hot/cold baths can be applied to the same aim [21].

#### *2.1.2. Decellularized extracellular matrices*

of a viable tissue by combining separate elements, i.e. cells plus an extracellular matrix (ECM), properly conditioned to attain the correct mature function [4]. This universal paradigm has to be applied also to the reconstruction of the valve tissue. The rationale is to achieve the ability to construct *in vitro* a valve with adequate biomechanical properties, good hemodynamic performance, vital competence, growth/remodelling permissiveness and lack of inflammato‐ ry/immunological reactions. Such researches have required the synergistic application of different scientific disciplines, from cell biology to engineering and surgery. *In vitro* valve creation is commonly pursued via two different methods diverging for the starter matrix.

The choice of the ECM scaffold is not only a distinction parameter among diverse approaches, but also essential for the successful realization of tissue-engineered heart valves (TEHVs). ECM is able to establish the necessary 3D configuration and guides cell attachment and structural

Synthetic materials have often represented the privileged option in TEHV formulations. Biopolymers as aliphatic polyesters, polyhydroxyalcanoates or different polyurethane compounds have been preferentially employed, providing scaffolds with controllable chemophysical characteristics as reproducibility, porousness and biodegradability rate. The chosen biomaterials have to respond to important requirements, as good cell-affinity and adequate structural architecture able to sustain the organ mechanics. To enable cell adhesion and spreading in the selected biopolymeric mesh, the modulation of porousness until 90%

In particular, most of these materials offer a further regenerative advantage also thanks to the good immunotolerance induced in the host body. In fact, after an initial guiding effect, the non-natural material is progressively degraded by colonizing cells, which in turn operate a new matrix synthesis. The process results in a newly produced tissue of completely autologous

First generation polymeric heart valves were designed to overcome the poor durability and excessive wear shown by Teflon fabric-composed substitutes. These more rigid valves, with caged-ball or low-profile design, were based either on metals, like titanium or stellite 21, or on silicone with fixed fabric sewing rings. While metallic devices mostly presented difficulties in the insertion phase, elastomeric ball valves demonstrated less stability in the mid/long-term evaluation [6]. Polyurethanes were lately proposed for their relatively good haemodynamic behaviour especially in contact with blood cells and indeed for their easy manipulable chemical

Again, biostability represented the major drawback associated to these elastomers togeth‐ er to the high calcification potential: polyester, polyether and polycarbonate urethanes were sequentially suggested for valve fabrication with minor biodegradation, but still insuffi‐ cient stability [9]. The introduction of further chemical groups and other modifications in

**2.1. Biomaterial scaffolds**

248 Calcific Aortic Valve Disease

*2.1.1. Polymeric materials*

development of the new tissue.

achievement is recommended [5].

origin.

structure [7,8].

Despite the evident ability of these biopolymers to undergo cell remodelling, a proper mature, cell-operated tissue architectural reconstitution might require a chemo-mechanical stimulation for a long time. Often, to the best of their biostability and mechanical behaviour, biopolymeric heart valves do not develop a trilaminate structure and remarkably, the elastin network in leaflet layer *ventricularis* and wall *media* can be inconsistently achieved.

On this account, another research stream inside the heart valve tissue engineering approach prefers the usage of animal-derived decellularized scaffolds. A suitable decellularization procedure allows the removal of xenogeneic cells, maintaining all the fibre composition and distribution of the natural ECM. In addition, as further advantage in respect to commonly produced bioprostheses, the treatment with glutaraldehyde can be skipped following the absence of xenogeneic cells. The avoidance of this cytotoxic agent enables remodelling processes to occur by providing a cell-friendly milieu, where viable engrafting elements are able to accomplish synthetic and contractile functions with extracellular matrix continuous remodelling.

Extracting processes should be hence developed ad hoc depending on the specific tissue to use as starter matrix, otherwise adversely affecting the original mechanical and bioactive proper‐

Novel Therapeutic Strategies for the Treatment of End-Stage Valvulopathies

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251

In addition, the realization of viable constructs must not be regardless of detergent scaffold retention. Incomplete washout of detergents could induce the creation of a toxic microenvir‐ onment for engrafting cells. *In vitro* evaluations should be critically performed in respect to

Cells embody the second key-component of a TEHV: it is this element that provides viability to the ECM and consequently permits its remodelling and maturation towards a functional organ. Furthermore, the use of cell elements of allogeneic or autologous source can prevent from non-self reactions towards synthetic biomaterials or decellularized natural tissues.

