Pharmaceutical. 3D scaffolds in tissue engineering and regenerative medicine: beyond structural templates? Review

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1 Pharmaceutical Pharm. Bioprocess. (2013) 1(3), D scaffolds in tissue engineering and regenerative medicine: beyond structural templates? The objective of this article is to systematically present the emerging understanding that 3D porous scaffolds serve not only as structural templates for tissue fabrication but also provide complex signaling cues to cells and facilitate oxygen and therapeutic agent delivery. Strategies in the field of tissue engineering and regenerative medicine often rely on 3D scaffolds to mimic the natural extracellular matrix as structural templates that support cell adhesion, migration, differentiation and proliferation, and provide guidance for neo-tissue formation. In addition to providing a temporary support for tissue fabrication, 3D scaffolds have also been used to study cell signaling that best mimics physiological conditions, thereby expanding our understanding beyond 2D cell cultures. It is now understood that cell responses to 3D scaffolds are distinctively different from 2D surfaces. Recently, 3D scaffolds emerged as a vehicle for improved oxygen transport to seeded cells and also to deliver relevant therapeutic agents to facilitate tissue formation and/or to regenerate damaged or otherwise compromised tissue functions. In this review, our goal is to provide recent advances made in 3D scaffolds to modulate tissue formation, cell signaling, mass transport and therapeutic agent delivery. Tierney GB Deluzio 1, Dawit G Seifu 1, Kibret Mequanint* 1,2 1 Department of Chemical & Biochemical Engineering, The University of Western Ontario, London, Ontario, N6A 5B9, Canada 2 Graduate Program of Biomedical Engineering, The University of Western Ontario, Canada *Author for correspondence: Tel.: Ext Fax: kmequani@uwo.ca The strategies of regenerative medicine and tissue engineering are conceptually simple and appealing, yet these have proven to be challenging engineering tasks. Despite rapid advances made in this field [1,2], success is still limited due to significant knowledge gaps in our ability to control, coordinate and direct tissue formation, which are the ultimate goals for both tissue engineering and regenerative medicine. One strategy of tissue engineering, among others, involves seeding cells onto a porous 3D scaffold that supports in vitro tissue formation and maturation. The resulting engineered tissue is implanted into a patient where it further grows, through self-repair remodeling. Therefore, the ultimate objective of tissue engineering is to develop responsive living tissues with properties similar to those of the native tissues they are intended to replace and, can be applied to many, if not all, tissues in the body. Not surprisingly, tissue engineering continues to be a promising alternative to current treatments for diseased and damaged organs, as well as applications in a variety of other areas such as drug research and discovery [3 5]. With the exception of some studies [6 8], exogenous porous 3D scaffolds that mimic the extracellular matrix (ECM) are required for the growth of cells to form engineered constructs [9,10]. Depending on the intended application, scaffolds may be designed to be biodegradable so that only the neo-tissue will remain after a given period of culture time following implantation, or they may be biostable where a composite tissue that provides long-term support is fabricated [11 14]. In the case of biodegradable scaffolds, cells /PBP Future Science Ltd ISSN

2 Deluzio, Seifu & Mequanint Key Term will remodel the scaffold with their Neo-tissue: Tissue that grows to own ECM proteins creating the eventually replace the scaffold intended tissue without compromising tissue structural integrity. This, template so that only the newly formed tissue remains. however, requires strict coordination of the scaffold biodegradation rate with the biosynthetic rate, and this is one of the major obstacles in the field today. In addition, a scaffold must have several required characteristics: biocompatibility, appropriate mechanical strength and compliance, optimal porosity for cell seeding, in vitro nutrient and oxygen transport, and the ability to bind to cells and release growth factors when needed. Although some of these criteria could be met with existing scaffolds, they do not always provide biological cues for the cells embedded in them and do not interact with the cells. In the body, cells reside within the ECM, which provides tissues with the appropriate architecture, as well as signaling cues that influence key cell function such as adhesion, migration, proliferation, differentiation and secretion of ECM components [15]. Fabrication of tissues in vitro thus requires that cells be given a more specific level of instruction so that tissue regeneration in the host is successful. With the discovery of cell adhesion peptide domains in fibronectin, collagen and laminin, the design of synthetic extracellular matrices with biological activity has become a valuable strategy to impart biomimetic properties [16 18]. The selection of scaffold type and material depends on the specific tissue engineering and regenerative medicine application, as well as the applicable design criteria. Natural materials such as collagen [19], chitosan [20] and hyaluronic acid [21] have the advantage of being generally nontoxic, in addition to providing biological cues to promote cell attachment and proliferation. They are, however, difficult to fabricate due to their limited processing parameters and often result in constructs with poor mechanical properties. Synthetic materials, on the other hand, are readily available and generally easy to modify, with minimal batch-to-batch variations. However, they do not innately possess appropriate sites to enhance cellular interactions and therefore lack biological activity. The method of scaffold fabrication also significantly impacts the physiochemical properties of the resulting tissue engineered construct. Other aspects, including reproducibility and cost effectiveness, should also be considered when selecting materials and fabrication techniques. Common methods for scaffold fabrication include solvent casting/particulate leaching and electrospinning, with more advanced techniques, such as rapid 3D plotting, solid free form and 3D projection stereolithography, being developed [22]. As the purpose of this review is to discuss the role 3D scaffolds play in tissue engineering and regenerative medicine, the reader is referred to dedicated literature reviews on scaffold fabrication strategies [23 25]. 