Engineering the extracellular matrix for clinical applications: endoderm, mesoderm, and ectoderm.

Biotechnology Journal

Biotechnol. J. 2014, 9

DOI 10.1002/biot.201300120

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Review

Engineering the extracellular matrix for clinical applications: Endoderm, mesoderm, and ectoderm Miguel L. Williams and Sujata K. Bhatia School of Engineering and Applied Sciences, Harvard University, Cambridge, MA, USA

Tissue engineering is rapidly progressing from a research-based discipline to clinical applications. Emerging technologies could be utilized to develop therapeutics for a wide range of diseases, but many are contingent on a cell scaffold that can produce proper tissue ultrastructure. The extracellular matrix, which a cell scaffold simulates, is not merely a foundation for tissue growth but a dynamic participant in cellular crosstalk and organ homeostasis. Cells change their growth rates, recruitment, and differentiation in response to the composition, modulus, and patterning of the substrate on which they reside. Cell scaffolds can regulate these factors through precision design, functionalization, and application. The ideal therapy would utilize highly specialized cell scaffolds to best mimic the tissue of interest. This paper discusses advantages and challenges of optimized cell scaffold design in the endoderm, mesoderm, and ectoderm for clinical applications in tracheal transplant, cardiac regeneration, and skin grafts, respectively.

Received 12 JUN 2013 Revised 09 OCT 2013 Accepted 27 NOV 2013

Keywords: Ectoderm · Endoderm · Extracellular matrix · Mesoderm · Tissue engineering

1 Tissue engineering as a clinical discipline Tissue engineering is the application of biotechnology to the repair or replication of organic structures. Medical science has long sought to replace damaged natural structures with synthetic analogs. Artificial toes have been found in the tombs of Egyptian mummies [1]. The leather and wood prosthetic had flattened base and joint to facilitate walking. The Egyptians were supplementing function with the materials available to them over two millennia ago. Current medicine has refined the resolution of manipulation from the macroscopic to the microscopic. Medical technologies are increasingly able to replace functions rather than replace parts [2], representing a

Correspondence: Dr. Sujata K. Bhatia, School of Engineering and Applied Sciences, Harvard University, 29 Oxford Street, Cambridge, MA 02138, USA E-mail: [email protected] Abbreviations: BMSC, bone marrow-derived stem cells; ECM, extracellular matrix; FDA, Food and Drug Administration; FGF, fibroblast growth factor; PCU, poly[carbonate-urea] urethane; PGA, poly(glycolic acid); POSS, polyhedral oligomeric silsesquioxane; TET, tissue engineered trachea; TGF-β, transforming growth factor-beta; VEGF, vascular epidermal growth factor

© 2014 Wiley-VCH Verlag GmbH & Co. KGaA, Weinheim

movement from prosthetics to truly regenerative medicine. Convergent capabilities in material science, biochemistry, and cellular culture are creating many of the properties and structures found in natural tissues. These engineered tissues are already proving to be of great use in therapies. Regulated design can minimize the risk of pathogenic or immunogenic contamination reducing inflammation and rejection after implantation [3]. Hypo-immunogenic products are especially beneficial to immunocompromised patients, who often require therapeutics the most. Therapies may be customized, incorporating autologous cell populations or built for specific patients for maximum biocompatibility [4]. Engineered tissues could be used to repair organs that traditional therapies are unable to supplement. Perhaps most importantly, many tissue engineered products are scalable, so production can be increased to meet demand. This could greatly reduce the reliance on scant donor organs which have never met demand, and which will supply a diminishing percentage of patients under current projections [5]. Though engineered tissues hold great promise, there are still major hurdles to be overcome prior to widespread clinical application; laboratory success has thus far translated to relatively few clinical therapies [6]. The novel