For a similar selection principle operated by those researchers preferring the more committed animal-derived matrix, endothelial cells, fibroblasts, myofibroblasts and/or smooth muscle cells isolated from vascular or valve conduits have been extensively utilized to seed nude matrices and obtain both endothelial coverage with antithrombotic activity and tissue

Endothelial cells (ECs) were first applied in the '90s to test the feasibility of endothelialisation on Biomer and Mitrathane thromboresistant polyurethane ureas in response to physiological shear stress [38]. Commonly harvested from cardiovascular structures, as adult saphenous or more immature umbilical veins, endothelial cells were also seeded onto bioprosthetic heart valves to increase their low thrombogenic properties [39]. In an analogous fashion, a previous endothelial coverage on harshly decellularized native tissues can avert *in vivo* thromboembolic

More frequently, ECs were employed in combination to other cell elements, as for example myo/fibroblasts. A sequential seeding of fibroblasts and ECs was demonstrated effective in the *in vitro* creation of tissue-engineered valve constructs endowed with appropriate cell topography [40]. However, the district of fibroblast cell origin is able to significantly affect the degree of recellularization with better outcome associated to the usage of arterial myofibro‐

Valve fibroblasts, known as valve interstitial cells (VICs) for their cusp origin, have been successfully employed as sole repopulating cell population, exhibiting unmodified prolifera‐

While dissimilar to VICs, smooth muscle cells of vascular derivation (vSMCs) can be chemi‐ cally manipulated with epidermal growth factor (EGF), platelet-derived growth factor (PDGF) and transforming growth factor beta-1 (TGF-beta1) for a phenotypic switch towards leaflet

tive and synthetic abilities once engrafted in decellularized scaffolds [24].

events, related to basal membrane damage/loss resulting in collagen fibres exposure.

each decellularization procedure currently applied [unpublished data, 35, 36].

ties of the natural ECM.

*2.2.1. Differentiated cells*

repopulation [24, 37].

cells [43].

blasts rather than of dermal fibroblasts [41, 42].

**2.2. Cells**

Cell-freed natural matrices can be realised by means of several methods: trypsin-based enzymatic and detergent decellularization procedures are only two examples of the proposed treatments. Most decellularizing protocols take advantage of a combined mechanical and chemical tissue handling to ease cell removal.

Grauss and colleagues interestingly compared various protocols currently applied to decellularize porcine aortic heart valves and verified that the combination of trypsin and Triton X-100, an anionic detergent, could be able to provoke a loss of matrix integrity [22]. Not only the enzymatic treatment with trypsin can induce elastin defragmentation, but also the use of sodium dodecyl sulphate can end out with a similar deleterious effect [23]. Conversely, the adoption of sodium cholate- and deoxycholate-based methods allows the achievement of fully nude matrices, able to be cell-recolonized *in vitro* and/or *in vivo* even for clinical implantation [23-28].

Besides native heart valves, another natural tissue has been regarded with attention for the production of animal-derived acellular scaffolds. Pericardium has been extensively applied in bioprosthetic manufacturing thanks to its biocompatible, mechanical and biological proper‐ ties, entirely suitable for long-lasting heart valve substitutes. More often of bovine origin, this tissue partially differs from a native heart valve for its low cellularity and extremely compacted ECM, raising the question on which decellularization process can best convey the optimal outcome [29]. With regards to this tissue, the comparative analysis developed by Yang et al. revealed a superior decellularizing and preserving effect of enzymatic/detergent treatment or trypsin alone on Triton X100/sodium-deoxycholate-based extraction.

Sodium cholate demonstrates instead a less aggressive behaviour towards the pericardial tissue providing analogous results to its application on native semilunar heart valves [30].

Together to the elimination of xenogeneic cells, a variable depletion in the glycosaminoglycan (GAG) content is usually observed after decellularization procedure. GAGs play a significant viscoelastic role by mitigating valve stress during flexion: their highly hydrophilic nature allows, in fact, the hydration of the *spongiosa* layer in this cycle phase. Indeed, several biological processes appear to be modulated by GAGs, therefore their loss has profound consequences in the mechanical behaviour and in the cellular functions of successively repopulated scaffolds [31]. Associated to GAG content is also the hydrated state of the tissue: a reduction in hydration following decellularization can induce collapse of collagen fibres and introduces a less suitable environment for colonizing cells [32], 33].