3D scaffolds as templates for clinically relevant & model tissue fabrication One obvious and extensively studied area of 3D scaffold application is their use as templates for the engineering of a variety of tissues including, but not limited to, bone, cartilage, cardiovascular, nerve, skin, bladder and airway tissue. A properly designed 3D scaffold acts as a support for the growing tissue, mimicking the role of the natural ECM, providing cues for proliferation and differentiation while providing the desired shape for the tissue (Figure 1A). Although the quest for the ideal scaffold is far from over, a number of studies have demonstrated the strategy of scaffold-guided tissue engineering where full scale and functional tissues have been developed [9,26 28]. In this section, the authors have selected specific examples to demonstrate the clinical relevance of scaffold-guided engineered tissues due to available clinical data. The first example is the advances in bladder reconstruction (when the native tissue loses its ability to store and empty effectively as a result of numerous medical conditions) [29]. A demonstration of autologous hollow organ reconstruction using tissue engineering methods was the creation of a transplantable urinary bladder neo-organ whereby, a bladder-shaped 3D biodegradable scaffold fabricated from polyglycolic acid coated with polylactic co-glycolic acid (PLGA), was used as a structural template [30]. Bladder urothelial and smooth muscle cells harvested from canine subjects were seeded onto the 3D scaffolds, with the urothelial cells on the luminal surface and the muscle cells on the exterior surface and then cultured in vitro. The resulting neo-tissues were implanted into the same animals from which the cells were harvested, after the animals underwent a trigone-sparing removal of their native bladder. The tissue-engineered neo-bladders were able to surpass the native bladder capacity and were structurally and functionally indistinguishable from native bladders. The resulting organ exhibited normal cellular organization with a trilayer of urothelium, submucosa and muscle as found in the native bladder. The immune response from scaffold fragments was limited with no observation of upper tract obstruction, lithogenesis, encrustation or other abnormalities. On a more practical clinical note, a composite of collagen and polyglycolic acid 3D scaffolds seeded with the patients own urothelial and muscle cells was successfully implanted into seven patients with adequate mechanical properties, structural architecture and phenotype after a 4-year followup [31]. Although this is a first step towards the goal of 268 Pharm. Bioprocess. (2013) 1(3)

3 3D scaffolds in tissue engineering & regenerative medicine D1 Amount released A1 A2 A3 Time D A B B1 3D scaffolds C C1 F F F F F F F F F F F F F F F F F O Si O O B2 B3 C2 C3 C4 Figure 1. The versatility of 3D scaffolds in tissue engineering and regenerative medicine applications. (A) As an illustrative example, tubular biodegradable fibrous scaffolds can be seeded with vascular cells to infiltrate the cross-section and remodel the scaffolds leading to a mature engineered neo-tissue. (B) Different signaling molecules (such as growth factors) can be conjugated onto 3D scaffolds (B1). As an example, when smooth muscle cells are seeded on laminin-modified scaffolds, a contractile phenotype was observed (B3) whereas fibronectin conjugation led to a synthetic phenotype characterized by the lack of SM α-actin expression (B2). (C) Scaffolds can be fabricated with embedded oxygen vectors in which fluorinated zeolite particles were incorporated (C2) and cells were seeded (C3) such that cell distribution throughout the scaffold thickness was uniform without being necrotic (C4). (D) Scaffolds can also be used to release therapeutic drugs in a sustained manner (D1). In figures A2, B2, B3, C3 and C4, blue is nuclei, green is F-actin. In figures B2 and B3, red is smooth muscle α-actin. engineering fully functional bladders, such engineered tissues show great promise in urologic tissue regeneration and underscores the utility of 3D scaffolds for fabricating clinically relevant tissues. The fabrication of 3D scaffolds for engineering full scale organs/tissues of clinical relevance is not restricted to synthetic biomaterials as decellularized donor tissues can also be seeded with the patient s cells. This method has been successfully used in a young patient born with long-segment congenital tracheal stenosis and pulmonary sling [32]. In fact, there are reported cases where a biologic scaffold composed of allogeneic extracellular scaffolds derived from heterologous sources and recellularized with autologous stem cells or differentiated cells are preferred over synthetic materials [33], as demonstrated by earlier clinical transplantations of tissue-engineered airways [34,35]. Furthermore, for tracheal tissue engineering, scaffolds with air- and liquidtight seals as well as adequate structural support (lateral rigidity and longitudinal flexibility) to maintain airway patency and allow rapid epithelialization are required [33]. Such directional variation on mechanical properties (i.e., anisotropy) is unmet by synthetic scaffolds fabricated from only one type of biomaterial. In view of this, only a composite scaffold was used to engineer an airway tissue in an only reported clinical test for a single patient [36]. Results from this study demonstrated proliferation of cells and the growth of an ECM-like coating. Levels of regenerative-associated plasma factors were found to be amplified, suggesting homing of stem cells, cell-mediated wound repair, remodeling of the ECM and neovascularization of the graft. In addition, large granulation areas with initial signs of epithelialization and more organized vessel formation were evident. The scaffold eventually recellularized into a functional tissue containing site-specific 269