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properties that make engineered tissues efficacious require extensive testing, adding to the time and resources required for clinical implementation [7]. Protocols must be validated for manufacture including handling, storage, cellular isolation, and recordkeeping. Due to their novel properties, synthetic tissues are often classified as biological drugs, medical devices, or combination products. The Food and Drug Administration (FDA) classifies each product based on the “primary mode of action”; the requirements for testing may only become clear after the application process has begun [8]. Developing a tissue that meets the criteria of proper function, ease of synthesis, cost, and biocompatibility is a challenging prospect. Many products rely on the tissue engineering triad, which consists of a selected cell population, a biomaterial scaffold, and directed signaling, to replicate natural tissue. The cell scaffold integrates the chosen cells into the host tissue while the signaling molecules promote the required cellular processes. Though natural, sometimes autogenous, cells, and signals can be utilized, cell scaffolds by their nature must be allogeneic or exogenous. The complexity of cell scaffolds is both advantageous and unfavorable for tissue engineers. Cell scaffolds can regulate cellular adhesion, host integration, and natural remodeling and turnover. Conversely cells will respond poorly to erroneous or absent signals. Modern engineers must design cell scaffolds to recapitulate the properties of the natural extracellular matrix (ECM) just as the Egyptians modeled their prosthetics on the natural toe. Cell scaffolds are important mediators of FDA approval, cellular engraftment, and biocompatibility. By replicating the properties of natural tissues, the products minimize rejection and encourage the inhabiting cell to exhibit desired properties.

2 Natural properties of the extracellular matrix The ECM is a tissue specific network of interconnected proteins that organizes the tissue ultrastructure and alignment. It is secreted over the course of development but remains a lifelong homeostatic participant. The composition and properties of the ECM affects and is affected by the surrounding cells [9]. Cellular processes can be regulated by intrinsic factors such as the composition, rigidity, topography and ultrastructure, or extrinsic signaling molecules such as growth factors, cytokines, and chemokines [10]. Together, these signals direct cellular differentiation, proliferation, localization, and homeostasis (Fig. 1). The cells in turn influence the state of the ECM through secreted proteases and inhibitors, creating a system of dynamic reciprocity [11]. These mechanisms help direct tissue response as a whole. The importance of the ECM is underscored by its role as a tumor suppressor, stem cell niche regulator, and repair mediator [12–14].

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Figure 1. ECM dynamic reciprocity. The local tissue environment is influenced by dynamic crosstalk between the ECM, cells, and signaling molecules of the extracellular milieu.

The ECM is mostly comprised of polymeric selfassembling fibrous proteins, glycoproteins, and proteoglycans; tissue specificity arises through precise combinations of the subunits [13]. The protein collagen is the foundation of the ultrastructure support system, three polypeptide chains form a triple helix through hydrogen bonds. Adhesion domains are present throughout the 19 different subtypes creating an array of variable integration sites. Though abundant in most tissues the tensile strength conveyed by fibrils in the tendon and cartilage are representative of the family [15]. Glycoproteins often operate as linkers, connecting ECM molecules to cells and disparate polymers together. Fibronectin, e.g., binds cells, collagen, heparin sulfate, and fibrin; its motifs were used to develop the RGD adhesion peptides used in cell scaffolds [16]. The glycoprotein elastin allows moving tissues to be simultaneously plaint and resilient; it is found in tissues such as the lungs and skin [15]. The proteoglycans are polymerized disaccharides with covalently bound core protein; in addition to structural diversity, properties are modified through bound co-receptors, mophogens, and cell adhesion proteins. The proteoglycans are a significant method of ECM crosstalk and signaling, and use hydrophobic domains to trap interstitial fluid between cells [17]. Signaling factors, such as vascular epidermal growth factor (VEGF) or fibroblast growth factor (FGF) may be bound to specific ECM polymers [17]. These signaling factors may directly interact with cells, or complex with other domains such as integrins in order to induce further signaling. It is worth noting that the spatial relationships of complexing receptors influence the overall signal cascade, introducing another layer of regulation [18]. Free floating signaling molecules may also be regulated by the ECM. Cytokines, such as transforming growth factor-beta





(TGF-β) can be bound to the ECM to be released once more upon remodeling or cleavage by matrix metalloproteases [19]. The ECM can also enable activation of TGF-β by binding inhibitors or presenting active sites to cellular receptors [20]. Adhesion molecules such as integrins and cadherins are built into the matrix, facilitating the localized attachment of specific cells [21]. Cellular focal adhesions connect receptors in the ECM to the cellular membrane and cytoskeleton, actively binding and signaling the cell simultaneously. The rate of ECM turnover is also controlled by the concentration of cleavage sites within the polymers. Cellular secreted proteases break down nearby ECM components, creating space for newly synthesized networks. Tissue engineers are just beginning to imbue cell scaffolds with the complex spatial and temporal signals seen in natural ECMs.