While GAG and water preservation is pivotal, an opposite decellularizing effect is expected as far as DNA/RNA content. Similar to cell membrane residues, the phosphate groups of the nucleic acid backbone behave as powerful calcification triggers [34]. Aspecific endonucleases are often applied as efficient tools for the complete removal of nucleic acid debris [25-26].

Extracting processes should be hence developed ad hoc depending on the specific tissue to use as starter matrix, otherwise adversely affecting the original mechanical and bioactive proper‐ ties of the natural ECM.

In addition, the realization of viable constructs must not be regardless of detergent scaffold retention. Incomplete washout of detergents could induce the creation of a toxic microenvir‐ onment for engrafting cells. *In vitro* evaluations should be critically performed in respect to each decellularization procedure currently applied [unpublished data, 35, 36].

#### **2.2. Cells**

processes to occur by providing a cell-friendly milieu, where viable engrafting elements are able to accomplish synthetic and contractile functions with extracellular matrix continuous

Cell-freed natural matrices can be realised by means of several methods: trypsin-based enzymatic and detergent decellularization procedures are only two examples of the proposed treatments. Most decellularizing protocols take advantage of a combined mechanical and

Grauss and colleagues interestingly compared various protocols currently applied to decellularize porcine aortic heart valves and verified that the combination of trypsin and Triton X-100, an anionic detergent, could be able to provoke a loss of matrix integrity [22]. Not only the enzymatic treatment with trypsin can induce elastin defragmentation, but also the use of sodium dodecyl sulphate can end out with a similar deleterious effect [23]. Conversely, the adoption of sodium cholate- and deoxycholate-based methods allows the achievement of fully nude matrices, able to be cell-recolonized *in vitro* and/or *in vivo* even

Besides native heart valves, another natural tissue has been regarded with attention for the production of animal-derived acellular scaffolds. Pericardium has been extensively applied in bioprosthetic manufacturing thanks to its biocompatible, mechanical and biological proper‐ ties, entirely suitable for long-lasting heart valve substitutes. More often of bovine origin, this tissue partially differs from a native heart valve for its low cellularity and extremely compacted ECM, raising the question on which decellularization process can best convey the optimal outcome [29]. With regards to this tissue, the comparative analysis developed by Yang et al. revealed a superior decellularizing and preserving effect of enzymatic/detergent treatment or

Sodium cholate demonstrates instead a less aggressive behaviour towards the pericardial tissue providing analogous results to its application on native semilunar heart valves [30].

Together to the elimination of xenogeneic cells, a variable depletion in the glycosaminoglycan (GAG) content is usually observed after decellularization procedure. GAGs play a significant viscoelastic role by mitigating valve stress during flexion: their highly hydrophilic nature allows, in fact, the hydration of the *spongiosa* layer in this cycle phase. Indeed, several biological processes appear to be modulated by GAGs, therefore their loss has profound consequences in the mechanical behaviour and in the cellular functions of successively repopulated scaffolds [31]. Associated to GAG content is also the hydrated state of the tissue: a reduction in hydration following decellularization can induce collapse of collagen fibres and introduces a less suitable

While GAG and water preservation is pivotal, an opposite decellularizing effect is expected as far as DNA/RNA content. Similar to cell membrane residues, the phosphate groups of the nucleic acid backbone behave as powerful calcification triggers [34]. Aspecific endonucleases are often applied as efficient tools for the complete removal of nucleic acid debris [25-26].

trypsin alone on Triton X100/sodium-deoxycholate-based extraction.

remodelling.

250 Calcific Aortic Valve Disease

chemical tissue handling to ease cell removal.

for clinical implantation [23-28].

environment for colonizing cells [32], 33].

Cells embody the second key-component of a TEHV: it is this element that provides viability to the ECM and consequently permits its remodelling and maturation towards a functional organ. Furthermore, the use of cell elements of allogeneic or autologous source can prevent from non-self reactions towards synthetic biomaterials or decellularized natural tissues.

