**5. Preformed gels and scaffolds**

#### **5.1. Rationale**

An injectable α-cyclodextrin/poly(ethylene glycol)–b-polycaprolactone-(dodecanedioic acid) polycaprolactone–poly(ethylene glycol) (MPEG–PCL–MPEG) hydrogel was used to deliver and encapsulate bone marrow stem cells into infarcted myocardium [119]. The CD/MPEG-PCL-MPEG hydrogel alone does not induce angiogenesis, but can serve as a support in the infarcted zone and contribute to inhibit the left ventricular remodeling. One month after the injection of the gel combined with cells, cell retention and survival and the density of vessels were increased when compared with cells injection alone; moreover, the gel was absorbed,

PEG-based temperature-sensitive hydrogels have also been combined with growth factors or other molecules. VEGF was mixed or conjugated with the aliphatic polyester hydrogel poly(δvalerolactone)-block-poly(ethylene glycol)-block-poly(δ-valerolactone) (PVL-b-PEG-b-PVL); the sustained VEGF release during the degradation time of the hydrogel translated into an improvement of the myocardial and functional recovery, in dependence of the preparation method [120]. In another work, a metalloproteinase-responsive PEG-based hydrogel was synthesized to be a thymosin β4 (a pro-angiogenic and pro-survival factor) delivering scaffold. It was implanted combined with endothelial and smooth muscle cells derived from human embryonic stem cells (hESC) in rats [121]. The gel provides structural organization and when was loaded with cells and thymosin b4 enhanced more contractile performance than when the hydrogel was only loaded with the factor, because of their paracrine effect. Another PEG-based hydrogel, α-cyclodextrin/MPEG–PCL–MPEG, was tested as a delivery system for erythro‐ poietin (EPO) [122], a hormone that plays a protective role in the infarcted myocardium. Rats treated with this system showed limited cell apoptosis and increased neovasculature forma‐

PEG in the format of nanoparticles has also been studied. They can be injected intravenously, circulate in the body for long periods and bind only to desired tissues. Nanoparticles targeting the infarcted myocardium were developed based on the overexpression of angiotensin II type 1 (AT1) receptor in the infarcted heart [123]. The system was formed by a vehicle and a targeter, a ligand specific to AT1 that will make the nanoparticles bind specifically. The vehicle was 142 nm diameter PEGylated liposomes, which could carry therapeutic molecules and release them in a controlled way. This system was proved to target the infarcted heart in mice model, but

SAPs are short peptides capable of forming hydrogels at physiological pH and osmolarity [124]. When the SAPs solution is placed in contact with ions or pH is changed, the charges are partially neutralized and a hydrophobic packing takes place forming beta-sheet structures, constituting fibers that build a 3D network if the concentration is high enough. Fibers shape is different depending on the nature of the employed peptides. In the particular case of the RAD16 ionic peptides family (R: arginine, A: alanine, D: aspartate) fibers thicknesses are of

Peptides can be combined with cells to encapsulate them within the peptide network [125]. RAD16-I (AcN-RADARADARADARADA-CNH2) has proved to be a useful synthetic gel

ventricular dilation was limited and the ventricular ejection fraction improved.

tion; also infarct size was reduced and cardiac function improved.

not the healthy.

5-10 nm.

*Self Assembling Peptides (SAPs)*

290 Regenerative Medicine and Tissue Engineering

An alternative approach in the field of cardiac tissue engineering involves the use of bioma‐ terials to produce patches *ex vivo* and implant them epicardially onto the infarcted tissue, conveniently adapted to its size and shape. These patches can be pre-loaded with cells (incorporated within their pores in the case of microporous scaffolds, or encapsulated in the case of a gel conformed before implantation, as shown in figure 1 b) and growth factors or drugs, and act as a cell supply, a mechanical reinforcement to the infarct scar to avoid ven‐ tricular dilation and a drug release system simultaneously.

### **5.2. Requirements of the scaffolds**

In this strategy a key aspect is to find a material that matches the required properties. The material needs also to be cell-friendly, non-cytotoxic and promote cell attachment and proliferation, and it must also be non-immunogenic [128]. The scaffolds should provide a 3D environment to the cells with a porous structure able to guide cardiomyocytes alignment and promote maturation, also induce the development of a contractile phenotype and the electromechanical coupling of the implanted cells among them, and also with the host tissue [129, 32] and need to be easily vascularized [37].

The mechanical properties exhibited by the scaffolds should be adequate to their application in heart tissue engineering. It implies that they should ideally be compliant with contractions and exhibit non-linear elasticity, as well as be capable to adapt to the shape of the heart in all phases of the heart beat. Anisotropy to mimic the directionally-dependent electrical and mechanical properties of the native myocardium is important too [130]. Besides, the stiffness of the material employed affects to a great extent the phenotype and contractile properties of the neonatal cardiomyocytes [131, 132], and has to be carefully tuned to match physiological conditions. During heart development, the ECM on which cardiomyocytes maturation takes place, stiffen 9 times. An interesting approach to mimic it is the development of materials with time dependant mechanical properties [133]. For instance, hyaluronic acid hydrogels that stiffen with time form more contractile units when compared with cultures in hydrogels without such time-dependant stiffness.