3 Cell scaffolds in tissue engineering Cell scaffold requirements are greater for clinical scaffolds and are often validated first in a preclinical setting. The ideal therapy would be maximally biomimetic, as similar to the undamaged tissue as possible. A tissue engineered scaffold’s purpose is threefold, first to adhere and integrate the transplanted cell population, second to facilitate the transition from implant to host tissue, and third to create the appropriate environment to restore normal function. Composition affects all three of these goals. Scaffolds are often grouped into top down or bottom up approaches. Top down approaches use natural products such as decellularized ECMs as scaffolds [22]. Though these techniques utilize biocompatible products they maintain the same reliance on donor tissue as conventional transplantation and risk pathogenic transmission. Scaffolds built from the bottom up can be built from natural substrates such as collagen or synthetic materials such as poly(glycolic acid) (PGA). Natural materials integrate well with host tissue but can be difficult to manipulate and are limited by impaired mechanical stability. Synthetic materials may be more tunable but are also more likely to create adverse responses in the surrounding tissues [9]. Cellular adherence and recruitment result from many factors including porosity and adhesion sites. The optimal porosity is a particularly fine parameter; too large and the cells will be non-cohesive, unable to form a tissue, too dense and the cells will be unable to infiltrate to the scaffold core [23]. Integrated ligands or adhesion motifs can be built into scaffolds to facilitate recruitment [24]. The three dimensional structure of a scaffold is of key importance to the therapeutic efficacy and engraftment. Current tissue engineered products are limited by poor vascularization. Oxygen will only diffuse across a distance of about 100 μm; cells further from oxygen exchange rapidly become necrotic [25]. A viable therapy must allow for



angiogenesis within the engineered tissue at a rate commensurate with growth and anastomosis with host tissues post implantation. Vascularization can be enhanced by growth factors such as VEGF, pre-culture in vivo prior to application, and in vitro bioreactor incubation [26, 27]. In vitro engineered tissues are limited to a cellular thickness of a few millimeters [28]. Though these methods have improved survival and engraftment of clinical tissues, vascularization remains the size limiting factor. To allow for the growth of host tissue and normal functions, scaffolds must maintain lifelong biocompatibility or be replaced by natural ECM. Bioabsorption facilitates the transition from repair to sustained tissue and minimizes immune response through reduced acute inflammation [29]. Degradable biomaterials have been used in products such as arterial stents before being adopted for tissue engineering [30]. Most biodegradable scaffolds however are broken down by stochastic hydrolysis while ECM turnover is mediated primarily by secreted proteases. The chaotic degradation of polymers and varying protease secretion can be challenging to model, leaving both scaffold types vulnerable. Hybrid biomaterials can be used to take advantage of both natural and synthetic substrates. Rates of degradation are usually determined pre-clinically; premature breakdown within clinical products results in tissue prolapse [31]. Advancements in tissue engineering have afforded researchers the ability to replicate some of the properties of natural tissues but no scaffold has successfully recapitulated all of the natural complexities. Cell scaffold customization in tracheal, cardiac, and dermal repair will be discussed here as representatives of the varied properties found within the endoderm, mesoderm, and ectoderm, respectively. These examples expound on the capabilities and challenges of tissue engineered cell scaffolds as a whole.

4 Tracheal replacement and endoderm cell scaffolds The trachea is an air delivery channel comprised of a series of C-shaped cartilage reinforced rings; it creates a passage from the larynx to the branched primary bronchi. The cartilage must be firm enough to avoid compression during lower pressure exhalation and malleable enough to allow for movement during swallowing or coughing. More precisely the convex ring portions must resist compressive forces, and the concave ring portions resist strain during expansion. The natural ECM is comprised of sheets of collagen fibers along the lumen, providing tensile strength, and a proteoglycan core resisting compressive forces. The tensile modulus of tracheal cartilage ranges between 1 and 20 MPa [32]. The luminal side is comprised of a mucosal layer of epithelial and goblet cells which remove foreign debris and prevent infections.