4 Deluzio, Seifu & Mequanint Key Term cells and integrated into the adjacent Extracardiac cavopulmonary tissue. The patient became asymptomatic, tumor-free and the ability connections: Process by which venous blood flows to the lungs, to breathe normally was restored [36]. bypassing the right ventricle. This The use of 3D scaffolds has also procedure is often performed to repair tricuspid atresia or for been successfully demonstrated to pediatric patients with congenital fabricate clinically relevant vascular substitutes. In view of this, heart problem. the first clinical application of a scaffold-guided tissue-engineered vascular construct was reported in 2001 for a pediatric patient with congenital single ventricle cardiac anomaly [37]. This was an exceptional case because no synthetic graft could be used with the capacity to grow, repair and remodel as required with normal development. Since then, tissue-engineered conduits have been used for a total of 25 extracardiac cavopulmonary connections (with a median patient age of 5.5 years) with only two late cardiac failure deaths reported [38,39]. The success of this procedure can be partially attributed to the relatively low-pressure environment (20 30 mmhg during systole) found in pulmonary circulation, which is less demanding than pressures found in coronary arteries ( mm Hg during systole). Despite the urgent need for an engineered vascular tissue substitute, success is still limited [40]. While engineering approaches for other tissue substitutes can rely on in vivo remodeling to approach functionality with time, tissue-engineered blood vessels must function immediately on implantation a significant challenge that contributed to its limited success. The long lead time associated with the fabrication of an autologous vascular substitute is commonly invoked as a major limitation to widespread clinical use. While this point may be valid for emergency coronary artery surgical procedures or critically ischemic limbs, in practice, most coronary and distal vascular bypass procedures can be predicted and delayed over extended periods, allowing sufficient time for tissue fabrication [40]. Clinical trials of engineered vascular substitutes for the adult population has been initiated to examine the use of these grafts as arteriovenous shunts, as well as coronary and lower limb bypass grafts [41]. The initial clinical trial was focused on the safety of the arteriovenous shunt model due to the lack of suitable vein for hemodialysis and the deplorable efficacy of synthetic vascular grafts. Although graft failure in this model is unlikely to be life- or limb-threatening, the high flow rates encountered ( 800 ml/min) generate considerable hemodynamic loads [40]. Notwithstanding both technological and regulatory challenges that lie ahead, engineered vascular tissues continued to be the holy grail of future vascular intervention. The above illustrative examples demonstrate the clinical feasibility of some engineered tissues, but it should also be pointed out that most, if not all, were used under extreme circumstances for which no alternative interventions were amicable. For these tissues, approval for clinical use is often restricted at the local hospital level. It is therefore understood that the clinical use at this stage is an exception and experimental rather than a standard treatment option. With relevant regulatory approvals, the clinical prospect of engineered tissues remains promising. This has been demonstrated for tissue-engineered skin, which is now a clinical reality for treating newborns with epidermolysis bullosa and burn patients [42 44]. Although the long-term goal of engineered tissues is for the clinic, the immediate significance of engineered tissues can also be viewed within the context of species-specific predictive organ model. For instance, the use of vascular cells and whole sections of a harvested artery in combination with animal models to study vascular diseases (e.g., atherosclerosis, post angioplasty restenosis and hypertension) in an attempt to develop therapeutics is not new, but employing engineered human vascular tissues for this role is a novel concept. While conventional 2D cell cultures are indispensable to our understanding of tissue morphogenesis and function in physiological and pathological states, they do not accurately replicate the 3D microenvironment of human tissues [15]. For example, 2D culturing of vascular cells for studying intimal hyperplasia without the arterial wall structure and ECM cannot recapitulate the intricate vascular wall mechanics and morphogenesis [45]. Similarly, animal organ cultures and whole animal models do not completely mimic human biology due to the inevitable inter-species difference [45]. Studies using closely related nonhuman primates are constrained by limited availability, legal restrictions, ethical concerns and high cost, making animal models impractical. When studying human vascular diseases and therapeutics, a realistic model is a human tissue but the inability to experiment directly on human subjects limits this progress. Thus, the need for an engineered human vascular tissue model to close this gap is of vital importance. Engineered human vascular tissues are unlikely to replace animal or human subjects; however, they have the potential to provide high throughput, substantive and detailed information regarding very specific conditions under controlled environments to study disease models and therapeutic outcomes that are not possible with animal-based models. The impact of successfully engineered human tissues is, therefore, not only restricted to the clinic but also fills a critical gap in the preclinical model tool chest, between traditional cell culture and whole animal experiments, and has the potential to accelerate 270 Pharm. Bioprocess. (2013) 1(3)