Attending to the type of strategy, three groups can be distinguished, in terms of the nature of the matrices: biologically-derived materials, synthetic (either biodegradable or biostable) materials and decellularized tissues. With the use of biodegradable scaffolds, it is expected that the matrix will degrade as the surrounding tissue is regenerated; the degradation products should not be toxic and metabolized by the body. By using permanent scaffolds, the idea is that they will be infiltrated by the host tissue and contribute to the regeneration, but also act as a permanent mechanical restraint to limit ventricular dilation. The approach of scaffolds derived from decellularized tissue is based on the use of tissues whose cells are removed and the remaining ECM maintains the architecture and mechanical properties similar to those of the native tissue. Obtaining a scaffold matching the desired properties is a hard task, as many different properties are required; thus, materials exhibiting different properties have been mixed in more advanced strategies to obtain a composite that combines them.

#### **5.3. Related problematic**

As all the approaches described so far, this one also has some advantages, disadvantages and unsolved problematic. An important disadvantage is that the application of a patch in the heart needs a much more invasive technique than a catheter-delivered system, as it requires a surgical procedure to be implanted. As advantage, the fact that the materials are synthesized and conveniently prepared out of the body can be outlined. It implies that there is no limitation in the preparation procedure and in the use of solvents (if they are properly removed at the end of the fabrication process and do not induce cytotoxicity). Therefore, the range of chem‐ istries and techniques available to obtain scaffolds with different architectures is broadening. Besides, cells can be pre-cultured *in vitro* within them prior to implantation if desired. In addition, the mechanical properties of polymer scaffolds may be tuned to match more closely those of the heart muscle than with gelly biomaterials.

Unlike native myocardium, where the greatest distance between capillaries is around 20 microns [69], scaffolds are not vascularized *a priori*. Then, cells seeded in the scaffolds have their oxygen and nutrients supply limited to their molecular diffusion through the thickness of the scaffold. Given the fact that cardiomyocytes have great consumption rates of nutrients and oxygen, diffusion is insufficient supply for thick constructs. Consequently, to obtain a thick engineered tissue with viable cells through all its thickness, pre-vascularization or improved diffusion throughout the scaffold until it is vascularized is key for the implant to succeed. Otherwise, cell density will be concentrated in the external parts and cell viability will be compromised in the center of the scaffold if the distance to the surface is greater than a critical value estimated around 100 microns [134]. For example, the influence of oxygen concentration in cell density and viability in collagen scaffolds has been studied, the former decreasing linearly with the distance to the surface and the latter exponentially [135]. These results indicate that in order to guarantee an appropriate oxygen concentration throughout the scaffold, additional measures need to be taken.

and exhibit non-linear elasticity, as well as be capable to adapt to the shape of the heart in all phases of the heart beat. Anisotropy to mimic the directionally-dependent electrical and mechanical properties of the native myocardium is important too [130]. Besides, the stiffness of the material employed affects to a great extent the phenotype and contractile properties of the neonatal cardiomyocytes [131, 132], and has to be carefully tuned to match physiological conditions. During heart development, the ECM on which cardiomyocytes maturation takes place, stiffen 9 times. An interesting approach to mimic it is the development of materials with time dependant mechanical properties [133]. For instance, hyaluronic acid hydrogels that stiffen with time form more contractile units when compared with cultures in hydrogels

Attending to the type of strategy, three groups can be distinguished, in terms of the nature of the matrices: biologically-derived materials, synthetic (either biodegradable or biostable) materials and decellularized tissues. With the use of biodegradable scaffolds, it is expected that the matrix will degrade as the surrounding tissue is regenerated; the degradation products should not be toxic and metabolized by the body. By using permanent scaffolds, the idea is that they will be infiltrated by the host tissue and contribute to the regeneration, but also act as a permanent mechanical restraint to limit ventricular dilation. The approach of scaffolds derived from decellularized tissue is based on the use of tissues whose cells are removed and the remaining ECM maintains the architecture and mechanical properties similar to those of the native tissue. Obtaining a scaffold matching the desired properties is a hard task, as many different properties are required; thus, materials exhibiting different properties have been

As all the approaches described so far, this one also has some advantages, disadvantages and unsolved problematic. An important disadvantage is that the application of a patch in the heart needs a much more invasive technique than a catheter-delivered system, as it requires a surgical procedure to be implanted. As advantage, the fact that the materials are synthesized and conveniently prepared out of the body can be outlined. It implies that there is no limitation in the preparation procedure and in the use of solvents (if they are properly removed at the end of the fabrication process and do not induce cytotoxicity). Therefore, the range of chem‐ istries and techniques available to obtain scaffolds with different architectures is broadening. Besides, cells can be pre-cultured *in vitro* within them prior to implantation if desired. In addition, the mechanical properties of polymer scaffolds may be tuned to match more closely

Unlike native myocardium, where the greatest distance between capillaries is around 20 microns [69], scaffolds are not vascularized *a priori*. Then, cells seeded in the scaffolds have their oxygen and nutrients supply limited to their molecular diffusion through the thickness of the scaffold. Given the fact that cardiomyocytes have great consumption rates of nutrients and oxygen, diffusion is insufficient supply for thick constructs. Consequently, to obtain a thick engineered tissue with viable cells through all its thickness, pre-vascularization or improved diffusion throughout the scaffold until it is vascularized is key for the implant to

mixed in more advanced strategies to obtain a composite that combines them.

those of the heart muscle than with gelly biomaterials.

without such time-dependant stiffness.

