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Pharm. Bioprocess. (2013) 1(3), 267–281

26710.4155/PBP.13.21 © 2013 Future Science Ltd ISSN 2048-9145

Review

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.

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 ulti-mate 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 ulti-mate objective of tissue engineering is to develop responsive living tissues with proper-ties 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 sur-prisingly, 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 pro-vides long-term support is fabricated [11–14]. In the case of biodegradable scaffolds, cells

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

Tierney GB Deluzio1, Dawit G Seifu1, Kibret Mequanint*1,2

1Department of Chemical & Biochemical Engineering, The University of Western Ontario, London, Ontario, N6A 5B9, Canada 2Graduate Program of Biomedical Engineering, The University of Western Ontario, Canada*Author for correspondence: Tel.: +1 519 661 2111 Ext. 88573 Fax: +1 519 661 3498 E-mail: [email protected]

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268 future science groupPharm. Bioprocess. (2013) 1(3)

will remodel the scaffold with their own ECM proteins creating the intended tissue without compromis-ing tissue structural integrity. This, 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: biocompatibil-ity, 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 pro-vides 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], chi-tosan [20] and hyaluronic acid [21] have the advantage of being generally nontoxic, in addition to providing biological cues to promote cell attachment and prolif-eration. 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 scaf-fold fabrication also significantly impacts the physio-chemical properties of the resulting tissue engineered construct. Other aspects, including reproducibility and cost–effectiveness, should also be considered when selecting materials and fabrication techniques. Com-mon 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 engineer-

ing 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 fabricationOne obvious and extensively studied area of 3D scaf-fold application is their use as templates for the engi-neering of a variety of tissues including, but not limited to, bone, cartilage, cardiovascular, nerve, skin, blad-der 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 prolifera-tion 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 stud-ies have demonstrated the strategy of scaffold-guided tissue engineering where full scale and functional tis-sues have been developed [9,26–28]. In this section, the authors have selected specific examples to demonstrate the clinical relevance of scaffold-guided engineered tis-sues 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 effec-tively as a result of numerous medical conditions) [29]. A demonstration of autologous hollow organ recon-struction 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 polylac-tic 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 lumi-nal 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 func-tionally 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 observa-tion of upper tract obstruction, lithogenesis, encrusta-tion or other abnormalities. On a more practical clini-cal 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, struc-tural architecture and phenotype after a 4-year follow-up [31]. Although this is a first step towards the goal of

Key Term

Neo-tissue: Tissue that grows to eventually replace the scaffold template so that only the newly formed tissue remains.

Deluzio, Seifu & Mequanint

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future science group www.future-science.com 269

3D scaffolds in tissue engineering & regenerative medicine

engineering fully functional bladders, such engineered tissues show great promise in urologic tissue regenera-tion 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 pulmo-nary sling [32]. In fact, there are reported cases where a biologic scaffold composed of allogeneic extracel-lular scaffolds derived from heterologous sources and recellularized with autologous stem cells or differen-tiated cells are preferred over synthetic materials [33], as demonstrated by earlier clinical transplantations of tissue-engineered airways [34,35]. Furthermore, for tra-cheal tissue engineering, scaffolds with air- and liquid-

tight seals as well as adequate structural support (lat-eral rigidity and longitudinal flexibility) to maintain airway patency and allow rapid epithelialization are required [33]. Such directional variation on mechani-cal 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, suggest-ing 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 for-mation were evident. The scaffold eventually recellu-larized into a functional tissue containing site-specific

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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.

Review


cells and integrated into the adjacent tissue. The patient became asymp-tomatic, tumor-free and the ability to breathe normally was restored [36].