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Tracheal tumors are rare, but often past the point of surgical intervention once identified [33]. Any constriction of the airway is considered a life threatening condition so maintaining proper structure is critical. Mortality rates range between 77 and 100% depending on the location of the tumor [34]. Surgical techniques are limited by a lack of materials with which to replace resected tissue; a mass producible engineered tissue would improve therapeutic options and lower mortality. The trachea is an excellent candidate for tissue engineering as it is only moderately vascularized and maintains similar properties over the patient’s lifetime. The natural cartilage’s collagen matrix is produced and maintained by chondrocytes. Tracheal replacement therapies attempted to provide structural support while chondrocytes augment cartilage production. The first transplants started in 1950, but all of the transplants prompted immune rejection, or necrosed due to limited vascularization [35, 36]. Prosthetic transplants were investigated into the early 1960s using materials such as silicone, Teflon, polyvinyl chloride; none were able maintain long term rigidity while allowing for epithelialization, often resulting in stenosis [37]. Early studies with biodegradable scaffolds such as collagen coated poly-Llactic acid demonstrated good biocompatibility but poor structural stability, leading to collapse and infections [33]. The tissue engineered trachea (TET) must be structurally secure, devoid of allogeneic antigens and able to anastomose with the host vasculature. The tensile and compressive properties must be soft enough to move when influenced by muscles, while resisting the stress of repeated compressions, with surfaces that can support the necessary epithelial and goblet cells. TETs are being applied clinically using cell scaffolds similar to the endogenous cartilage but long-term studies must be performed prior to widespread use. Both bioresorbable and permanent cell scaffolds have been investigated in research. The bioresorbable scaffolds are purposefully designed to be replaced by ECM as transplanted cells proliferate. Initial TET pre-clinical studies made use of seeded fibroblasts and chondrocytes on a hybrid PGA and collagen matrix, this technique resulted in airway stenosis and scarring [36]. Though structurally sound many biomaterials produced insufficient endothelialization and excess inflammation for widespread clinical use [28]. The first successful clinical application of an engineered trachea was in 2008. It was made with a nonbiodegradable propylene Marlex tube surrounded by a collagen sponge, seeded with autologous gingival fibroblasts and adipose stem cells. No discrete signaling molecules were used though adipose stem cells had previously been shown to enhance vascularization and endothelial cell differentiation through secretion of VEGF and other growth factors [38]. These transplants were found to be stable up to two years after implantation with good luminal epithelialization. Though successful the small graft

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could have been achieved using conventional methods [39]. The protocol also required an in vivo pre culture in order to promote adhesion and vascularization. A permanent, full-sized TET scaffold was designed using seeded autologous bone-marrow mononuclear cells adherent to a polyhedral oligomeric silsesquioxane (POSS) poly-[carbonate-urea] urethane (PCU) scaffold in 2011 [4]. Previous studies on POSS-PCU scaffolds, also known as UCL Nano scaffolds, have noted their excellent biocompatibility, chemical stability, and nanostructured features [40]. Positron emission tomography (PET) scans of the patient’s trachea were used to make an airtight seal when implanted. The cell scaffold rigidity was mediated by controlling the ratio between the soft polymer PCU and hard POSS. The tissue was custom designed for the patient and cultured in a bioreactor where autologous mononuclear cells adhered and expanded on the scaffold. TGF-β3, granulocyte colony stimulating factor, and epoetin beta were applied directly to induce the formation of epithelial layers. The POSS-PCU TET has been successfully maintained in the years post-transplant. Though excellent progress, the procedure’s utility is limited due to high cost, lack of FDA approval, and complex protocol. These challenges could certainly be overcome, indeed they could be applied to most novel medicines, but they are indicative of hurdles both before and after successful clinical use. FDA approval of POSS-PCU and a streamlined protocol will surely further the adoption of this tissue.

5 Cardiac regeneration and mesoderm cell scaffolds Cardiac health is one of the most important topics in modern medicine. Coronary artery disease, the buildup of flow-occluding plaque along the arterial wall, is the leading cause of death worldwide [41]. The ischemia resulting from a cardiac blockage, or myocardial infarction, kills a portion of the heart muscle. The damage reduces contractile tissue and the infarcted wall is remodeled into a collagen fibrotic abscess of scar tissue. The adjacent tissue is strained, increasing cardiomyocyte apoptosis and metalloprotease activation. This disrupts the ECM, diminishing overall symmetry and contractile function. The post-infarcted left ventricle has a vulnerable, thinned wall, with fewer myocytes, under greater stress, providing less contractile force. The cardiomyocytes that make up the heart tissue are non-proliferative and have only a minimal repair capacity; it is believed that myocardial infarction damage is irreversible [42]. This sets up a cascade of diminishing function and increasing damage. Heart transplants remain the best treatment option for a failing heart but are severely constrained by donor availability [43]. It is hoped that cardiomyocytes could be supplemented, either though