292 Regenerative Medicine and Tissue Engineering

**5.3. Related problematic**

Many attempts have been done in this direction, like the addition of oxygen carriers to the culture medium to simulate the effect of the hemoglobin in the blood. Their addition contrib‐ uted to improve mass transport and to increase cell density [136]. Another strategy includes the use of scaffolds releasing growth factors to enhance the vascularization process, like basic fibroblast growth factor [137], vascular endothelial growth factor (VEGF) [138] and Thymosin beta-4 [139]. Another approach is the addition of the growth factor platelet derived growth factor BB to the culture medium to protect cardiomyocytes from apoptosis [140]. In a different methodology, channeled scaffolds were produced to simulate the capillary structure of the native tissues and guide endothelial cells growth. The porosity might be adjusted to increase capillary infiltration but it is limited to the maximum size of the pores on which endothelial cells can form vascular structures [141]. An alternative involves the use of decellularized tissues that already provide a native vascular network [142, 143]. The culture of endothelial cells prior to implantation of cardiac myocytes has also been explored [144], and reduced cardiomyocytes apoptosis and necrosis was found. Another possibility is to pre-implant the scaffold to pre-vascularize it prior to its implantation in the final site: alginate scaffolds loaded with angiogenic and pro-survival factors (Matrigel, SDF-1, VEGF and IGF-1) were preimplanted into the omentum of rats [145]. It proved to be a very interesting *in vivo* "bioreactor", providing to the patch a functional vascular network that maintained the viability of the transplanted cells.

Pre-culturing the scaffolds *in vitro* in bioreactors has also been a considered an option. There are many types of bioreactors (stirring, spinning flasks rotating, perfusion, etc.), but not all of them improve enough the diffusion to lead to uniform cell density and compact tissue formation. As an example, in a study where rotating bioreactors were used to culture poly‐ glycolic acid (PGA) scaffolds [146], functional and interconnected cells only were found in the peripheral parts, where there was a better diffusion of the oxygen. Perfusion bioreactors have been developed to try to reduce diffusional limitations by establishing interstitial flow through the scaffolds in order to allow the formation of thick tissues with uniform cell density through‐ out them. The effect of culturing scaffolds in perfusion bioreactors was compared with culturing them in spinner flasks [134] or orbital mixed dishes [147]. In both studies results were improved with the perfusion bioreactors; when cultured in the others, high cell density was only found in the outer layers. However, a limitation of perfusion bioreactors is the medium flow rate, because of the hydrodynamic shear the interstitial flow inflicts to the cells, which could maintain them in a rounded morphology or even wash them out if it is too high. This finding led to the combination of the perfusion culture with the use of channeled scaffolds that provided separated compartments for medium flow [148]. Even more, this strategy has been successfully combined and used simultaneously with a selective pre-seeding of the scaffold in the channels with endothelial cells using a perfusion seeding technique, which provides uniform seeding throughout the entire scaffold without the use of cell carriers [149].

Another step was made when the pulsatile perfusion bioreactor [150] was developed. It was expected that the pulsatile interstitial medium flow would provide mechanical conditioning and improved mass transport, intending that all together would lead to a tissue with better contractile properties. Indeed, scaffolds cultured under these conditions had enhanced contractile properties. A different type of bioreactor, with bidirectional slow flow perfusion obtained with an oscillatory system was tested with culture medium loaded or not with Insuline-like growth factor-I [151]. The advantage of the combined strategies was revealed.

However, despite the great efforts put and the improvements achieved, obtaining vascularized constructs is still an unsolved problem.

#### **5.4. Preparation techniques**

Many different techniques have been proposed to obtain 3D porous structures with different topographies and porosities, basically based in phase separation procedures or the use porogen templates to create the pores. Now with the introduction of controlled computer assisted systems, new possibilities are open. Next, a brief description of the main techniques employed to prepare scaffolds for heart tissue engineering is outlined.

The electrospinning technique is based in the application of a high voltage to a polymer melted or in a solution that leads to the formation of ultrathin nonwoven fibers [152], which are projected on a collector giving rise to fiber mats with controlled thicknesses. The fibers diameters can be obtained in the range of the ECM proteins. This technique also allows the preparation of aligned fibers, which can be applied to obtain aligned cardiac cells [153].

The particle leaching technique is based on the use of a porogen that is mixed with a polymer solution or a melted polymer. This porogen is removed after the solvent has been eliminated (solvent casting, freeze extraction) or the polymer has solidified after cooling, leaving empty spaces (pores) with the size and shape of the porogen template (and also small pores for the elimination of the solvent, if used). Porosity and pores interconnection can be tuned by changing the porogen-polymer ratio. Gas foaming avoids the use of solvents and high temperatures, because the pores are obtained by exposition to a high pressure gas followed by a pressure decrease with nucleation and growth of pores. The freeze-drying technique consists in freezing a polymeric solution and then lyophilize it to remove the solvent in the frozen state and obtain a solid porous structure [154]. Different morphologies can be obtained by changing the freezing conditions [155].

