**3. Innovative tissue engineering approaches**

#### **3.1. Decellularized matrices**

The Extracellular Matrix (ECM) represents the three-dimensional fibrilar protein scaffold, produced by cells of each tissue and organ, which surrounds and anchors them. It is kept in a state of dynamic reciprocity with those cells, in response to changes in the microenvironment. ECM has been shown to provide cues that affect cell migration, proliferation, gene expression and differentiation [32, 37].

The ECM is obviously the optimal support for tissue engineering, as it provides the perfect chemical composition, surface topology and physical properties experienced by cells *in vivo* *in their niche* [32]. Even though sometimes that`s exactly what is needed to be avoided (e.g. Central Nervous System ECM has been shown to contain molecules which inhibit axonal growth and hinders tissue regeneration [33,34]), ECM has been considered a great option for tissue engineering.

usually large, though, and cells within the scaffold are not able to fully fill it and achieve cell density similar to natural tissues. Therefore, it is almost as if cells were still in twodimensional surfaces [31]. Actually, extracellular matrix molecules can be washed out from 3D porous scaffolds in the same way as in 2D cultures, and may not provide means

**4.** Recent reports on the effect of matrix rigidity on (stem) cell differentiation can undermine the value of solid rigid biodegradable scaffolds at least for certain tissue applications

Stem cell differentiation has traditionally employed cocktails of various growth factors, but recently, mechanobiological concepts have been described as important to cell fate decision. The mechanism underlying cellular response to tension comprises the force generated by myosin bundles sliding along actin filaments and transmission to the ECM. Transduction of these signals link the extracellular and intracellular worlds, ultimately affected by proteins such as Rho GTPases, which not only regulate contraction of stress fibers, but also regulate

Actually, matrix rigidity has been involved in embryonic development, as well as adult stem cell differentiation. As expected, rigid surfaces facilitates adult stem cell differentiation into bone, and soft surfaces lead to differentiation of adult stem cells into soft tissues, such as fat

Even though it makes sense that biomaterials should be absorbed by the body in order to give space to neotissue formation, the same is not true when whole tissue engineering is planned. There is no use in spending efforts in order to build a construct, which will be invaded by inflammatory cells and vessels and disorganized, previously to being substituted by neotissue. Therefore, even though current tissue engineering techniques are fairly successful in treating bone, skin and cartilage loss, they are extremely limited in treating large tissue loss, as well as in regenerating complex tissues, such as heart annd kidneys, among sev‐

The Extracellular Matrix (ECM) represents the three-dimensional fibrilar protein scaffold, produced by cells of each tissue and organ, which surrounds and anchors them. It is kept in a state of dynamic reciprocity with those cells, in response to changes in the microenvironment. ECM has been shown to provide cues that affect cell migration, proliferation, gene expression

The ECM is obviously the optimal support for tissue engineering, as it provides the perfect chemical composition, surface topology and physical properties experienced by cells *in vivo*

gene expression by acting over their effector target proteins, [45].

for real tridimensional tissue formation.

302 Advances in Biomaterials Science and Biomedical Applications

or central nervous system (brain) [45].

**5.** Biodegradability of constructs

eral other tissues and organs.

**3.1. Decellularized matrices**

and differentiation [32, 37].

**3. Innovative tissue engineering approaches**

[35-37]

Recently, it has been shown that cell sensibility towards ECM chemical composition is higher than previously expected. For instance, Tsai et al. showed that MG63, an osteoblast like cell lineage, behaves differently when grown in collagen or gelatin electrospun matrices. When grown in electrospun collagen, MG63 did not show variation on cell attachment or proliferation rates. On the other hand, cells seeded on electrospun collagen showed increased expression of osteogenic genes such as Osteopontin and alkaline phosphatase. Collagen and gelatin present high chemical composition similarity, varying mainly in secondary and tertiary structure. Such fact underscores the strikingly cell sensitivity to all aspects of ECM chemical and physical composition [62]. It also underscores the potential of decellularized matrices on tissue engineering.