injected cells, cellular cardiomyoplasty, or the surgical engraftment of a preassembled cardiac patch, tissue cardiomyoplasty [44, 45]. The cardiac ECM is chiefly composed of collagen, elastin, and proteoglycans but the contribution of individual components to the physical properties remains unknown [46]. Collagen distributes contractile force during systole and limits overexpansion during inflow [47]. Arterial elastin resists stretching at low pressures while collagen provides elasticity at higher pressures, but it is unknown if elastin in heart tissue has a similar function [48]. Elastin knockout mice exhibit normal function at birth but die soon after due to hypertrophy. Cellular grafts overexpressing elastin have been shown to prevent damage post-infarction in rats [49, 50]. The ECM bridles the myocytes’ contractile force. Though myocytes occupy 70–75% of the volume, these cells are only 25–30% of the cellular population. Other components such as fibroblasts make up the largest portion of the non-myocyte cells, underscoring importance of the ECM. The cardiac ECM aligns cells, maintaining a tight link between connexins, the junctions that facilitate ion transfer between cells [51]. Proper location of integrins and cadherins are necessary for optimal conductance and, by extension, synchronization. The natural ECM binds myocytes with collagen type I and III fibrils which run parallel to the myocytes, providing elastic reinforcement and support [52]. A second perpendicular network of collagen weaves around the myocytes, binding them to fibroblasts and the basement membrane [53]. Fibronectin also connects collagen with cellular receptors or actin filaments sarcolemma [51]. By


aligning myocytes along the Z line, the ECM propagates the contractile force [54]. Cell scaffolds should have a Young’s modulus similar to that of natural ECMs so that contractile resistance is uniform across the tissue. Adhesion between engineered and natural tissue should be strong and uniform, most biomaterials have properties that are non-conducive to cardiac function. The natural heart ECM has a low Young’s modulus, about a few hundred kPa, most polymers have a high young modulus in the range of MPa or GPa, which makes integration more challenging [2, 55]. Excitation is propagated between cells through extracellular ions, most bio-polymers are not conductive. A perfect cardiac scaffold would allow for the propagation of ions through the interstitial space, materials such as nanowires have been used to improve conductance [56]. Scaffolds have been designed that meet many, but not yet all of the requirements. A proangiogenic scaffold has been designed that satisfies the requirements for signaling, mechanical properties, cellular alignment and engraftment [57]. Cardiac cell sheets have also shown promise for creating layered, dense cardiac tissue (Fig. 2). The ideal engineered heart product would have an even, synchronized, contractile force, identical to the adjacent host tissue. The heart is highly anisotropic, excitation signals are transmitted longitudinally across gap junctions, connexins, and desmosomes. Misalignment disrupts signal propagation while increasing tissue strain [58]. A product that does not maintain sinus rhythm may work against the surrounding tissue [59]. Additionally therapies must align to host tissue while responding to the same environmental cues, functionality must be

Figure 2. Histological analyses of mouse embryonic stem cell-derived layered cardiac cell sheets. Stacked cell sheets were subjected to hematoxylin and eosin (H&E) and azan staining. Left panels are images of monolayered sheets. Right panels are images of triplelayered cell sheets. Scale bars, 50 μm [104].


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maintained over the range of heart rates and exertions. Application is as important as design; in vitro patches require a thoracotomy, and should quickly fasten securely to the heart to minimize surgery times. By contrast, in situ tissue engineered therapies can set during the hour it takes to catheterize a heart; the catalyst and rate of polymerization must be considered before an in situ product can be tested. Both products must consider the vibrations of normal sinus rhythm and hemodynamic shear that they will withstand. The complexity of the heart makes it a very challenging organ to repair. No therapy has yet successfully combined all of the properties necessary for long term proper function. Many human clinical trials have been performed but have not translated into a standard of care. There are several surgical patches currently available, including nylon, polytetrafluoroethylene and glutaraldehyde crosslinked bovine pericardium [60]. However, these products are non-bioresorbable and only provide structural support, delaying heart failure, not stopping it. Cellular cardiomyoplasty techniques have used stem cell populations to induce repair. Clinical trials have shown improvement in the left ventricular ejection fraction and reduce lesion area when >108 autologous bone marrow-derived stem cells (BMSC) were applied within 7 days post infarction [61]. Such applications reduced the rate of secondary infarction and death [62]. These initial studies underscore the potential for stem cell therapies when applied shortly after an infarction, but not all applications of BMSC have been as successful. Other studies have found that BMSC applications only improved the most severely remodeled tissue [63]. These disparate results indicate that the protocol can be just as important as the application itself. Differences in cell isolation, culture, and application can produce drastically affect integration [64]. Only a small fraction of therapeutic cells are maintained in the target tissue, some estimates are as low as 2–5% within the first few hours of application [64]. Unanchored cells are believed to go through anoikis, programmed cell death from lack of ECM adhesion [65]. Because so few cells adhere to the target, some researchers question whether the stem cells are integrating into the damaged tissue, or merely increasing activity due to paracrine signaling [66]. It is hoped that cell scaffolds can improve engraftment by linking the stem cells to the host tissue. Recent work has indicated that human embryonic stem cells are able to synchronize with the host myocardial tissue [67]. In situ polymerizing cell scaffolds have been utilized as a method to improve adhesion in animal models. Fibrin glue has been used as a scaffold and adhesive in rats; animals given both the fibrin and stem cells maintained structure along the infarcted wall better than control groups [66]. PGA and PLA synthetic scaffolds have been used but are limited by degradation kinetics and reduced elasticity compared to natural tissue [68]. Hydrogels con-