5 3D scaffolds in tissue engineering & regenerative medicine the pace of basic biomedical research [46]. A number of important physiological characteristics of native tissues appear to be preserved in an engineered tissue; thus providing models for specific disease conditions such as elevated contractility of vascular smooth muscle cells in hypertension, elevated proliferation in atherosclerosis and post-angioplasty restenosis, and fibrosis and cardiovascular remodeling [3,47 50]. In this context, engineered vascular tissue technology may be used both to validate drug targets and to optimize loads. This allows for cardiovascular drug screening in a more controlled and efficient way than can be performed using a traditional whole animal approach, thereby minimizing the number of laboratory animals used and decreasing the overall cost of performing research [46]. In recent years, engineered 3D tissue models such as cardiac patch [4,10], lung tissue [51], cornea [52 54] and solid tumor [55] have emerged as powerful tools for drug discovery. Although clinical applications of engineered 3D tissues attract the most media attention, it is evident that engineered tissue models can serve as platforms for tightly controlled, high-content screening of drugs and for pharmacodynamic analyzes. 3D scaffolds as ECM-mimicking microenvironments for signal transduction & gene regulation studies In tissue engineering and regenerative medicine strategies, signals that are stimulated when cells are seeded to scaffolds provide important information on early cellular events. In addition to the chemical composition, cell signal transduction has been predicted to be highly dependent on the dimensionality and architecture of scaffolds [56,57]. The merits of culturing cells in 3D rather than on 2D flat surfaces have been accelerated by the difference it makes to cell s behavior, which is much closer to their state in vivo [15]. A number of studies have compared biological responses and signaling during cell interactions with standard 2D cell cultures versus 3D matrices [57,58]. There are two general approaches that have been taken to investigate the role of 3D matrices in cell signaling (Figure 2). The first approach is the utilization of cell-derived 3D matrices whereby fibroblasts are cultured in a confluent state to produce the ECM as structural support for their adhesion, migration and tissue organization [57,59]. The ECM components produced by cultured fibroblasts are then denuded of cells using detergent, followed by the removal of cellular debris (Figure 2A & B) [60]. This treatment produces a 3D fibrillar ECM comprised of fibronectin with varying levels of collagen, heparan sulfate proteoglycan and laminin that is intact and cell-free. Fibroblast responses to these 3D matrix components were strikingly different compared with the same ECM components in 2D forms [15,57]. Contrary to 2D surfaces, initial fibroblast cell adhesion increased on 3D matrices and was solely dependent on α5b1 integrin engagement. Furthermore, tyrosine phosphorylation of FAK at residue 397 was poor in cells cultured on 3D matrices while activation of the MAPk ERK1/2 was enhanced. Despite unchanged activation levels of Rho and Cdc42, Rac activation was considerably low. Given that the cellderived 3D matrices are fibrillar and porous, it is reasonable for these matrices to be more pliable than compressed 2D rigid films of the same ECM components (Figure 2C). This may, in turn, explain the observed signaling differences between 3D matrices and 2D ECM films derived from cells. While FAK phosphorylation is often needed for the sustained downstream ERK1/2 phosphorylation in 2D cultures, 3D matrices do not appear to require the same upstream signal. Instead, it is shown that 3D matrices induce sustained activation of ERK1/2 via Src/Ras/Raf signaling pathway [57]. The second approach for studying cell signaling and gene regulation in 3D cultures is the utilization of either a naturally occurring (such as collagen gel) or a synthetic 3D scaffold. Using a genome-wide gene expression analysis of osteoblast-like cells grown in the 3D collagen matrix, the partial suppression of genes associated with cell adhesion and cell cycling compared with 2D surfaces is reported [61]. Western blot analysis revealed that the expression of the phosphorylated p130cas, FAK and ERK1/2 was reduced in cells grown in the 3D matrix. Conversely, phosphorylation of p38 MAPK was elevated in the 3D matrix and its upregulation was linked to an increase in mrna levels of DMP1 and bone sialoprotein [61]. Recent progress in miniaturizing scaffold-guided engineered tissues and associated physiological assay systems are postulated to accelerate our knowledge regarding interaction of the genome with both the phenome and the environment [46]. For example, using DNA microarrays with 9600 genes, Chien and coworkers [62] demonstrated that 77 vascular smooth muscle cells (VSMC) genes were expressed more than twofold and 22 genes were expressed in levels less than one-half in 3D matrix when compared with the 2D culture condition. Specifically, cells in 3D had less stress fibers and focal adhesions, and a lower level of tyrosine phosphorylation of FAK. The cyclin-dependent kinase inhibitor 1 (p21) was differentially upregulated in 3D cultures leading to lower VSMC proliferation. Collagen Type I expression was also higher in 3D suggesting that VSMC cultured on 3D matrix have increased ECM synthesis. In addition, Mequanint and coworkers [63] demonstrated upregulation of the elastin gene and downregulation of differentiation marker genes when VSMCs were cultured on 3D scaffolds compared with 271