#### *2.2.1. Differentiated cells*

For a similar selection principle operated by those researchers preferring the more committed animal-derived matrix, endothelial cells, fibroblasts, myofibroblasts and/or smooth muscle cells isolated from vascular or valve conduits have been extensively utilized to seed nude matrices and obtain both endothelial coverage with antithrombotic activity and tissue repopulation [24, 37].

Endothelial cells (ECs) were first applied in the '90s to test the feasibility of endothelialisation on Biomer and Mitrathane thromboresistant polyurethane ureas in response to physiological shear stress [38]. Commonly harvested from cardiovascular structures, as adult saphenous or more immature umbilical veins, endothelial cells were also seeded onto bioprosthetic heart valves to increase their low thrombogenic properties [39]. In an analogous fashion, a previous endothelial coverage on harshly decellularized native tissues can avert *in vivo* thromboembolic events, related to basal membrane damage/loss resulting in collagen fibres exposure.

More frequently, ECs were employed in combination to other cell elements, as for example myo/fibroblasts. A sequential seeding of fibroblasts and ECs was demonstrated effective in the *in vitro* creation of tissue-engineered valve constructs endowed with appropriate cell topography [40]. However, the district of fibroblast cell origin is able to significantly affect the degree of recellularization with better outcome associated to the usage of arterial myofibro‐ blasts rather than of dermal fibroblasts [41, 42].

Valve fibroblasts, known as valve interstitial cells (VICs) for their cusp origin, have been successfully employed as sole repopulating cell population, exhibiting unmodified prolifera‐ tive and synthetic abilities once engrafted in decellularized scaffolds [24].

While dissimilar to VICs, smooth muscle cells of vascular derivation (vSMCs) can be chemi‐ cally manipulated with epidermal growth factor (EGF), platelet-derived growth factor (PDGF) and transforming growth factor beta-1 (TGF-beta1) for a phenotypic switch towards leaflet cells [43].

#### *2.2.2. Stem cells and progenitors*

Marrow stromal cells, umbilical cord myofibroblasts and progenitor cells, chorionic villiderived cells and placenta or amniotic fluid progenitors share not only the potential to transdifferentiate in valve phenotypes after appropriate stimulation, but are also associat‐ ed to several positive aspects, making them particularly attractive for bioengineering applications [25, 44-46].

the addition of specific chemicals can further train tissue-engineered valve constructs towards

Novel Therapeutic Strategies for the Treatment of End-Stage Valvulopathies

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Firstly developed devices were very simple systems based on common or modified petri

The introduction of a dynamic conditioning was hence operated to improve cell engraftment and differentiation through the assemblage of two chambers: one for lodging the valvular construct, the other mimicking ventricular function. Connected to an air pump, the ventriclelike compartment can exert defined hydrodynamic settings in terms of flow and pressure. Diaphragms of different materials, as for example silicone, separate the two chambers and are

The small size of the device has been primarily regarded for ensuring long-term culturing in defined temperature/gaseous environment. A compact system can, in fact, easily fit in an incubator, where CO2 saturation, humidity and temperature are already set for cell culture [16]. To this aim, the implementation with gas sensors directly placed in the device can result in a

Single specific mechanical stresses were successively applied to better induce tissue matura‐ tion and, furthermost, to activate endothelial-mesenchymal transformation. Laminar flow-, flexure- and cyclic strain-based stimulations could have profound effects on mechanical

A three-chamber bioreactor was developed by Sierad et al. in 2010 to respect previously indicated criteria and other important conditions, as easy valve mounting, physiological stimulation (transvalvular pressures, pulsatile forces, flow rate, frequency, stroke rate and shear stresses) and full control over parameters. Absence of toxic or degradable fabrication materials, maximum visibility, together to ease of sterilization and waste removal, further increase the yield of repeatable results. Compliance and reservoir tanks with sterile filters for gas exchange, one-way and resistance valves, pressure transducers, a web-cam and a ventilator

The reconstruction of a heart valve tissue in a plate dish was first endeavoured by the Bostonian teams of R. Langer, J.P. Vacanti and J.E. Jr Mayer, who published in 1995 the results of this pioneering work [12]. Vascular cells, outgrown from ovine neonatal femoral artery explants, were divided into two populations by LDL selection. LDL-negative SMCs/fibroblasts and LDL-positive ECs were sequentially seeded onto polyglycolic acid/polyglatin scaffolds in a 2 week-long procedure. Later, different cells and polymers were combined in the most efficient valve tissue combinations [41, 73, 74]. The introduction of PHAs, PH4B/PHAs and pulsatile flow conditioning in a bioreactor for heart valve construction allowed the attainment of more pliable scaffolds. After dynamic seeding, cells demonstrated to be actively involved in GAG and collagen synthesis, leading to an autologous replacement of the polymeric mesh [16, 75, 76]. PH4B/PHA became, in particular, the most promising scaffold for several successful TEHV

stiffness, collagen synthesis and alignment in the tissue-engineered valve [69-71].