Microfluidic patterning consists in forcing a polymer solution through a channeled mould previously obtained with the desired geometry. Once the polymer is consistent, the mould is removed and the scaffold or patterned surface is ready. Selective laser sintering is a technique based in the use of a CO2 laser to sinterize selectively the powder of a material to form the cross section of each layer of a 3D object.

Microcontact printing is a technique that allows cell adhesion guidance [156]. It consists in the use of a stamp, with the pattern to be followed by the cells. The stamp is inked with the solution that is expected to promote the adhesion (laminin, ECM proteins, etc.) and then pressed against the substrate to transfer the solution. By loading the solution with growth factors, cell differ‐ entiation can also be induced in patterns [157].

#### **5.5. Biomaterials employed as scaffolds**

Many different types of materials have been considered for cardiac tissue engineering. According to their origin we can distinguish: biologically-derived materials, decellularized tissues and synthetic materials. Natural materials include collagen, gelatin, fibrin, silk and alginate; and synthetic materials include polyurethane (PU), polylactide acid (PLA), polygly‐ colic acid (PGA), polycaprolactone (PCL), or polyglicerolsebacate (PGS), among others.

#### *5.5.1. Natural materials*

#### *Collagen*

successfully combined and used simultaneously with a selective pre-seeding of the scaffold in the channels with endothelial cells using a perfusion seeding technique, which provides

Another step was made when the pulsatile perfusion bioreactor [150] was developed. It was expected that the pulsatile interstitial medium flow would provide mechanical conditioning and improved mass transport, intending that all together would lead to a tissue with better contractile properties. Indeed, scaffolds cultured under these conditions had enhanced contractile properties. A different type of bioreactor, with bidirectional slow flow perfusion obtained with an oscillatory system was tested with culture medium loaded or not with Insuline-like growth factor-I [151]. The advantage of the combined strategies was revealed.

However, despite the great efforts put and the improvements achieved, obtaining vascularized

Many different techniques have been proposed to obtain 3D porous structures with different topographies and porosities, basically based in phase separation procedures or the use porogen templates to create the pores. Now with the introduction of controlled computer assisted systems, new possibilities are open. Next, a brief description of the main techniques

The electrospinning technique is based in the application of a high voltage to a polymer melted or in a solution that leads to the formation of ultrathin nonwoven fibers [152], which are projected on a collector giving rise to fiber mats with controlled thicknesses. The fibers diameters can be obtained in the range of the ECM proteins. This technique also allows the preparation of aligned fibers, which can be applied to obtain aligned cardiac cells [153].

The particle leaching technique is based on the use of a porogen that is mixed with a polymer solution or a melted polymer. This porogen is removed after the solvent has been eliminated (solvent casting, freeze extraction) or the polymer has solidified after cooling, leaving empty spaces (pores) with the size and shape of the porogen template (and also small pores for the elimination of the solvent, if used). Porosity and pores interconnection can be tuned by changing the porogen-polymer ratio. Gas foaming avoids the use of solvents and high temperatures, because the pores are obtained by exposition to a high pressure gas followed by a pressure decrease with nucleation and growth of pores. The freeze-drying technique consists in freezing a polymeric solution and then lyophilize it to remove the solvent in the frozen state and obtain a solid porous structure [154]. Different morphologies can be obtained

Microfluidic patterning consists in forcing a polymer solution through a channeled mould previously obtained with the desired geometry. Once the polymer is consistent, the mould is removed and the scaffold or patterned surface is ready. Selective laser sintering is a technique

laser to sinterize selectively the powder of a material to form the cross

employed to prepare scaffolds for heart tissue engineering is outlined.

constructs is still an unsolved problem.

by changing the freezing conditions [155].

based in the use of a CO2

section of each layer of a 3D object.

**5.4. Preparation techniques**

294 Regenerative Medicine and Tissue Engineering

uniform seeding throughout the entire scaffold without the use of cell carriers [149].

There are a number of commercial collagen patches, widely used by clinicians for other purposes, which are now under study as epicardial patches, because it has been reported to be a good substrate for cell attachment and infiltration [158]. They have been combined with different cell types and molecules. Unfortunately, collagen sponges have a great swelling rate and poor mechanical performance in aqueous medium.

Collagen can be used in two formats, as a porous scaffold or as a hydrogel. To obtain the scaffold a collagen solution is lyophilized and then rehydrated and seeded with cells. In the case of hydrogels, a collagen solution is mixed with cells *ex vivo* and then gelled. As a gel entrapping embryonic chick cardiomyocytes [159], it was found to beat and arrange as a highly organized tissue-like when pulses with different frequency were applied.

The potential of collagen scaffolds as an attractant for neovascularization was demonstrated in a study with rats [160]. Collagen sponges implanted in both healthy and cryoinjured hearts were almost absorbed after 2 months, but the remaining structures were populated by new arterioles and capillaries. In another study, collagen has been combined with chondroitin 6 sulfate to obtain porous scaffolds. These scaffolds delivered MSC in the infarcted region in a rat model, promoting neovascularization [161].