6 Deluzio, Seifu & Mequanint A 1. Cells plated at high density 2. Synthesis of 3D extracellular matrix C Extracellular signals Plasma membrane 3. Detergent extraction of cells 4. Plate new cells??? Pliable matrix Ras Rigid matrix Raf B 3D matrix Purified matrix component Rho MEK Compressed (with weight) Solubilized matrix Rac/ cdc42 ERK Cytoskeletal actin fibers Elastin transcription Flattened to a 2D matrix Plated as a uniform 2D matrix Pharm. Bioprocess. Future Science Group (2013) Figure 2. 3D scaffolds play an important role in cell signaling. (A) Preparation of fibroblast-derived 3D matrices composed mainly of fibronectin fibrillar lattices and (B) 2D controls with the same molecular composition as the 3D matrix. (C) Synthetic electrospun 3D scaffolds for studying cell signaling. The exact signal that is transmitted from the 3D scaffold to the plasma membrane is not fully understood but significant downstream signal divergence exists based on the pliability of the 3D matrix. (A & B) Adapted with permission from [60]. Key Terms Biomimetic scaffolds: Fabricated from either natural or synthetic materials and often surface-modified with different proteins to mimic the natural extracellular matrix environment in order to mediate and enhance the cell material interactions. Ligand-conjugated scaffolds: Fabricated by conjugating a specific ligand for integrin signaling onto in order to facilitate interactions with cells such that they elicit a specific cellular response. 2D surfaces. As exemplified by the above studies, differential signaling and gene expression are not limited to vascular cells. Several groups have reported preliminary evidence that gene expression of different cells cultured in 3D more closely parallels the in vivo situation [64 66]. Gene expression profiling experiments in various cell types clearly demonstrated close correlation between engineered and native tissues in tumor cells [64], tendon [67], bone [61] and skin [68]. Collectively, these studies point to engineered tissues as model systems that could be used to test gene expression and to study the effect of altered gene expression on function in vitro. Given the availability and ease of fabrication techniques, a number of synthetic 3D scaffolds are frequently used to culture different cell types for signaling studies. In a study that utilized a 3D electrospun polyamide nanofiber scaffold, fibroblasts activated Rho, Rac and Cdc42 three of the most characterized members of the Rho GTPases that are regulators of the actin cytoskeletal assemblies [56]. In particular, the 3D polyamide surface caused a significant increase in preferential and sustained activation of Rac. This increased activation was coupled with observations of other changes in the cells, such as morphology and proliferation. It is worth noting that, contrary to this study, Rac was downregulated when cell-derived fibrillar 3D matrices were used [69]. At a glance, this may appear to indicate differences in composition between the cell-derived matrix and the synthetic polyamide matrix. This, however, does not 272 Pharm. Bioprocess. (2013) 1(3)

7 3D scaffolds in tissue engineering & regenerative medicine appear to be the case. Data supporting this argument comes from studies of VSMC cultured on 3D polyurethane scaffolds where it was clearly demonstrated that Cdc42, which is a Rho GTPase like Rac, was undetectable [58]. In view of the aforementioned studies, scaffold pliability is a very likely factor where an elastomeric polyurethane which has a glass-transition temperature of approximately 1 C is more pliable at cell culture conditions whereas the polyamide fibers are rigid and crosslinked with glass-transition temperature exceeding 37 C making them stiff. Collectively, these studies suggest that signaling cascades are activated differently in 3D than their 2D counterparts. Although more studies are still needed, these 3D signaling events appeared to be regulated, not by the type of cells studied, but by the pliability of 3D scaffolds whereby rigid 3D matrices promote Rho signaling while Ras signaling is activated by pliable 3D scaffolds for all the cells studied so far (Figure 2C). Ligand-functionalized biomimetic scaffolds As an ECM surrogate, scaffolds should provide more than a structural support to both the seeded cells and to the growing engineered tissue. The ECM is an intricate network of macromolecules and plays a very complex and dynamic role in cell functions. In addition to the obvious mechanical role, the ECM serves as a reservoir for a number of growth factors and cell signaling components including ligands for integrin signaling, thus suggesting the possibility of both spatial and temporal information exchange between cells and the matrix. The complex signals experienced by cells include biochemical molecules (proteins, hormones and growth factors), mechanical forces, and cell cell interactions. Biomimetic scaffolds offer a means to mediate and enhance cell-material interactions for tissue engineering, and can be fabricated by modifying (either chemically or physically) synthetic polymers with bioactive peptides, growth factors and other biomolecules. Alterations in the bulk chemistry of the scaffold material can consequently affect the mechanical properties, whereas surface modifications with biomolecules do not compromise the mechanical integrity of scaffolds extensively. Physically adsorbed bioactive molecules on a 3D scaffold have the disadvantage of being leached into the surrounding media with a gradual loss of the bioactive compound. The covalent attachment of bioactive molecules to 3D scaffolds overcomes this problem; however, it may be difficult to chemically attach the desired molecule, especially without affecting its biological activity [70]. Cellular interaction with adhesion proteins in the ECM influences the morphology, motility, gene expression and ultimately the survival of cells [71]. Since the discovery of the essential cell adhesion RGD sequence in fibronectin [72 74], there has been considerable research to covalently attach this protein to synthetic polymers to impart biomimetic properties. Instead of fibronectin, many studies have immobilized the short chain RGD peptide to polymers through an amide bond between an activated surface carboxylic acid group on the polymer and a nucleophilic N-terminus of the peptide (detailed in [75]). The choice between using a protein versus a shorter peptide sequence in engineering biomimetic scaffolds can be difficult, as both options have positive and negative aspects. For example, proteins must be isolated and purified from other organisms, and may induce undesirable immune responses. Proteins are also subject to proteolytic degradation, thereby limiting their long-term use [75]. Proteins (either adsorbed or immobilized on a scaffold) experience hydrophobic forces different from their native environment, thus any resulting conformational change in the protein may alter the presentation of key motifs, if not bury them. This would reduce the effective biological activity of biomimetic scaffolds. Small peptides, on the other hand, tend