The use of 3D scaffolds has also been successfully demonstrated to fabricate clinically relevant vas-cular substitutes. In view of this, the first clinical application of a

scaffold-guided tissue-engineered vascular construct was reported in 2001 for a pediatric patient with con-genital 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 car-diac failure deaths reported [38,39]. The success of this procedure can be partially attributed to the relatively low-pressure environment (20–30 mmHg during sys-tole) found in pulmonary circulation, which is less demanding than pressures found in coronary arteries (100–140 mm Hg during systole). Despite the urgent need for an engineered vascular tissue substitute, suc-cess is still limited [40]. While engineering approaches for other tissue substitutes can rely on in vivo remodel-ing to approach functionality with time, tissue-engi-neered 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 sub-stitute 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 arterio-venous shunt model due to the lack of suitable vein for hemodialysis and the deplorable efficacy of syn-thetic 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]. Notwithstand-ing 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 inter-ventions 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 spe-cies-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 indispens-able to our understanding of tissue morphogenesis and function in physiological and pathological states, they do not accurately replicate the 3D microenviron-ment of human tissues [15]. For example, 2D cultur-ing 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 dif-ference [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 vas-cular 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 therapeu-tic 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 ani-mal experiments, and has the potential to accelerate

Key Term

Extracardiac cavopulmonary connections: Process by which venous blood flows to the lungs, bypassing the right ventricle. This procedure is often performed to repair tricuspid atresia or for pediatric patients with congenital heart problem.


Review



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 athero-sclerosis 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 engi-neered 3D tissues attract the most media attention, it is evident that engineered tissue models can serve as plat-forms for tightly controlled, high-content screening of drugs and for pharmacodynamic analyzes.

3D scaffolds as ECM-mimicking microenvironments for signal transduction & gene regulation studiesIn tissue engineering and regenerative medicine strate-gies, signals that are stimulated when cells are seeded to scaffolds provide important information on early cellular events. In addition to the chemical composi-tion, cell signal transduction has been predicted to be highly dependent on the dimensionality and architec-ture of scaffolds [56,57]. The merits of culturing cells in 3D rather than on 2D flat surfaces have been acceler-ated 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 sig-naling 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 sig-naling (Figure 2). The first approach is the utilization of cell-derived 3D matrices whereby fibroblasts are cul-tured in a confluent state to produce the ECM as struc-tural 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 fibril-lar 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 com-

pared 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. Further-more, tyrosine phosphorylation of FAK at residue 397 was poor in cells cultured on 3D matrices while acti-vation of the MAPk ERK1/2 was enhanced. Despite unchanged activation levels of Rho and Cdc42, Rac activation was considerably low. Given that the cell-derived 3D matrices are fibrillar and porous, it is rea-sonable for these matrices to be more pliable than com-pressed 2D rigid films of the same ECM components (Figure 2C). This may, in turn, explain the observed sig-naling 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 syn-thetic 3D scaffold. Using a genome-wide gene expres-sion analysis of osteoblast-like cells grown in the 3D col-lagen 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 miniaturiz-ing scaffold-guided engineered tissues and associated physiological assay systems are postulated to acceler-ate 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 vascu-lar 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. Col-lagen Type I expression was also higher in 3D suggest-ing 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

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2D surfaces. As exemplified by the above studies, dif-ferential signaling and gene expression are not limited

to vascular cells. Several groups have reported preliminary evidence that gene expression of different cells cul-tured in 3D more closely parallels the in vivo situation [64–66]. Gene expres-sion profiling experiments in various cell types clearly demonstrated close correlation between engineered and native tissues in tumor cells [64], ten-don [67], bone [61] and skin [68]. Col-lectively, these studies point to engi-neered 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 tech-niques, a number of synthetic 3D scaffolds are frequently used to culture different cell types for signaling stud-ies. 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 cyto-skeletal 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 downregu-lated 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

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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.


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appear to be the case. Data supporting this argument comes from studies of VSMC cultured on 3D polyure-thane scaffolds where it was clearly demonstrated that Cdc42, which is a Rho GTPase like Rac, was undetect-able [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 sug-gest that signaling cascades are activated differently in 3D than their 2D counterparts. Although more stud-ies 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 scaffoldsAs 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 addi-tion 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 inter-actions. Biomimetic scaffolds offer a means to mediate and enhance cell-material interactions for tissue engi-neering, and can be fabricated by modifying (either chemically or physically) synthetic polymers with bio-active peptides, growth factors and other biomolecules. Alterations in the bulk chemistry of the scaffold mate-rial 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 bio-active molecules to 3D scaffolds overcomes this prob-lem; however, it may be difficult to chemically attach the desired molecule, especially without affecting its biological activity [70]. Cellular interaction with adhe-sion proteins in the ECM influences the morphology, motility, gene expression and ultimately the survival of cells [71]. Since the discovery of the essential cell adhe-