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taining therapeutics delivered within an aqueous polymer matrix have been used for cardiac repair. Acellular hydrogels decrease the stress on rodent hearts postmyocardial infarction and improve contractile force of adjacent tissues [69]. Alginate, collagen, matrigel, and fibrogen have all decreased cardiac stress when applied as acellular hydrogel cardiac therapy [70]. Though the mechanisms are still unknown, it is likely that these cell scaffolds work by sending anti-apoptotic signals to cells around the infarction site while recruiting healthy cells. Scaffolds also act as a bulking agent, thickening the infarcted wall and preventing the ECM breakdown that occurs after an infarction [71]. Much of the current cardiac repair research makes use of decellularized ECM or natural ECM components. Decellularized ECM retain growth factors, adhesion proteins, and basement membranes which facilitate recruitment [72]. Protocols must balance the removal of epitopes and the retention of signaling proteins [73]. Engineered whole organ replacements have been performed but are not yet at the clinical stage; reseeding decellularized matrices remains a promising avenues of research [74]. Though some therapies have been found to be beneficial, pre-clinical progress must be made before a conclusive cardiac repair therapy is developed. Current technology is not able to meet the necessary requirements of application, adhesion, alignment, conductance, and elasticity. These points must be addressed before the promise of cardiac regeneration is fully realized.

6 Skin grafts and ectoderm cell scaffolds The three dermal layers of the skin constitute the largest organ of the body, providing structural support while creating a barrier to pathogens. Though the skin is quite regenerative, repair of all three dermal layers is only possible over a small surface area [75]. The epidermis constitutes the outermost layer and is comprised of keratinocytes, forming a barrier from pathogens. A sheet of heparan sulfate fibers with proteoglycans connects the dermis and epidermis, this basement membrane influences the differentiation of dermal cells into keratinocytes through growth factors and cytokines [76]. The dermis provides the elasticity and tensile strength with an ECM of collagen fibrils, elastin fibrils, and microfibrils embedded in proteoglycans; it also houses ECM secreting fibroblasts and macrophages to resist infection [34]. The hypodermis hosts macrophages, blood vessels, fibroblasts, and adipocytes, while connecting the dermal layers to muscle or bones via an elastin network. Grafts and injuries are broken into full and half thickness, depending if they include full or partial dermal layers. Vitaligo, diabetic ulcers, and traumatic burns are poignant examples of conditions that produce full-thickness wounds; even small gaps can instigate serious