Decellularized tissues have been used in regenerative medicine approaches since the early eighties [38], specially focused on treating cardiovascular diseases by engineering vascular grafts. Most of the grafts produced, derived from synthetic and natural sources suffered from several limitations. When the issue of natural graft calcification and immunological recognition were related to residual cellular components of unmodified biological materials, decellulari‐ zation techniques began to be developed [38,39].

Initially, decellularization was considered for tissue grafts. Developed techniques are con‐ tinuously evolving, as every cell removal agent and method currently available alters ECM composition and cause some degree of ultrastructure disruption. Decellularization agents include chemical, biological and physical agents, each of them with different mechanisms of action.

More specifically: acids and bases promote hydrolytic degradation of biomolecules; hypotonic solutions lyses cells through osmosis with minimal changes in matrix molecules and tissue architecture; hypertonic solutions dissociates DNA from proteins; ionic, non-ionic, and zwitterionic detergents solubilize cell membranes leading to effective removal of cellular material from tissue; solvents, such as alcohol and acetone, promote either cell lysis by dehydration or solubilization and removal of lipids and biological agents, such as enzymes, and chelating agents act through protein cleavage and disrupting cell adhesion to ECM. Finally, physical agents promote cell lysis through freezing and thawing cycles, electropora‐ tion or pressure [32].

The most effective agents for decellularization of each tissue and organ will depend upon many factors, including the tissue's cellularity, density, lipid content, and thickness [32].

Lately, whole organ decellularization began to be performed, offering an interesting option for modular organs such as the heart, lung and kidneys. In 2008, Ott et al. not only performed whole heart decellularization, but also recellularized the organ with neonatal cardiomyocytes and obtained organ function [17]. This groundbreaking work highlighted the possibilities of decellularized matrix-based whole organ tissue engineering.

The major breakthrough of organ decellularization is to obtain scaffolds with perfect (or very similar) chemical composition and tridimensional structure, compared to natural organs. In addition, the vascular bed is completely preserved, facilitating *in vitro* maintenance of the construct via perfusion bioreactors, as well as *in vivo* viability of the construct, which may be reconnected to the circulatory system of host, also shown by Ott et al [17]. As decellularization is performed making use of vascular system of organs, virtually any vascularized organ may be decellularized, disregarding its size, as depicted in Figure 4. Acellular organs, such as tracheas, may also be decellularized through different protocols [15].

**Figure 4.** Decellularized pig heart. As published by Ott et al., perfusion decellularization is feasible in small rat organs, but also in bigger size organs, such as the pig heart. Illustration owned by Miromatrix Medical Inc., available at http:// miromatrix.com/technology/perfusion-decellularization-recellularization/, accessed on September 2012.

#### **3.2. Biomimetic scaffolds**

As already mentioned, an optimal scaffold attempts to mimic the function of the natural extracellular matrix [52]. The functionality of most tissues is related to their complex architecture, therefore mimicking and recapitulating this complexity *in vitro* is paramount for successful tissue engineering. In the *in vivo* condition, cells are surrounded by other cells and by the extracellular matrix (ECM), whose components, such as collagen, elastin, and laminin, are organized in nanostructures (i.e., fibers, triple helixes, etc) with specific bioactive motifs that regulate cell homeostasis [42].

Following the initial concept of mimicking the ECM chemical composition, it has been shown that the structure of the cell-surrounding niche is also paramount for optimal cell function. In accordance, modulating the scaffold microarchitecture is one of the most potent ways of achieving biomimetic tissues. Advances in microfabrication technologies have been exploited by an increasing number of research groups. Often technologies from other engineering disciplines have been translated and used in creating microfeatures in engineered scaffolds in a controlled manner. These include photolithographic approach of the electrical engineering, electrospinning tools of the textile industry, emulsification and fluid dynamics principles of chemical engineering and rapid prototyping methods of mechanical engineering. The latter will be covered in further detail in the next section [40].