their proper functionality [66].

displaced periodically by air influx [16].

pump complement this efficient system [72].

approaches relying on diverse stem cells [44, 45, 49, 77-79].

*2.2.4. In Vitro applications*

supports, in which TEHVs could be statically cultured.

superior chemical control of CO2, N2, O2, glucose and lactate [67, 68].

Stem and progenitor cells are now of relatively safe isolation from foetal and neonatal tissues, such as amniotic fluid, chorionic villi or umbilical cord. Most of them demon‐ strate mesenchymal properties, potentiated by higher cell plasticity in relation to their immature state. The embryonic-like phenotype, possessed by these early precursors, can be 'frozen' in cell banks at its genuine isolation state without further differentiation/ maturation and loss of stemness [45, 47-51].

While neonatal progenitor cells can be cryopreserved at birth in view of a future use, adult bone marrow-derived cells, and especially their mesenchymal compartment, can be easily harvested from the same cardiopathic patient for a fully autologous TEHV or even employed for the creation of allogeneic constructs [25], with no risk of cell rejection thanks to their beneficial immunomodulatory properties [52, 53]. MSCs reside virtually in all post-natal body departments, as for example adipose tissue or dermis [54, 55]. MSCs obtained from bone marrow (BM-MSCs) offer some advantages over other stem or progenitor cells in terms of their prospects for use in routine clinical practice, i.e. relatively simple protocols for their isolation, storage and *in vitro* expansion and a surprising phenotypic resemblance to valve cells [56, 57]. Indeed, their phenotypic convertibility into ECs, fibroblasts/myofibroblasts and SMCs might allow in a single step-seeding procedure to reconstruct the cell geographic distribution typical of valve cusps [25, 47, 58-60]. hBM-MSCs display a repertoire of molecules that may be relevant to their adhesion and penetration in synthetic or decellularized scaffolds, including b1-integrin (which plays a pivotal role by mediating cell–ECM interactions), CD54, CD105 and CD44 (which act cooperatively in cell homing via binding to hyaluronan, the major non-protein glycosaminoglycan of the ECM) [25].

Another cell type with attractive potential for heart valve tissue engineering is the equally rare, but more easily harvestable circulating fraction of endothelial progenitor cells (EPCs). Commonly isolated from the peripheral blood with a simple venous drawing (i.e. umbilical vein), these progenitors have a high proliferation activity and can commit to transform into a mesenchymal phenotype, reminiscent of the endothelial-mesenchyme transition during embryonic valve development [61-63].

#### *2.2.3. Bioreactors*

Scaffolds and cells do not represent alone sufficient components for a successful TEHV, but conditioning is indispensable to achieve a perfect maturation of the construct prior to implan‐ tation. This last step in manufacturing a viable valve can be guaranteed by the use of bioreactors able to submit it to physiological pressures and flow [64, 65]. Besides mechanical stimulation, the addition of specific chemicals can further train tissue-engineered valve constructs towards their proper functionality [66].

Firstly developed devices were very simple systems based on common or modified petri supports, in which TEHVs could be statically cultured.

The introduction of a dynamic conditioning was hence operated to improve cell engraftment and differentiation through the assemblage of two chambers: one for lodging the valvular construct, the other mimicking ventricular function. Connected to an air pump, the ventriclelike compartment can exert defined hydrodynamic settings in terms of flow and pressure. Diaphragms of different materials, as for example silicone, separate the two chambers and are displaced periodically by air influx [16].

The small size of the device has been primarily regarded for ensuring long-term culturing in defined temperature/gaseous environment. A compact system can, in fact, easily fit in an incubator, where CO2 saturation, humidity and temperature are already set for cell culture [16]. To this aim, the implementation with gas sensors directly placed in the device can result in a superior chemical control of CO2, N2, O2, glucose and lactate [67, 68].