health problems. Without an effective barrier, opportunistic pathogens assail the immune system unchecked [77]. Full thickness damage requires a graft which will link the dermis with the internal tissues below [78]. Tissue engineered skin is designed to create a barrier to prevent pathogenic infection while the body initiates repair. Autologous split skin transplant is possible but limited by the surface area of the wound and the patient’s overall health. Cadaverous skin transplant is a possibility but transplant rejection is still a detraction. Tissue engineered dermal layers must integrate into the host structure, be non-immunogenic, provide a temporary barrier, and allow for self-repair throughout the patient’s life. Over 20 tissue engineered skin products are currently on the market [79]. Tissue grafts are grouped into acellular and cellular varieties. Acellular examples include Alloderm, introduced in 1994, and Integra’s wound matrix system which debuted in 2002. Both are often used as base layer for a later split thickness graft. Alloderm is created from decellularized dermal tissue while Integra is based on bovine collagen and shark chondroitin sulfate with a temporary silicon membrane to prevent desiccation [80, 81]. As the skin heals, the silicon barrier is moved and an autologous epidermal graft is added. Matrix proteins stabilize unbound cells reducing inflammatory signaling, anoikis, and scar formation [82]. The treatment has been shown to recruit fibroblasts and facilitate revascularization of the dermal layer prior to repair or further transplantation [83]. Exogenous matrices will be remodeled during turnover due to the conserved nature of ECM proteins [84]. Epidermal keratinocytes have been applied therapeutically since 1981 and are now being applied as aerosols [85]. Dermagraft delivers neonatal foreskin fibroblasts seeded within a polyglactin mesh, it also contains ECM proteins such as collagen to assist with dermal repair [86]. Epicel utilizes a similar model with autologous keratinocytes. While an effective barrier for epidermal wounds, the scaffold only supports keratinocytes a few weeks, and fails to address wounds below the epidermal layer; multi-layer products are much closer to full thickness grafts. Apligraf, introduced in 2000, is the only FDA approved dermal skin substitute which delivers fibroblasts. The fibroblasts are seeded within type 1 collagen with a layer of differentiated keratinocytes [87]. The tissue contains growth factors and cytokines and is a solid barrier against pathogens and physical perturbation. When Apligraf was applied to patients with venous ulcers, the patients exhibited better wound closure than did control groups [88]. Interestingly deoxyribonucleic acid (DNA) from Apligraf cells was not found past 4 weeks when applied to patients, yet the wound closure was improved when compared to controls. This may suggest that Apligraf’s benefits are not conferred by cells [87]. Many of these partial thickness products are applied in tandem or with a transplant [89].



Figure 3. Stained paraffin section of a MyDerm bilayer engineered tissue. Autologous keratinocytes and fibroblasts are seeded upon a fibrin scaffold. This product has demonstrated excellent engraftment and smooth, even distributions across both layers [91].

Pathogenic transmission through naturally derived matrices or allogeneous cell populations is still a concern; as a result some therapies have designed protocols using autologous cell populations. Two treatments utilize autologous cell populations, PolyActive developed in 2004 and TissueTech’s Autograft System introduced in 2005 [90]. Both expand autologous cell populations and incorporate these cells into the graft. PolyActive utilizes a polyethylene oxide terephthalate and a non-bioabsorbable polybutylene terephthalate component which must be removed during treatment. The keratinocytes also take several weeks to grow to confluence, which precludes applications for burn treatments. These treatments have made the leap into clinical work and future products will continue the advancement. There are several full thickness preclinical tissues. MyDerm is a complete bilayer derived from autologous sources first reported in 2007. Keratinocytes and fibroblasts are seeded on a fibrin cell scaffold derived from the patient’s blood plasma [91]. A layer of silk along the epidermal side helps maintain graft cohesion. Grafts exhibit a smooth transition between skin layers, similar to natural tissues with excellent host integration (Fig. 3). Permaderm (also known as the Cincinnati Skin Substitute) is a second full thickness replacement, introducing autologous dermal and epidermal layers on a bovine matrix. Though promising, it has not finished clinical trials despite being in development since the 1990s [92, 93]. These preclinical steps are impressive but the autologous pre-culture limits the applications for burn victims, and broader methodologies are still required. Integration of other cell types such as hair, sweat glands, and melanocytes will better mimic natural skin. This will help overcome the limitations of current products including high cost and small surface area [89]. Many of the grafts are brittle and challenging for surgeons to implant [94]. Technologies such as scaffolds which transfect recombinant VEGF upon engraftment will improve host integration by creating larger, thicker grafts [95]. Though current

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products are imperfect, they are clear indications of the potential of tissue engineered skin. Future advancements will move synthetic skin to clinical applications of all dermal layers, on demand.

Sujata K. Bhatia, MD, PhD, PE is a physician, bioengineer, and professionally licensed chemical engineer who serves on the biomedical engineering