The therapeutical potential of collagen as epicardial patch has been compared with injectable approaches. Collagen matrices loaded with mesenchymal stem cells (MSC) [162], and collagen scaffolds loaded with human umbilical cord blood cells (hUCBCs) [25], gave better results than the injection of cells alone in mice. In the MAGNUM phase I clinical trial [163], intrainfarct cell therapy of autologous BMC was combined with collagen scaffolds loaded with BMC. This treatment was found to be safe and contribute to limit left ventricular remodeling by increasing the thickness of the ventricle wall and then reducing the stress of the wall.

Collagen has been modified to incorporate bioactive molecules to improve its biological behavior. Its scaffolds have been modified with RGD [164] and cardiac markers of cardio‐ spheres derived from cardiac progenitor cultured on them were upregulated. Collagen functionalized with interleukin-10 plasmid [165] (an anti-inflammatory plasmid) increased 5 times cell retention and modulated inflammation.

#### *Gelatin*

Gelatin is obtained from chemical denatured collagen; it is therefore weaker and degrades faster than it [27]. It has been reported to provoke unspecific inflammatory response upon degrading; at first this can be considered an undesired effect, but for certain applications it might be beneficial for the positive impact that can have on angiogenesis [166]. A commercial gelatin sponge bare or cultured either with fetal or adult rat heart cells was implanted to replace the resected right ventricular outflow tract (ROVT) of rats [167]. After 4 weeks a great inflam‐ mation was observed and after 12 weeks the patches had endothelial cells on the endocardial surface. Nonetheless, the authors concluded that a material inducing less inflammatory response is needed.

#### *Fibrin*

Fibrin can be used as an injectable gel, but can also be preformed *ex vivo*, which broad‐ ens the possibilities of fabrication. For example, SDF-1 (a factor that is up-regulated for a period of time after a myocardial infarction, and contributes to mobilize cells from bone marrow and peripherial blood to the damaged tissue) was covalently bound to a PEGylat‐ ed fibrin patch [168] and implanted in an AMI mouse model; the SDF-1 loaded patch reduced more significantly the scar area expansion and improved the left ventricular function than the un-loaded patch.

#### *Alginate*

Alginate scaffolds obtained by the freeze drying technique have been extensively ex‐ plored in myocardial regeneration. Loaded with fetal cardiac cells and implanted in infarcted rats, they limited left ventricular dilation [169]. However, cultured with neona‐ tal or fetal cardiomyocytes in static conditions, cell aggregates were formed due to the nonadhesive nature of the alginate [170].

To improve cell adhesion and survival modifications of alginate scaffolds have been investi‐ gated. For example, it has been modified to incorporate the adhesion peptide RGD [171], which improved cell adhesion, reduced apoptosis, accelerated tissue regeneration and led to the organization of cardiomyocytes in myofibers *in vitro*, and also with a combination of RGD and the heparin-binding peptide G4SPPRRARVTY (HBP) [172], with better results.

#### *Polysaccharides*

Polysaccharide-based scaffolds have also been investigated with myocardial regeneration purposes. The effectiveness of freeze-dried pullulan and dextran patches was compared to mesenchymal stem cells endocardial delivery alone in a rat myocardial infarction model [173], the scaffolds improving the cell engraftment and survival at 1 and 2 months.

Because of silk fibroin good mechanical properties, biological performance, and its easy processing to obtain different morphologies, it has generated interest in the tissue engineering field. Silk is produced by some insects like spiders or silkworms, and is considered a nondegradable material by the FDA [174]. Silk fibroin has been combined by chitosan and hyaluronic acid to produce microparticles that were pressed and crosslinked with genipin to obtain cardiac patches [175]. MSC cultured on the composite patches exhibited greater proliferation and cardiomyogenic differentiation than in silk patches.

Recently, non-mulberry silk fibroin from Antheraea mylitta has been investigated as a material for cardiac tissue engineering [176]. It has better mechanical properties than mulberry silk, contains RGD sequences, is non-cytotoxic and induces low level of inflammatory response. When neonatal rat cardiomyocytes were seeded in an Antherea mylitta silk lyophilized scaffold, the results were better than those obtained with a mulberry silk.

#### *Decellularized-tissue derived scaffolds*

Decellularized extracellular matrices have been used as scaffolds in many studies and also in preclinical and human clinical applications [177]. The decellularization process consists in a set of washes to remove the cells but maintain as much as possible the architecture, proteins and adhesion molecules. The more aggressive the washes and treatments are, the lower the risk of allogenic immune reaction is, but undesired washout of adhesion proteins and archi‐ tecture damage can be associated [65].

Decellularized sheets have been tested in combination with fibrin, TGF-beta, and MSC and tested in a nude rat model of infarction with positive results [178]. A patch of urinary bladderderived extracellular matrix (UBM) was implanted in pigs, as a left ventricular wall replace‐ ment after infarction, and compared with a polytetrafluoroethlyene (ePTFE) [177] one. At three months, the results were better with the UBM: it was reabsorbed and a cellularized and vascularized tissue rich in collagen was formed.