The major breakthrough of organ decellularization is to obtain scaffolds with perfect (or very similar) chemical composition and tridimensional structure, compared to natural organs. In addition, the vascular bed is completely preserved, facilitating *in vitro* maintenance of the construct via perfusion bioreactors, as well as *in vivo* viability of the construct, which may be reconnected to the circulatory system of host, also shown by Ott et al [17]. As decellularization is performed making use of vascular system of organs, virtually any vascularized organ may be decellularized, disregarding its size, as depicted in Figure 4. Acellular organs, such as

**Figure 4.** Decellularized pig heart. As published by Ott et al., perfusion decellularization is feasible in small rat organs, but also in bigger size organs, such as the pig heart. Illustration owned by Miromatrix Medical Inc., available at http://

As already mentioned, an optimal scaffold attempts to mimic the function of the natural extracellular matrix [52]. The functionality of most tissues is related to their complex architecture, therefore mimicking and recapitulating this complexity *in vitro* is paramount for successful tissue engineering. In the *in vivo* condition, cells are surrounded by other cells and by the extracellular matrix (ECM), whose components, such as collagen, elastin, and laminin, are organized in nanostructures (i.e., fibers, triple helixes, etc) with specific bioactive motifs

Following the initial concept of mimicking the ECM chemical composition, it has been shown that the structure of the cell-surrounding niche is also paramount for optimal cell function. In accordance, modulating the scaffold microarchitecture is one of the most potent ways of

miromatrix.com/technology/perfusion-decellularization-recellularization/, accessed on September 2012.

**3.2. Biomimetic scaffolds**

that regulate cell homeostasis [42].

tracheas, may also be decellularized through different protocols [15].

304 Advances in Biomaterials Science and Biomedical Applications

Vacanti and coworkers pioneered the concept of engineering a vasculature using photolitho‐ graphic techniques (which use light, e.g. UV, to selectively remove parts of a thin film or the bulk of a substrate), literally generating channels within biomaterials [43].

Electrospining techniques have also been considered and tested for tissue engineering application, due to their potential of producing polymer fibers with nano to micrometer diameter scale that are physically and topographically comparable to the collagen fibers, commonly found in the natural ECM, as shown in Figure 5. It has been extensively employed in tissue engineering strategies, including vascularization strategies [41].

**Figure 5.** Similarity between ultrastructure of natural collagen fibers (A) and electrospun biomaterials (B). Source: A – author`s unpublished data; B - http://en.wikipedia.org/wiki/Electrospinning, accessed on September 2012.

Electrospun biomaterials must also be carefully fabricated, as even fiber diameter variation results in different cell behavior, as observed in endothelial cells cultured on electrospun poly(l-lactide-co-ε-caprolactone) with different fiber diameters. In contrast to cells cultured on fibers of 0.3 or 1.2μm, cells cultured on 7μm presented lower cell adhesion, spreading and proliferation [63].

Even though electrospining has presented promising results, it also suffers from some limitations, such as poor cell invasion, as usually the electrospun biomaterials are highly compacted, impeding cell migration towards the inner side of the scaffold. The search for different solvents and electrospining conditions may solve this issue, promoting less fiber compactation.

Still, as cited, many excellent constructs have been built using those techniques, and in association with the other tissue engineering strategies presented in the present chapter, strongly contribute to novel advances in the field.

Another exciting tissue engineering strategy, which have gained growing interest since the nineties, has been the hydrogel approach. Hydrogels are 3D cross-linked insoluble, hydrophilic networks of polymers that partially resemble the physical characteristics of native ECM. The biocompatibility of various hydrogels (e.g., collagen, agarose and polyethylene glycol) is well characterized, and the possibility of optimizing their physico-chemical and mechanical properties to levels that are desirable for tissue scaffolds, in order to achieve cell encapsulation, immobilization, and drug delivery turn hydrogels into an extremely promising technique [64].