to have higher stability towards heat sterilization, ph variation, storage and are easier to characterize and are cost-effective [75]. Furthermore, since peptides occupy less space than proteins, peptides can be packed at higher densities on the modified scaffold surface [75]. Still, a single protein contains several motifs that are recognized by cells, whereas a small peptide fragment represents only a single motif. For the case of fibronectin, it has been shown that this protein contains a multitude of recognized domains other than the RGD motif, all of which play a role in the complexity involved in signaling pathways. For example, the REDV sequence in the CS5 region (a total of 20 amino acids) of fibronectin supports endothelial cell attachment and spreading [76]. Different approaches to conjugate fibronectin on various types of biomaterials have been reported [77 81], although the majority of these studies were conducted on films rather than on porous 3D scaffolds that would ultimately be used to fabricate viable tissues. Despite this, some insightful studies have demonstrated fibronectin conjugation on PEG diacrylate and methacrylic acid 3D scaffolds [82], on a fibrous 3D mesh of nonwoven PET via a gluaraldehyde crosslinker [83], cylindrical 3D scaffolds fabricated from PCL PEG PCL block copolymers [84], and more recently to 3D polyurethane scaffolds [16]. All these studies documented favorable cell interaction with the fibronectin conjugated scaffolds presumably due to integrin ligand binding. Most studies conducted regarding ligandconjugated scaffolds focused on either RGD or fibronectin conjugation with the primary goal of achieving cell adhesion to 3D scaffolds. However, as advances are 273

8 Deluzio, Seifu & Mequanint Key Terms 3D cell signaling: Stimulation/ activation of cell functions such as adhesion, migration, proliferation, differentiation and secretion of matrix components using cues provided by 3D tyopography as opposed to conventional 2D surfaces. Prevascularization: Process by which engineered tissue constructs are rendered to have a capillary-like structure to facilitate host integration and perfusion of nutrients. made in 3D cell signaling, conjugation of other signaling molecules to scaffolds is inevitable. While some ligands can be presented to cells in a soluble form, for example through the addition via culture medium, other signaling molecules needed to be conjugated. For instance, in stem cell fate decisions and in vascular tissue engineering, the Notch signaling pathway plays an integral role in cell differentiation, proliferation, and apoptosis [85]. Notch transmembrane receptors can be activated by the ligands Jagged-1 and -2 and Delta-1, -3 and -4 through cell-to-cell contact or by immobilization of Notch ligands to scaffolds [86]. Thus, Notch ligand functionalized scaffolds would allow for enhanced therapeutic strategies with the ability to control cell signaling and gene expression. In addition, development of bioengineered surfaces with immobilized Notch ligands would provide improved methods for studying the Notch signaling pathway in an isolated way [87,88]. A number of studies have emerged regarding the immobilization of Notch ligands onto the surfaces of 3D porous scaffolds. In one such study, the Notch signaling ligand Jagged-1 was attached to self-assembled monolayers at different densities with the objective of controlling the orientation and conformation [85]. x-ray photoelectron spectroscopy and ELISA were used to confirm the presence of Jagged-1 on the surface at varying concentrations. To test the bioactivity, HL-60 leukemic cell line was cultured on the functionalized constructs and the resulting gene expression, specifically the activation of the Notch receptor for the Hes-1 family, was observed with real-time PCR studies. Interestingly, results showed that an increase in the concentration of Jagged-1 did not directly cause an increase in the levels of Notch target gene expression. Instead, it was determined that the favourable exposure of the ligands with the desired orientation and improved availability increased the levels of corresponding gene expression. In a further study, Notch ligand coated 3D scaffolds were investigated for studying human hemotopoietic stem cell niches in vitro [89]. A layer-by-layer method was used to fabricate inverted colloidal crystal scaffolds with immobilized Delta-1 Notch ligand on the surface of the pores. Scanning electron microscopy confirmed the presence of open, uniform and interconnected pores that were unaffected by the ligand incorporation. Stem cells derived from blood from bone marrow and umbilical cord were cultured on the modified construct using a rotary cell culture vessel to facilitate physical contact between the ligands and cells [89]. The corresponding biological activity was confirmed by the T-cell differentiation of the human hemotopoietic stem cells through the progression of differential stage-specific surface marker expressions. Such functionalized scaffold could, therefore, be applied as an improved tool to investigate human stem cell biology on the molecular and tissue level. It should be, however, noted that experimental conditions including ligand density may lead to variations in cell activities. 3D scaffolds as oxygen reservoirs in tissue engineering The success of engineered tissues has been limited by the inability to deliver sufficient oxygen to the growing constructs. Oxygen delivery, in particular, is a limiting step for clinically relevant size tissues because of its low solubility in culture media. This is further exacerbated by the fact that cells consume 5 6 moles of oxygen per mole of monosaccharide according to the following mole balance: C 6 H 12 O6 + 6O 2 6CO 2 + 6H 2 O Clearly, the delivery of this much oxygen to cells requires the development of an oxygen delivery system that is more efficient than molecular diffusion alone. Previous attempts to overcome this limitation involve perfusion bioreactors where oxygen dissolved in the culture medium diffuses to the scaffold interior. For this reason, bioreactors have been developed with the intent to mechanically stimulate the growing tissue constructs with physiologically relevant forces and to improve mass