Sliced decellularized porous scaffolds of acellular bovine pericardia have been combined with cell sheets from bone marrow stem cells, cultured and implanted in rats replacing the resected infarcted myocardium [179]. The patch pores were filled by cells, new vessels and new muscle fibers, indicating that the graft was integrating. Cardiac function was improved and the dilated left ventricle was restored after implantation. In a revolutionary study entire rat hearts were decellularized, and then re-cellularized with neonatal cardiac cells [180]. The architecture was conserved and the preserved vasculature was perfusable. Seeded cardiomyocytes coupled electromechanically and after 8 days under external electrodes stimulation the re-cellularized heart beat and was capable to pump blood.

#### *5.5.2. Synthetic materials*

Synthetic materials are prepared in the laboratory, allowing precise control over their mechanical properties, degradation, morphology and porosity that can be tuned as desired

#### *Silk*

Collagen has been modified to incorporate bioactive molecules to improve its biological behavior. Its scaffolds have been modified with RGD [164] and cardiac markers of cardio‐ spheres derived from cardiac progenitor cultured on them were upregulated. Collagen functionalized with interleukin-10 plasmid [165] (an anti-inflammatory plasmid) increased 5

Gelatin is obtained from chemical denatured collagen; it is therefore weaker and degrades faster than it [27]. It has been reported to provoke unspecific inflammatory response upon degrading; at first this can be considered an undesired effect, but for certain applications it might be beneficial for the positive impact that can have on angiogenesis [166]. A commercial gelatin sponge bare or cultured either with fetal or adult rat heart cells was implanted to replace the resected right ventricular outflow tract (ROVT) of rats [167]. After 4 weeks a great inflam‐ mation was observed and after 12 weeks the patches had endothelial cells on the endocardial surface. Nonetheless, the authors concluded that a material inducing less inflammatory

Fibrin can be used as an injectable gel, but can also be preformed *ex vivo*, which broad‐ ens the possibilities of fabrication. For example, SDF-1 (a factor that is up-regulated for a period of time after a myocardial infarction, and contributes to mobilize cells from bone marrow and peripherial blood to the damaged tissue) was covalently bound to a PEGylat‐ ed fibrin patch [168] and implanted in an AMI mouse model; the SDF-1 loaded patch reduced more significantly the scar area expansion and improved the left ventricular

Alginate scaffolds obtained by the freeze drying technique have been extensively ex‐ plored in myocardial regeneration. Loaded with fetal cardiac cells and implanted in infarcted rats, they limited left ventricular dilation [169]. However, cultured with neona‐ tal or fetal cardiomyocytes in static conditions, cell aggregates were formed due to the non-

To improve cell adhesion and survival modifications of alginate scaffolds have been investi‐ gated. For example, it has been modified to incorporate the adhesion peptide RGD [171], which improved cell adhesion, reduced apoptosis, accelerated tissue regeneration and led to the organization of cardiomyocytes in myofibers *in vitro*, and also with a combination of RGD and

Polysaccharide-based scaffolds have also been investigated with myocardial regeneration purposes. The effectiveness of freeze-dried pullulan and dextran patches was compared to mesenchymal stem cells endocardial delivery alone in a rat myocardial infarction model [173],

the heparin-binding peptide G4SPPRRARVTY (HBP) [172], with better results.

the scaffolds improving the cell engraftment and survival at 1 and 2 months.

times cell retention and modulated inflammation.

296 Regenerative Medicine and Tissue Engineering

*Gelatin*

*Fibrin*

*Alginate*

*Polysaccharides*

response is needed.

function than the un-loaded patch.

adhesive nature of the alginate [170].

[181]. However, they may not have as good biological performance as biologically derived materials [4].

#### *Polylactic Acid and Polyglycolic Acid (PLA and PGA)*

Polylactic acid is a biocompatible, biodegradable and FDA-approved polymer; it de‐ grades into lactic acid (non-cytotoxic), and has been widely used in patients, for example as sutures. However, its degradation products can induce a slight, undesired, acidifica‐ tion of the microenvironment [65]. Polyglycolic acid is a thermoplastic too; it has also been used in the clinic and degrades into non-toxic products. However, neither PLLA nor PGA exhibit the desired elasticity to match that of native heart tissue. In many studies PLA and PGA have been combined as poly(lactic–co-glycolic acid) (PLGA), or other polyesters, to modify their properties as desired. Electrospun PLGA fibrous membranes with different compositions (having different hydrophobicity and degradation rates) [4] were found to align cardiomyocytes in the direction of the nanofibers, the best results being those of the slightly hydrophobic copolymers. Porous beads of PLGA seeded with human amniotic fluid stem cells (hAFSCs) have been tested as a cell delivery vehicle or "cellularized microscaffold" [182]; after implantation by intramyocardial injection in a rat infarct model, they showed good retention of the cells in the site of interest. PLGA has been treated with laminin [183] to improve its biological development and combined with carbon nanofib‐ ers (CNF) to increase its conductivity and cytocompatibility [184]. PLLA-PLGA scaffolds loaded with Matrigel have been co-cultured with endothelial cells, cardiomyocytes and embryonic fibroblasts simultaneously [185], for EC to provide vasculature and act synergi‐ cally with cardiomyocytes to improve cell survival and proliferation.