Single specific mechanical stresses were successively applied to better induce tissue matura‐ tion and, furthermost, to activate endothelial-mesenchymal transformation. Laminar flow-, flexure- and cyclic strain-based stimulations could have profound effects on mechanical stiffness, collagen synthesis and alignment in the tissue-engineered valve [69-71].

A three-chamber bioreactor was developed by Sierad et al. in 2010 to respect previously indicated criteria and other important conditions, as easy valve mounting, physiological stimulation (transvalvular pressures, pulsatile forces, flow rate, frequency, stroke rate and shear stresses) and full control over parameters. Absence of toxic or degradable fabrication materials, maximum visibility, together to ease of sterilization and waste removal, further increase the yield of repeatable results. Compliance and reservoir tanks with sterile filters for gas exchange, one-way and resistance valves, pressure transducers, a web-cam and a ventilator pump complement this efficient system [72].

#### *2.2.4. In Vitro applications*

*2.2.2. Stem cells and progenitors*

252 Calcific Aortic Valve Disease

applications [25, 44-46].

maturation and loss of stemness [45, 47-51].

glycosaminoglycan of the ECM) [25].

embryonic valve development [61-63].

*2.2.3. Bioreactors*

Marrow stromal cells, umbilical cord myofibroblasts and progenitor cells, chorionic villiderived cells and placenta or amniotic fluid progenitors share not only the potential to transdifferentiate in valve phenotypes after appropriate stimulation, but are also associat‐ ed to several positive aspects, making them particularly attractive for bioengineering

Stem and progenitor cells are now of relatively safe isolation from foetal and neonatal tissues, such as amniotic fluid, chorionic villi or umbilical cord. Most of them demon‐ strate mesenchymal properties, potentiated by higher cell plasticity in relation to their immature state. The embryonic-like phenotype, possessed by these early precursors, can be 'frozen' in cell banks at its genuine isolation state without further differentiation/

While neonatal progenitor cells can be cryopreserved at birth in view of a future use, adult bone marrow-derived cells, and especially their mesenchymal compartment, can be easily harvested from the same cardiopathic patient for a fully autologous TEHV or even employed for the creation of allogeneic constructs [25], with no risk of cell rejection thanks to their beneficial immunomodulatory properties [52, 53]. MSCs reside virtually in all post-natal body departments, as for example adipose tissue or dermis [54, 55]. MSCs obtained from bone marrow (BM-MSCs) offer some advantages over other stem or progenitor cells in terms of their prospects for use in routine clinical practice, i.e. relatively simple protocols for their isolation, storage and *in vitro* expansion and a surprising phenotypic resemblance to valve cells [56, 57]. Indeed, their phenotypic convertibility into ECs, fibroblasts/myofibroblasts and SMCs might allow in a single step-seeding procedure to reconstruct the cell geographic distribution typical of valve cusps [25, 47, 58-60]. hBM-MSCs display a repertoire of molecules that may be relevant to their adhesion and penetration in synthetic or decellularized scaffolds, including b1-integrin (which plays a pivotal role by mediating cell–ECM interactions), CD54, CD105 and CD44 (which act cooperatively in cell homing via binding to hyaluronan, the major non-protein

Another cell type with attractive potential for heart valve tissue engineering is the equally rare, but more easily harvestable circulating fraction of endothelial progenitor cells (EPCs). Commonly isolated from the peripheral blood with a simple venous drawing (i.e. umbilical vein), these progenitors have a high proliferation activity and can commit to transform into a mesenchymal phenotype, reminiscent of the endothelial-mesenchyme transition during

Scaffolds and cells do not represent alone sufficient components for a successful TEHV, but conditioning is indispensable to achieve a perfect maturation of the construct prior to implan‐ tation. This last step in manufacturing a viable valve can be guaranteed by the use of bioreactors able to submit it to physiological pressures and flow [64, 65]. Besides mechanical stimulation, The reconstruction of a heart valve tissue in a plate dish was first endeavoured by the Bostonian teams of R. Langer, J.P. Vacanti and J.E. Jr Mayer, who published in 1995 the results of this pioneering work [12]. Vascular cells, outgrown from ovine neonatal femoral artery explants, were divided into two populations by LDL selection. LDL-negative SMCs/fibroblasts and LDL-positive ECs were sequentially seeded onto polyglycolic acid/polyglatin scaffolds in a 2 week-long procedure. Later, different cells and polymers were combined in the most efficient valve tissue combinations [41, 73, 74]. The introduction of PHAs, PH4B/PHAs and pulsatile flow conditioning in a bioreactor for heart valve construction allowed the attainment of more pliable scaffolds. After dynamic seeding, cells demonstrated to be actively involved in GAG and collagen synthesis, leading to an autologous replacement of the polymeric mesh [16, 75, 76]. PH4B/PHA became, in particular, the most promising scaffold for several successful TEHV approaches relying on diverse stem cells [44, 45, 49, 77-79].