transfer within the tissue. Although the former is largely successful, the latter has shown to be a formidable engineering task. Therefore, the delivery of nutrients and the removal of metabolic waste materials remained to be a fundamental consideration for fabricating engineered tissues [90 92]. In the case of flow-induced mass transfer, the high flow rate required to maintain an adequate oxygen concentration for cell viability often surpasses the shear stress tolerance of the cells [93]. One feasible approach involves the use of perfluorocarbon (PFC) emulsions as an oxygen carrier emulating the role of hemoglobin in the physiologic system [92]. However, unlike oxygen chemically bound to hemoglobin, solubilized oxygen can be rapidly and extensively extracted from PFC molecules. The oxygen unloading of the PFC emulsion is facilitated by its increased surface area. The disadvantage of these emulsions is that their high density causes them to settle in the medium. From this, it can be inferred that binding the oxygen carrier molecules to the scaffold would be advantageous. One method for improving the delivery of oxygen to cells is incorporating oxygen generating chemicals into the 3D scaffolds. Calcium peroxide (CPO), which 274 Pharm. Bioprocess. (2013) 1(3)

9 3D scaffolds in tissue engineering & regenerative medicine decomposes in water to form calcium hydroxide and hydrogen peroxide, which further decomposes to oxygen and water, was embedded into PLGA scaffolds [94]. These porous 3D constructs were prepared via particulate leaching with paraffin as a porogen and contained calcium peroxide at concentrations of 0, 1, 5 and 10 wt%. Scanning electron microscopy data suggested that the pore size and porosity of the scaffold were unaffected by the incorporation of CPO, as the scaffold maintained a highly porous and open pore structure. To test the effect on oxygen delivery, NIH3T3 fibroblasts were seeded onto PLGA and PLGA CPO scaffolds with the addition of catalase to capture the hydrogen peroxide by-products that may be toxic to cells. Under normoxic conditions (21% oxygen, 5% carbon dioxide), significant cell viability with the incorporation of 5% CPO was shown. When scaffolds containing 5% CPO were seeded with cells and cultured in a hypoxic environment (1% oxygen, 5% carbon dioxide), metabolic activity on the PLGA CPO scaffold increased significantly compared with control scaffolds. The hypoxic condition was chosen so that metabolic results are attributed to oxygen generation from the CPO scaffold as opposed to 21% oxygen. A similar study of incorporating CPO into 3D polycaprolactone (PCL) nanofibers for the antibacterial properties of the calcium hydroxide produced during oxygen generation was carried out [95]. The reported results were not consistent with the previous study [94] since the scaffolds were reported to be cytotoxic to human osteoblast cells. The cytotoxicity, however, appeared to be temporary as cells regained a healthy status and spread over the nanofiber mesh. This negative effect may be attributed to the initial burst release of calcium hydroxide. The method of incorporating calcium peroxide into 3D porous scaffolds demonstrates potential for enhancing oxygen delivery to seeded cells. However, calcium hydroxide and hydrogen peroxide are a strong base and oxidizing agent, respectively, which could be detrimental to cells seeded on these scaffolds. Inclusion of catalase or ascorbic acid to protect against oxidative stress and remove potentially harmful byproducts could help to alleviate these effects. Based on the inconsistency in the reported cytotoxicity of these by-products, further studies are required to determine the amount of cell protector required to fully prevent toxic effects. In addition, this method delivers oxygen for a limited time, until the CPO is completely consumed. Further studies must be done to determine if this is enough time to extend cell viability until tissue maturation is established and if the amount of oxygen delivered could support larger constructs, with higher densities of cells [96]. A recent study in the authors laboratory tested the oxygen delivery capability of fluorinated porous zeolite Y particles embedded into a 3D polyurethane scaffold [96]. The zeolite Y particles were prepared and then fluorinated with 1 H, 1 H, 2 H, 2 H-perfluorodecyltriethoxysilane resulting in fluorinated-zeolite (FZ) particles with particle diameters between 850 and 1000 nm (Figure 1C). 3D polyurethane scaffolds containing 2 wt% embedded FZ particles were fabricated by the solvent casting and particulate leaching method with NH 4 Cl porogens. Data confirmed that the FZ particles were not leached out during the fabrication process. In addition, results demonstrated open and well-defined pores with high interconnectivity, uniform distribution of FZ particles throughout the scaffold, and that the FZ particles were mostly contained at the surface of the scaffold, which is crucial for oxygen extraction by cells. This was important for efficient cell infiltration and nutrient transport. To test the effect of the FZ particles on oxygen delivery, human coronary artery smooth muscle cells (HCASMC) were seeded onto the FZ-modified and control scaffolds and were cultured for 4, 7 and 14 days. Cell number increased significantly after 4 and 7 days of culture on the PCU FZ scaffolds compared with control scaffolds likely attributed to increased oxygen delivery. Although the HCASMC attached and spread on both control scaffolds, the infiltration depth was double on the FZmodified scaffolds, suggesting enhanced availability of oxygen at greater depths in the scaffold. As reiterated in this section, the delivery of oxygen to cells seeded on scaffolds is vital for successfully developing tissues of clinical relevance. The preceding cited studies have provided platforms to incorporate oxygen delivery vectors into 3D scaffolds without adversely affecting the porosity or morphology. An important factor is the length of oxygen delivery time provided by these vectors. Oxygen must be delivered to the cells seeded on the scaffold in sufficient quantities until the engineered tissue matures. Development of an improved method for oxygen delivery to cells and preventing cell necrosis would allow for enhanced tissue structures, ultimately leading to the fabrication of clinically relevant tissues. Although the supply of oxygen to thick tissue constructs in vitro (such as muscle and bone) can, in part, be addressed by strategies discussed above, combination of these with prevascularization before transplantation is conceptually an attractive approach. Such an approach relies on the seeding of endothelial cells to form primitive capillary-like tubes within the constructs that may improve the vascularization, blood perfusion and survival of the tissue constructs after transplantation [97,98]. Both naturally occurring and 275