Hydrogels have been successfully used in mimicking ECM of simple tissues, composed of one cell type. In many cases, hydrogels provide means for nutrient diffusion, facilitating cell maintenance *in vitro* and *in vivo*. Still, maintaining the viability of high cell density constructs remains a challenge as well as promoting cell organization within the scaffold. In many cases, construct implantation near host rich vascularization sites may be an effective strategy to promote construct viability *in vivo*.

In the cell`s point of view, hydrogels possess the advantage of completely surround encapsulated cells, and providing tridimensional substrate for cell interaction. This strategy prevents cell polarization which is common to regular scaffolds. Usually, as already stated, even though most scaffolds may be tridimensional macroscopically, they are commonly seen as bidimensional surfaces by cells.

As expected, truly tridimensional environments promote several effects over cultured cells. Some of them include, but are not limited to: cell morphology/spreading [65], cell motility and proliferation [66], and metabolic rate [67]. Obviously, all of those cell behaviors reflect differential gene expression.

Even though cell gene expression is a paramount factor to be evaluated in tissue engineering, it must be clearly noted that an exact gene expression profile is not essential for tissue engineering effectiveness. Cells will always interact with their microenvironment. What is important for tissue engineering is to maintain cell plasticity in a manner that, once in vivo, implanted cells start to behave as host cells would.

Finally, a third highly modern and innovative approach for tissue engineering, which shares the truly tridimensional environment for cells as provided by hydrogels, but which doesn`t suffer from vascularization and cell organization limitations presented by the latter, is the organ printing tissue engineering technique, which will be more thoroughly described in the next section.

#### **3.3. Organ printing**

As listed before, the main limitations of the solid scaffold approach include the low level of precision in cell placement, especially when engineering multicellular constructs, considering the intrinsic problem of vascularization of thick tissue constructs [53]. Ideally, a possible way to solve many of those aforementioned problems would be to assemble cells and ECM elements at the same time and in an organized way, in order to obtain the most similar structure found in a functional organ as possible. Over the years, technology evolution turned tissue engineers unreachable dream into a feasible objective to be fulfilled in the next years or decades. This strategy is known as organ printing and/or robotic biofabrication, and offer interesting alternatives to solid scaffold-based tissue engineering.

Still, as cited, many excellent constructs have been built using those techniques, and in association with the other tissue engineering strategies presented in the present chapter,

Another exciting tissue engineering strategy, which have gained growing interest since the nineties, has been the hydrogel approach. Hydrogels are 3D cross-linked insoluble, hydrophilic networks of polymers that partially resemble the physical characteristics of native ECM. The biocompatibility of various hydrogels (e.g., collagen, agarose and polyethylene glycol) is well characterized, and the possibility of optimizing their physico-chemical and mechanical properties to levels that are desirable for tissue scaffolds, in order to achieve cell encapsulation, immobilization, and drug delivery turn hydrogels into an extremely promising

Hydrogels have been successfully used in mimicking ECM of simple tissues, composed of one cell type. In many cases, hydrogels provide means for nutrient diffusion, facilitating cell maintenance *in vitro* and *in vivo*. Still, maintaining the viability of high cell density constructs remains a challenge as well as promoting cell organization within the scaffold. In many cases, construct implantation near host rich vascularization sites may be an effective strategy to

In the cell`s point of view, hydrogels possess the advantage of completely surround encapsulated cells, and providing tridimensional substrate for cell interaction. This strategy prevents cell polarization which is common to regular scaffolds. Usually, as already stated, even though most scaffolds may be tridimensional macroscopically, they are commonly seen

As expected, truly tridimensional environments promote several effects over cultured cells. Some of them include, but are not limited to: cell morphology/spreading [65], cell motility and proliferation [66], and metabolic rate [67]. Obviously, all of those cell behaviors reflect

Even though cell gene expression is a paramount factor to be evaluated in tissue engineering, it must be clearly noted that an exact gene expression profile is not essential for tissue engineering effectiveness. Cells will always interact with their microenvironment. What is important for tissue engineering is to maintain cell plasticity in a manner that, once in vivo,

Finally, a third highly modern and innovative approach for tissue engineering, which shares the truly tridimensional environment for cells as provided by hydrogels, but which doesn`t suffer from vascularization and cell organization limitations presented by the latter, is the organ printing tissue engineering technique, which will be more thoroughly described in the

As listed before, the main limitations of the solid scaffold approach include the low level of precision in cell placement, especially when engineering multicellular constructs, considering

strongly contribute to novel advances in the field.