#### *Poly (epsilon-caprolactone) (PCL)*

Poly(epsilon-caprolactone) is a FDA-approved biocompatible polyester, as PLA and PGA. It is more elastic because of its lower glass transition temperature, and behaves as a rubber at body temperature. Its degradation does not produce acidification because it occurs more slowly [158]. It has been proposed for myocardial regeneration for example in 3D constructs obtained by overlapping electrospun PCL nanofibrous mats (up to 5 layers) on which neonatal cardiomyocytes were cultured [186]. The layers established morphologic and electrical connections between them and exhibited synchronized beating, and no ischemia was found in the center of the constructs.

It is usually combined with PLA, PGA or its copolymer. Poly-glycolide-co-caprolactone (PGCL) biodegradable porous scaffolds have been studied as cell vehicles for bone marrowderived mononuclear cells (BMMNC) in rat myocardial infarction models [187]. BMNC migrated from the scaffold and neovasculature over the implant was detected; left ventricular function improvement and limitation of the progression of the left ventricular dilation was also observed. Scaffolds made of poly(DL-lactide-*co*-caprolactone) (PLACL), PLGA, and type I collagen [158], cultured with neonatal rat heart cells, have been compared. The composite scaffolds gave better results than controls (collagen and PLGA sponges) in terms of cellularity, contractility and cardiac markers expression (Tn-I and Cx-43). Perfusion culture improved cell density distribution.

#### *Polyurethanes (PU)*

[181]. However, they may not have as good biological performance as biologically derived

Polylactic acid is a biocompatible, biodegradable and FDA-approved polymer; it de‐ grades into lactic acid (non-cytotoxic), and has been widely used in patients, for example as sutures. However, its degradation products can induce a slight, undesired, acidifica‐ tion of the microenvironment [65]. Polyglycolic acid is a thermoplastic too; it has also been used in the clinic and degrades into non-toxic products. However, neither PLLA nor PGA exhibit the desired elasticity to match that of native heart tissue. In many studies PLA and PGA have been combined as poly(lactic–co-glycolic acid) (PLGA), or other polyesters, to modify their properties as desired. Electrospun PLGA fibrous membranes with different compositions (having different hydrophobicity and degradation rates) [4] were found to align cardiomyocytes in the direction of the nanofibers, the best results being those of the slightly hydrophobic copolymers. Porous beads of PLGA seeded with human amniotic fluid stem cells (hAFSCs) have been tested as a cell delivery vehicle or "cellularized microscaffold" [182]; after implantation by intramyocardial injection in a rat infarct model, they showed good retention of the cells in the site of interest. PLGA has been treated with laminin [183] to improve its biological development and combined with carbon nanofib‐ ers (CNF) to increase its conductivity and cytocompatibility [184]. PLLA-PLGA scaffolds loaded with Matrigel have been co-cultured with endothelial cells, cardiomyocytes and embryonic fibroblasts simultaneously [185], for EC to provide vasculature and act synergi‐

cally with cardiomyocytes to improve cell survival and proliferation.

Poly(epsilon-caprolactone) is a FDA-approved biocompatible polyester, as PLA and PGA. It is more elastic because of its lower glass transition temperature, and behaves as a rubber at body temperature. Its degradation does not produce acidification because it occurs more slowly [158]. It has been proposed for myocardial regeneration for example in 3D constructs obtained by overlapping electrospun PCL nanofibrous mats (up to 5 layers) on which neonatal cardiomyocytes were cultured [186]. The layers established morphologic and electrical connections between them and exhibited synchronized beating, and no ischemia was found

It is usually combined with PLA, PGA or its copolymer. Poly-glycolide-co-caprolactone (PGCL) biodegradable porous scaffolds have been studied as cell vehicles for bone marrowderived mononuclear cells (BMMNC) in rat myocardial infarction models [187]. BMNC migrated from the scaffold and neovasculature over the implant was detected; left ventricular function improvement and limitation of the progression of the left ventricular dilation was also observed. Scaffolds made of poly(DL-lactide-*co*-caprolactone) (PLACL), PLGA, and type I collagen [158], cultured with neonatal rat heart cells, have been compared. The composite scaffolds gave better results than controls (collagen and PLGA sponges) in terms of cellularity, contractility and cardiac markers expression (Tn-I and Cx-43). Perfusion culture improved cell

*Poly (epsilon-caprolactone) (PCL)*

in the center of the constructs.

density distribution.

materials [4].

298 Regenerative Medicine and Tissue Engineering

*Polylactic Acid and Polyglycolic Acid (PLA and PGA)*

Polyurethanes are synthetic biocompatible materials widely used in the biomedical field. Their mechanical properties and biodegradability can be tuned by changing their composition. PU degrades *in vivo* through hydrolytic chain scission, which is accelerated by the enzymes action and loads, among other factors [188], but with the appropriate composition non-biodegradable polyurethanes can be obtained [189]. This family of polymers can be used to obtain fibrous scaffolds by electrospinning with different mechanical properties depending on the fibers orientation [190] or porous elastic scaffolds [191]. Polyester urethane urea (PEUU) elastic porous scaffolds have been implanted in sub-acute infarctions in rats and were found to promote the formation of smooth muscle bundles, to increase the ventricle thickness and to improve contractile function [192]. Cell attachment on polyurethane-based porous scaffolds can be improved by pre-treating them with laminin [193].