10 Deluzio, Seifu & Mequanint synthetic scaffolds that were prevascularized appear to have better host integration and vascularization in animal experiments compared with control scaffolds [98 100]. It is, however, not clear if the in vivo vascular networks formed are nascent with invested pericytes and/or smooth muscle cells. Despite this, it is likely that these emerging data and shared strategies accelerate functional tissue fabrication and host integration. 3D active scaffolds for drug delivery In addition to cell signaling and ligand presentation, porous 3D scaffolds can be fabricated and functionalized for the delivery of biological agents, such as drugs and growth factors. As the broader application of scaffolds for drug and cell delivery is beyond the scope of the present manuscript, the reader is referred to a recent review on this subject [101]. Here, only recent relevant advances in scaffold-mediated drug delivery strategies are reviewed. The conventional wisdom for controlling the delivery of drugs involves either modifying the rate of scaffold degradation or swelling behavior of the drugloaded scaffolds. The drug release occurs by diffusion, often resulting in profiles with an initial unwanted burst effect followed by a slow release of the remaining drug. Additional drawbacks of this method are that the release cannot be controlled and it cannot be applied to structurally sustainable scaffolds (i.e., to biostable scaffolds) [102,103]. Core-shell laminated 3D scaffolds fabricated from hydrophobic PCL and hydrophilic poly(ethylene oxide)/drug combination fibers and embedded in the normal 3D PCL scaffold, which was fabricated by a melt-plotting system, appeared to eliminate the burst release problem [103]. In this study, a linear relationship between the thickness of the embedded electrospun mats and the resulting model drug release was observed. It is also possible to alter the release profile by developing a drug-delivery system independent of the polymer degradation rate and to control the initial burst release via manipulation of the physical structure and fabrication technique of a 3D scaffold [104]. This is often achieved by fabricating either a multi-component [104] or core-shell scaffold [105,106]. It is desirable that the drug release rate is independent of both time and concentration in the scaffold. This would imply zero-order kinetics, ensuring a therapeutically relevant steady release of drug over time, minimizing potential fluctuations and side effects. Concurrently, this maximizes the amount of time the drug concentrations remain within the therapeutic window. This is greatly beneficial for applications where long-term release of biological agents is required. However, since the pore size and distribution and hydrophobic properties of the scaffold can also affect the drug release, further research is necessary to develop a fully functional model for clinical applications. Although the aforementioned strategies improved the delivery of drugs, it would be beneficial to be able to trigger and/or regulate the delivery of these molecules using external cues. Active scaffolds that respond to external stimuli such as temperature, ph, enzymes and physical fields have been explored for the purposed of controlled delivery and can be developed via appropriate design and tailoring of porous biomaterials [102]. A smart drug-delivery polymeric system should undergo a complex chain of responses to survive in vivo, deliver the cargo, release the drug into the target cells and match the desired kinetics of the release [107]. Photoswitchable nanoparticles undergoing reversible volume change from 150 to 40 nm upon phototriggering with UV light and allowing repetitive dosing from a single administration have been reported [108]. These particles are thought to provide spatiotemporal control of drug release and enhanced tissue penetration, useful properties in many disease states, including cancer [108]. A stimuliresponsive macroporous ferrogel scaffold consisting of magnetic field responsive particles embedded in polymer gels is viable since both magnetic particles and fields show clinical acceptance and ferrogels have been made biodegradable and injectable [102]. The macroporous ferrogel fabricated from alginate with a homogeneous distribution of embedded iron oxide nano particles, upon the application of a magnetic field, undergoes gel deformation, causing water to flow through the interconnected pores and triggering the release of biological agents by diffusion. Removal of the magnetic field causes the gel to return to its original configuration as water is reabsorbed from the surroundings. Following a model drug release study, the release profile of human dermal fibroblasts seeded onto the gel, which was modified with RGD peptide to provide integrin-mediated cell adhesion, was also investigated. The ability to externally regulate drug delivery with reversible and on-demand distribution could potentially improve the safety and efficiency of numerous treatment processes. In addition, the large pore size and high interconnectivity allows for the delivery of high-molecular-weight molecules, such as proteins, plasmid DNA and cells, in addition to lowmolecular-weight molecules. In this context, ferrogels are ideal as they can be developed in numerous shapes and sizes to be tailored to specific applications [102]. Conclusion & future perspective In this article the authors have attempted to provide a summary of recent research findings on how 3D scaffolds modulate tissue formation, cell signaling, oxy- 276 Pharm. Bioprocess. (2013) 1(3)