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With the challenges discussed here, it is no surprise that very few of the developed tissue engineering products progress to clinical use [96]. These scaffolds have made initial forays into cellular response, transitioning from structural foundation to homeostatic director, but they also highlight the limits of current technologies. Both full thickness skin grafts and tracheal cartilage are only minimally vascularized or immune responsive. Vascularization remains a limiting factor, particular in for the highly metabolic cardiac tissue. Methodologies that improve angiogenesis both in situ and in vitro are a sine que non for many future therapies. Clinical methodologies must innovate just as research capabilities do. The large investments required for FDA approval of novel biomaterials limits utilization of products such as POSS-PCU. With continued advancements in chemical manipulations at the nanoscale, increasingly biomimetic cell scaffolds will better mediate cellular functions. Biocompatibility will be improved through new materials such as DNA origami and synthetic peptides [97, 98]. Precision signaling will allow for high resolution control of cellular adhesion and interactions through engineered integrins and precision applications [99, 100]. Integrated cleavage sites will help control remodeling, facilitating natural turnover, and repair [101]. Scaffolds will also influence their environment by both cellular signaling and drug elution, utilizing physical and temporal cues more closely modeling the dynamic ECM [102, 103]. These initial clinical treatments are just the start of tissue engineering’s foray into clinical treatments. By studying the success and setbacks of these clinical applications, tissue engineers can better design future scaffolds.

The authors are grateful to the Harvard School of Engineering and Applied Sciences. The authors declare no conflicts of interest.

8 References [1] Nerlich, A., Zink, A., Szeimies, U., Hagedorn, H., Ancient Egyptian prosthesis of the big toe. Lancet 2000, 356, 2176–2179. [2] Lakshmanan, R., Krishnan, U. M., Sethuraman, S., Living cardiac patch: The elixir for cardiac regeneration. Expert Opin. Biol. Ther. 2012, 12, 1623–1640.

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faculty at Harvard University. She is

Conclusions, future research, and applications

Assistant Director for Undergraduate Studies in Biomedical Engineering at Harvard; she is the academic advisor for all Harvard undergraduate students in bioengineering and biomedical engineering. She is a Lecturer on Biomedical Engineering. In addition, she is an Associate of the Harvard Kennedy School of Government for the Science, Technology, and Globalization Project. She is a faculty member in the Harvard Kennedy School Executive Education program on Innovation for Economic Development.

Miguel Williams is a degree candidate at Harvard University Extension School pursuing a Biotechnology Masters of Liberal Arts, focusing on Bioengineering and Nanotechnology. He will be performing research for his thesis to develop neuronal hydrogels for the treatment of secondary brain injury. Miguel is currently a Research Assistant at RainDance Technologies investigating droplet based digital PCR. Previous work has included epigenetic research within the Freitag laboratory of Oregon State University and studying the organelle acidification of professional phagocytes within the Soybel laboratory at Brigham and Women’s Hospital.

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Isolation of a homogeneous morphogenetic factor, inducing mesoderm and endoderm derived tissues in Triturus ectoderm.

The extracellular matrix: Structure, composition, age-related differences, tools for analysis and applications for tissue engineering.

Extracellular matrix microarrays to study inductive signaling for endoderm specification.

Therapeutic Applications of Extracellular Matrix.

Regeneration of endoderm by ectoderm isolated from mouse blastocysts.

Gut endoderm takes flight from the wings of mesoderm.

Extracellular Matrix Revisited: Roles in Tissue Engineering.

Mechanocompatible Polymer-Extracellular-Matrix Composites for Vascular Tissue Engineering.

Cartilaginous extracellular matrix-modified chitosan hydrogels for cartilage tissue engineering.

Extracellular Matrix-Based Biohybrid Materials for Engineering Compliant, Matrix-Dense Tissues.

[Dysgenesis of the neural crest, ectoderm, mesoderm and fetal alcohol syndrome].

Xwnt-8 modifies the character of mesoderm induced by bFGF in isolated Xenopus ectoderm.

Extracellular matrix production in vitro in cartilage tissue engineering.

Erratum: Extracellular Matrix Revisited: Roles in Tissue Engineering.

Engineering TGF-β superfamily ligands for clinical applications.

Effect of Urea and Thiourea on Generation of Xenogeneic Extracellular Matrix Scaffolds for Tissue Engineering.

Abstract 94: decellularized porcine stomach extracellular matrix for tissue engineering and repair.

A Retinoic Acid-Hedgehog Cascade Coordinates Mesoderm-Inducing Signals and Endoderm Competence during Lung Specification.

Origin of cells giving rise to mesoderm and endoderm in chick embryo.

Cardiac extracellular matrix-fibrin hybrid scaffolds with tunable properties for cardiovascular tissue engineering.

Feasibility of autologous bone marrow mesenchymal stem cell-derived extracellular matrix scaffold for cartilage tissue engineering.

Tissue Engineering for Pediatric Applications.

Changes in glycosaminoglycan structure on differentiation of human embryonic stem cells towards mesoderm and endoderm lineages.

Design of an osteoinductive extracellular fibronectin matrix protein for bone tissue engineering.