#### *Poly(glycerol sebacate) (PGS)*

Poly(glycerol sebacate) is a biocompatible and biodegradable elastomer capable of recovering from deformation. It can be obtained by polycondensation of glycerol and sebacic acid. By changing the synthesis temperature, the properties of the resulting material can be tuned to match the desired mechanical properties. The degradation rates can also be adjusted from fast degradation to nearly inert [194].

By the use of excimer laser microablation, 3D porous PGS scaffolds with anisotropic structural and mechanical properties were obtained [195, 130]. These scaffolds induced neonatal cardiac cells alignment in the absence of external stimuli and matched the mechanical properties of adult rat right ventricle. Moreover, they allowed cell contractili‐ ty when stimulated. For its interesting mechanical properties, PGS has been coaxially electrospun with gelatin to form a nanofibrous mat with PGS in the core and gelatin in the shell [196] to enhance cell adhesion and proliferation. PGS has been modified to incorporate acrylic groups in different number (to modify its mechanical properties and degradation) and electrospun in combination with gelatin [197].

#### *Acrylate based materials*

Acrylate based materials have not been widely exploited for cardiac tissue engineering yet but the interest on them is increasing, for their versatility of processing and variety of properties obtained. For example, scaffolds made of poly(2-hydroxyethyl methacrylate-co-methacrylic acid) (P(HEMA-co-MAA) hydrogel have been obtained by fibers and microspheres templating to obtain spherical pores and parallel channels [198], which allow simultaneously mass transfer and guidance of the cardiomyocyte bundles. Mechanical properties were adjusted intention‐ ally for the elastic modulus to be lower than that of native myocardium in order to make possible the mechanical stimulation of the cells when implanted *in vivo*. In [199], poly(ethyl acrylate) (PEA) scaffolds are filled with HA gel; the scaffolds provide the three-dimensional environment and mechanical properties and the gel may act as an encapsulating medium for the cells and may be also used as a medium for drug or growth factors release. RAD16-I gel may also be used as a filler in PEA scaffolds, where it acts as a diffusion medium and improves cell seeding efficiency (figure 2).

**Figure 2.** (a) Scanning electron microscopy (SEM) image of poly(ethyl acrylate), PEA, elastomeric membranes with in‐ terconnected spherical pores. (b) CryoSEM image (cross section) of a PEA scaffold whose pores are filled with the selfassembling peptide (SAP) gel RAD16-I. (c) Adipose stem cells (nuclei stained in blue and actin cytoskeleton stained in green) seeded in a PEA scaffold with a SAP gel filling. Confocal laser scanning microscopy image of a 50 μm thick in‐ ternal slice.

#### **5.6. Electrical and mechanical stimulation**

#### *Electrical stimulation*

External electrical fields have been shown to contribute to the differentiation towards cardio‐ myocytes of different cell types, such as embryonic stem cells (ESC) [200] or BMSC [201] seeded in collagen scaffolds, and to the development of conductive and contractile properties of neonatal cardiac cells, in this case seeded with Matrigel in a collagen porous scaffold [202]. It has been proposed that the intracellular endogenous reactive oxygen species (ROS) produced when an electric field is applied contribute to the hESC differentiation [203].

In an attempt of optimizing the electrical stimulation parameters [204], it has been determined that the electrode material is very important, and best results have been obtained for carbon electrodes. Amplitude and frequency of the stimulation have also a great influence in the cultured cardiac tissue. Micropatterned electrodes can be of interest as they allow spatial control of the electric field [205].

Polymeric scaffolds limit cardiomyocytes electric communication, what restricts the synchro‐ nous beating of the engineered tissue. To improve it, gold nanowires were incorporated to a porous alginate scaffold [206]. Another approach to obtain elastic and electrical conductive scaffolds consisted in impregnating thiol-HEMA/HEMA scaffolds with gold nanoparticles [207]. In both cases even without electrical stimulation the improvement in the scaffold conductivity had positive physiological effects.

#### *Mechanical stimulation*

Mechanical stress has a great impact on cell proliferation, ECM formation and hypertrophy (increased cell size), and has been intensively studied in the field of cardiac tissue engineering. Embryonic chick and neonatal rat cardiac myocytes mixed with collagen and mechanically stimulated exhibited hypertrophy and improvement of contractile function [208]. Cardiac myocytes from neonatal rats mixed with collagen I and Matrigel and casted in rings subjected to mechanical stretch [209] showed histological characteristics of adult cardiac tissue. Action potential measurements indicated electrophysiological behavior akin to cardiac tissue. Constructs produced by simultaneously electrospinning PU and electrospraying mesenchy‐ mal stem cells [210] were cultured in spinner flasks with stretching, which led to cells align‐ ment, cardiac markers increase and ion channels development. Similarly, cells isolated from neonatal rat hearts seeded in chitosan-collagen I channeled porous scaffolds [211] and cultured under high mechanical stimulation induced cell alignment, elongation and the presence of gap junctions connecting the cells. Mechanical stress applied to human cardiac cells cultured in a gelatin scaffold improved cell distribution and proliferation within the scaffold, increased the production of the ECM, and the structure and organization was similar to normal myocardium, likely because the stretching of the scaffold favors nutrients and oxygen exchange improving cell microenvironment [212].