306 Advances in Biomaterials Science and Biomedical Applications

technique [64].

promote construct viability *in vivo*.

as bidimensional surfaces by cells.

implanted cells start to behave as host cells would.

differential gene expression.

next section.

**3.3. Organ printing**

According to the First International Workshop on Bioprinting and Biopatterning, organ printing was defined as "The use of material transfer processes for patterning and assembling biologically relevant materials (molecules, cells, tissues, and biodegradable biomaterials) with a prescribed organization to accomplish one or more biological functions" [45]. In fact, this technology could be defined as computer-aided, layer-by-layer deposition of biologically relevant materials [53].

The ultimate goal of organ-printing technology is to fabricate 3D vascularized functional living human organs suitable for clinical implantation in reasonable time scales. Other applications of this technology are in histogenesis and organogenesis, pharmacological tests and disease research [45, 46, 47, 48].

Wilson and Boland (2003) showed protein and cell printing using a commercial ink-jet device can be possible. In this technology, either individual cells or small clusters are printed over ECM hydrogels, designed for involve printed cells and to provide them with desired signals. Therefore, organ printing has derived from hydrogel classical approach, described in the previous section. The method is rapid, versatile and cheap. Its disadvantage is that it is difficult to assure high cell density needed for the fabrication of solid organ structures. Furthermore, due to the high speed of cell deposition, considerable damage is caused to cells, although the latest developments in the field have led to considerable improvement in cell survival [54, 55]. In the other approach, mechanical extruders are used to place 'bio-ink' particles, multicellular aggregates of definite composition into a supporting environment, the 'biopaper', according to computer-generated templates consistent with the topology of the desired biological structure. Organoids are formed by postprinting fusion of the bio-ink particles and the sorting of cells within the bio-ink particles. The advantage of this technology is that the bio-ink particles represent small 3D tissue fragments. Thus, cells in them are in a more physiologically relevant arrangement, with adhesive contacts with their neighbors, which may assure the transmission of vital molecular signals. Both inkjet and extruder bioprinting are compatible with rapid prototyping [55].

In this biomanufacturing a precise layer-by-layer placement of self-assembled tissue spheroids in sprayed tissue fusion permissive hydrogels is used to obtain an organ or tissue. The hydrogels work as "biopapers" and cell blocks or tissue spheroids work as a "bioink". Both biopaper and bioink must be optimized in order to obtain viable tissues. For instance, biopapers vary according to each "printed" organ, and cell spheroids vary in properties according to their composition. Cell viability during and after printing is an obvious goal for bioprinting [45]. Preliminary studies of both ink-jet and laser forward transfer indicated that cells can survive deposition condition forces. Problems associated with ink-jet delivery of cell suspensions may also come about from the high shear stresses observed during ejection and impact of a fluid drop [50, 51, 52].

In many cases, the bioprinting process requires that before and during printing, cells and molecules must be carried in a fluid vehicle that shortly after printing requires consolidation and should consequently behave as a viscoelastic solid. This phase change must occur without damage to the biochemical, cells, or more complex units within the fluid, which presents a considerable challenge. Concurrently, tissue printed mustn't be too solid, or cell spheroids won`t interact and form a continued tissue.

Organ printing is a technology that promises to transform tissue engineering into a commercially successful biomedical industry. Unlike other tissue engineered approaches, organ printing involves the high throughput generation of organs, relying on automated cell sorters, cell and organ bioreactors and robotic bioprinters, most of them which are already commercially available [46]. However, much research is necessary to turn this technology into reality of clinical application.