Elastomeric poly(glycerol sebacate) scaffolds treated with multiple coating strategies based on ECM-derived proteins allowed adhesion and transdifferentiation of EPCs [61, 80].

had anatomical and histological resemblance to a quite mature blood vessel and it could be

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Also offering the opportunity to create tissue banks for ready-to-use devices at the moment of clinical need, the investigation on tissue-guided regenerated heart valves (TGRHVs) has

**4. Preclinical and clinical applications of tissue engineering and tissue**

The preclinical proof-of-principle of TEHVs as valve substitutes has been demonstrated in the lamb model already with the polymeric bioconctructs firstly produced *in vitro* [12]. Vascular cell-repopulated polyglycolic acid/polyglactin matrices were implanted in the pulmonary position up to 21 days. Function assessment by Doppler echocardiography demonstrated no

Each subsequent modification in scaffold or cell types, as introduced by the same group, was generally tested *in vivo*, validating progressive functional improvements in transplanted lambs

Further TEHVs applications in preclinical models were substantially based on the use of P4HB/ PHA with few exceptions, as electrospun polydioxanone [88]. In combination with stem cells of various stromal origins, P4HB/PHA-formulated engineered tissues were evaluated in a long-term animal model, showing replacement of the exogenous matrix after nearly 8 weeks

A MSC-engineered mesh of polyglycolic and polylactic acids was evaluated as autologous pulmonary valve replacement in juvenile sheep. The good performance of this *in vitro* generated construct could be appreciated in a long follow-up of 4 months with restoration of

Despite biomechanical stimulation induced optimal results in term of cell viability and differentiation almost independently from the cytotype utilised, combined polymer/cell-based efforts to obtain a valve substitute have usually failed in recreating the fibre arrangement of a native ECM. In fact, trilaminate distribution of collagens, GAGs and elastin has been reported

A finely organized ECM already exists in native heart valves and can be conserved after cell removal. After decellularization with trypsin/EDTA, heart valve conduits were seeded with ECs and myofibroblasts. Allogenically implanted in orthotopic position, they performed adequately. Ex vivo tissue analyses revealed surface endothelium reconstitution, myofibro‐ blasts-mediated repopulation and ECM synthesis with no signs of inflammation and calcifi‐

stenosis or regurgitation signs, even if a substantial leaflet thickness was reported.

hence considered an optimal vascular substitute [87].

particularly increased in recent years.

**regeneration approaches**

or sheep [41, 73, 74].

only in few cases [88].

cation [89].

a native-like pulmonary heart valve [59].

*in vivo* [16].

In respect to the application of bioabsorbable polymers, the other TEHV modality, founded on natural ECMs, was experimented some years later. After the development of various decellularizing treatments, the combination with differentiated cells, as ECs and VICs, was able to generate directly *in vitro* by static conditioning surrogates of early heart valve tissues [24, 81]. As well as polymeric TEHVs, cell-repopulated decellularized ECMs were positively remodelled after dynamic stimulation with proper mechanical signals. In this case, actually, also elastin content was demonstrated to increase [82-84].

The usage of stem cells as cell source for the engineering of plain ECMs led to even better *in vitro* outcomes. Multipotent differentiation potential of human bone marrow MSCs can represent the ideal characteristic for complete repopulation of natural valve matrices. MSC engrafting ability was evaluated on decellularized porcine and human scaffolds. In both considered interactions, stem cells were able to adhere, spread within the ECM and transdif‐ ferentiate towards typical valve phenotypes (ECs, VICs). Collagen, GAG and elastin synthesis was indeed activated in engrafted cells, which tend to distribute similarly to the original valve cell topography. It was, however, the homotypic combination to better favour MSC-to-SMC conversion in the *ventricularis* layer [25].