sion 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-termi-nus 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 deg-radation, thereby limiting their long-term use [75]. Pro-teins (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 effec-tive biological activity of biomimetic scaffolds. Small peptides, on the other hand, tend to have higher sta-bility towards heat sterilization, pH variation, storage and are easier to characterize and are cost-effective [75]. Furthermore, since peptides occupy less space than pro-teins, peptides can be packed at higher densities on the modified scaffold surface [75]. Still, a single protein con-tains 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 biomateri-als 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 gluaralde-hyde 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 ligand-conjugated scaffolds focused on either RGD or fibro-nectin conjugation with the primary goal of achieving cell adhesion to 3D scaffolds. However, as advances are

Review


made in 3D cell signaling, conjuga-tion 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 tis-sue 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 function-alized scaffolds would allow for enhanced therapeutic strategies with the ability to control cell signaling and gene expression. In addition, development of bioengi-neered surfaces with immobilized Notch ligands would provide improved methods for studying the Notch sig-naling 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 scaf-folds. 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 photoelec-tron spectroscopy and ELISA were used to confirm the presence of Jagged-1 on the surface at varying concen-trations. To test the bioactivity, HL-60 leukemic cell line was cultured on the functionalized constructs and the resulting gene expression, specifically the activa-tion 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 incorpora-tion. Stem cells derived from blood from bone mar-row 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 hemotopoi-etic stem cells through the progression of differential stage-specific surface marker expressions. Such func-tionalized scaffold could, therefore, be applied as an improved tool to investigate human stem cell biology on the molecular and tissue level. It should be, how-ever, noted that experimental conditions including ligand density may lead to variations in cell activities.

3D scaffolds as oxygen reservoirs in tissue engineeringThe 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:

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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 deliv-ery 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 con-centration 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 emul-sion is facilitated by its increased surface area. The dis-advantage 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

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.


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decomposes in water to form calcium hydroxide and hydrogen peroxide, which further decomposes to oxy-gen and water, was embedded into PLGA scaffolds [94]. These porous 3D constructs were prepared via particu-late 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 scaf-fold maintained a highly porous and open pore struc-ture. 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 cul-tured in a hypoxic environment (1% oxygen, 5% car-bon 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 poly-caprolactone (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 nega-tive 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. Inclu-sion of catalase or ascorbic acid to protect against oxidative stress and remove potentially harmful by-products 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 con-sumed. 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 flu-orinated with 1H,1H,2H,2H-perfluorodecyltriethoxysi-lane 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

4Cl 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 distribu-tion 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 infiltra-tion 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 scaf-folds, the infiltration depth was double on the FZ-modified 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 preced-ing cited studies have provided platforms to incorpo-rate 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 quan-tities until the engineered tissue matures. Develop-ment 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 transplanta-tion is conceptually an attractive approach. Such an approach relies on the seeding of endothelial cells to form primitive capillary-like tubes within the con-structs that may improve the vascularization, blood perfusion and survival of the tissue constructs after transplantation [97,98]. Both naturally occurring and

Review


synthetic scaffolds that were prevascularized appear to have better host integration and vascularization in animal experiments compared with control scaf-folds [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 strate-gies accelerate functional tissue fabrication and host integration.

3D active scaffolds for drug deliveryIn addition to cell signaling and ligand presentation, porous 3D scaffolds can be fabricated and functional-ized for the delivery of biological agents, such as drugs and growth factors. As the broader application of scaf-folds 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 deliv-ery of drugs involves either modifying the rate of scaffold degradation or swelling behavior of the drug-loaded scaffolds. The drug release occurs by diffusion, often resulting in profiles with an initial unwanted burst effect followed by a slow release of the remain-ing 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 hydro-philic 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 sys-tem 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 fabricat-ing either a multi-component [104] or core-shell scaf-fold [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, ensur-ing 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 thera-peutic window. This is greatly beneficial for applica-tions where long-term release of biological agents is required. However, since the pore size and distribu-

tion and hydrophobic properties of the scaffold can also affect the drug release, further research is neces-sary 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 appropri-ate 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]. Photoswitch-able nanoparticles undergoing reversible volume change from 150 to 40 nm upon phototriggering with UV light and allowing repetitive dosing from a single administra-tion 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 stimuli-responsive macroporous ferrogel scaffold consisting of magnetic field responsive particles embedded in poly-mer 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 inter-connected pores and triggering the release of biologi-cal 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 modi-fied 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 low-molecular-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 perspectiveIn this article the authors have attempted to provide a summary of recent research findings on how 3D scaf-folds modulate tissue formation, cell signaling, oxy-


Review



gen mass transport and therapeutic agent delivery in the rapidly advancing fields of tissue engineering and regenerative medicine. Due to the rapid advances made in this field and the focus of this manuscript, it was not possible to cover all aspects; however, every effort was made to ensure that relevant seminal works and signifi-cant research findings are included with minimal bias. The past 20 years saw a staggering number of publica-tions on 3D scaffolds for tissue engineering applica-tions, with the majority focused on new bio materials synthesis, scaffold fabrication and remodeling in culture conditions. It is only recently that molecular mechanistic studies in 3D cultures have emerged.

It is the authors’ belief that the next decade will bring a significant leap in elucidating the molecular mecha-nisms by which 3D scaffolds modulate tissue forma-tion. Another future direction of equal importance is the need to optimally prevascularize clinically relevant tissue constructs for accelerated perfusion and host integration. In order to translate engineered constructs to the clinic, safety and efficacy standards need to be provided by regulatory agencies such as the US FDA.

In the interim, however, successfully engineered tis-sues will continue to serve as tools for predictive tissue models and drug delivery and discovery. Ultimately, the recent progress in these fields needs to be com-bined to develop 3D scaffolds that would respond to the ever-increasing complex interplay between cells and materials.

AcknowledgementsCurrent and former laboratory members of K Mequanint’s laboratory are acknowledged.

Financial & competing interests disclosureThe authors acknowledge the financial support provided by the Natural Sciences and Engineering Research Council Canada, and the Heart and Stroke Foundation of Canada. The authors have no other relevant affiliations or financial involvement with any orga-nization or entity with a financial interest in or financial conflict with the subject matter or materials discussed in the manuscript apart from those disclosed.

No writing assistance was utilized in the production of this manuscript.

Executive summary

Background » As its central tenet, tissue engineering and regenerative medicine involves the in vitro seeding and cultivation

of cells to 3D biodegradable scaffolds that are capable of providing both mechanical and biological cues. » 3D scaffolds could be fabricated from naturally occurring and synthetic materials using a number of strategies.

3D scaffolds as templates for clinically relevant & model tissue fabrication » Although not widespread, engineered tissues have been clinically utilized for selected patients where

standardized interventions are not amicable. » With relevant regulatory approvals, the clinical prospect of engineered tissues remains promising. » Engineered tissue models can serve as platforms for tightly controlled, high-content screening of drugs and

for pharmacodynamic analysis.3D scaffolds as extracellular matrix mimicking microenvironments for signal transduction & gene regulation studies » Signaling events in cells seeded on 3D matrices is significantly different from 2D surfaces, which may facilitate

in vitro data correlation with in vivo environments. » The pliability of 3D scaffolds (both in naturally occurring and in synthetic scaffolds) plays a role in preferentially

activating one signaling event over another.Ligand-functionalized biomimetic scaffolds » 3D scaffolds can be modified by conjugating both short chain peptides and full length proteins for specific

cell–material interactions. » Most studies focused on fibronectin and RGD peptide, however Notch signaling molecules can potentially be

conjugated. 3D scaffolds as oxygen reservoirs in tissue engineering » Novel scaffold fabrication strategies allows the incorporation of oxygen vectors to improve oxygen delivery » Fluorinated molecules and calcium peroxide have a potential to facilitate oxygen delivery when embedded into

the scaffolds.3D active scaffolds for drug delivery » Although it is a recent endeavor, the use of fibrous scaffolds to deliver therapeutic agents is a viable strategy

and could have an impact in future drug delivery.Conclusion & future perspective » The combined advances in both cell biology and 3D scaffolds along with appropriate regulatory standards will

accelerate the widespread clinical use of engineered tissues.

Review


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