ReviewTissue Engineering

Engineering Complex Tissues

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Science Translational Medicine  14 Nov 2012:
Vol. 4, Issue 160, pp. 160rv12
DOI: 10.1126/scitranslmed.3004890


Tissue engineering has emerged at the intersection of numerous disciplines to meet a global clinical need for technologies to promote the regeneration of functional living tissues and organs. The complexity of many tissues and organs, coupled with confounding factors that may be associated with the injury or disease underlying the need for repair, is a challenge to traditional engineering approaches. Biomaterials, cells, and other factors are needed to design these constructs, but not all tissues are created equal. Flat tissues (skin); tubular structures (urethra); hollow, nontubular, viscus organs (vagina); and complex solid organs (liver) all present unique challenges in tissue engineering. This review highlights advances in tissue engineering technologies to enable regeneration of complex tissues and organs and to discuss how such innovative, engineered tissues can affect the clinic.


A tremendous clinical need exists for the development of technologies to facilitate the regeneration of injured or diseased tissues and organs. The unrelenting prevalence of trauma, congenital defects, and diseases such as cancer drives the demand, which becomes increasingly urgent as the global population expands and ages. A wide variety of tissues and organs would benefit from engineering-based repair or regeneration, from musculoskeletal tissues, such as bone and cartilage, to entire organs, including the bladder and liver. The field of tissue engineering is at the interface of bioengineering, materials science, chemistry, biology, and medicine, poised to meet these unmet clinical needs through the development of new technologies and refinement of existing ones.

Increasing levels of complexity in the tissues or organs targeted for repair generally necessitate a concomitant increase in the complexity of the associated tissue engineering approach. For example, solid organs, such as the kidney, would require several essential structures to restore function, whereas tubular hollow organs, such as the urethra, are more easily recreated from basic cells and materials (1). Similarly, complexity can be found at the interfaces between tissues, such as the transition from cartilage to bone in the osteochondral interface in articulating joints. Such interfaces are receiving increasing attention as targets for repair, given the prevalence of injuries affecting them (2, 3). Indeed, a complex tissue injury or defect may involve multiple tissue types, may be associated with compromised vascularity, or may be at risk for infection.

Regardless of the complexity of the target for repair, tissue engineering strategies generally involve the application of combinations of biomaterials, cells, and biologically active factors to effect tissue formation. This process can involve de novo growth in tissue culture (in vitro, ex vivo) or induction of tissue regeneration in vivo at sites or under conditions where it otherwise would not occur. Increasing emphasis is being placed on the development of tissue engineering approaches within the context of the injury or disease underlying the defect. For example, traumatic injuries to the extremities may involve multiple tissue types (bone, muscle, vasculature, lymphatics, nerve), and biomaterial-based approaches for regeneration are being developed and evaluated using preclinical composite tissue defect models (4). The present review will focus on advances in tissue engineering and regenerative medicine that may enable the repair of tissues with high complexity, while highlighting bottlenecks to clinical translation of these technologies.

Tissue Engineering Scaffolds

Biomaterials can provide a three-dimensional (3D) structure to support tissue growth. These scaffolds define and maintain the space in which the target tissue will form and can be tailored to support the attachment and proliferation of cells to effect the desired tissue formation (5). Ideally, a scaffold should serve as a transient structure that will degrade or resorb with time, such that it is replaced with the tissue of interest. Advances in biomaterials science combined with increasing knowledge of extracellular matrix (ECM) biology and the role of environmental factors in tissue formation have led to the development of scaffolds tailored to provide appropriate structural support and, in some cases, biological and mechanical cues to promote tissue regeneration in vivo (69). Moreover, scaffold biomaterials can be modified to present biologically active signals, including cell-adhesion peptides and growth factors, to facilitate cell attachment and to direct tissue formation (1012). In some instances, the scaffolds depend entirely on the migration of cells from the body into the defect for tissue formation to occur, whereas other approaches leverage the scaffolds for the transplantation of cell populations to supplement the body. In either case, tissue engineering scaffolds seek to mimic key elements of the ECM and local microenvironment to support and perhaps induce tissue formation.

Naturally derived polymeric materials, including polypeptides (for example, collagen) and polysaccharides (for example, hyaluronic acid), have been explored extensively in the development of tissue engineering scaffolds for applications ranging from cartilage repair to functional pancreatic replacements (13). Indeed, a key advantage associated with naturally derived polymers is the general capacity of these materials to support the attachment, proliferation, and differentiation of cells (14, 15). Although naturally derived polymers are typically enzymatically degradable, the kinetics of degradation may not be easily controlled or predicted. The generally weak mechanical strength associated with naturally derived polymers is also a limitation, but it may be improved through the introduction of intermolecular cross-links (16). However, cross-linking may prolong the degradation of the materials (17). A concern with naturally derived polymeric materials is the variability inherent in the production of the materials as well as the potential, albeit small, of the materials to evoke an immune response.

Synthetic polymers present several key advantages relative to naturally derived polymers. Synthetic polymers can be reproducibly manufactured with a wide range of mechanical properties and degradation kinetics to enable the production of scaffolds with properties tailored for a particular application (18). For example, scaffolds comprising poly(lactic-co-glycolic acid) (PLGA) have been investigated for the regeneration of tissues ranging from blood vessels to bone. Many synthetic polymers undergo hydrolytic degradation, which may be more readily predicted and controlled than enzymatic degradation in vivo, given the lack of dependence on local enzyme concentrations. Certain classes of synthetic polymers, such as poly(α-hydroxy esters), produce acidic products upon degradation (19), which may elicit a prolonged inflammatory response (20). Nevertheless, synthetic polymers themselves typically do not carry a risk for inducing an immune response owing to a lack of biologically functional domains. This is also a limitation because synthetic polymers cannot facilitate cell attachment or direct phenotypic expression as a natural polymer would. However, a variety of synthesis techniques have been developed to incorporate biologically active domains into synthetic polymer scaffolds, thereby enabling the production of biomimetic scaffolds with a defined and tunable composition (21).

Materials derived from the native ECM have also been explored as scaffolds. Tissues like the urinary bladder submucosa or the small intestinal submucosa can be processed through mechanical and chemical manipulation to remove the cellular components, yielding a collagen-rich matrix, in a process called “decellularization.” The structures and arrangement of the various ECM proteins in the resulting acellular matrices are largely conserved, which results in a general retention of the mechanical properties of the original tissue (22). Moreover, acellular tissue matrices have been shown to support the ingrowth of cells and tissues in several applications, without inducing a gross immune response (23, 24). Indeed, given the natural origin of the matrices, the materials degrade slowly after implantation and are replaced or remodeled with matrix produced by cells (25). Decellularized matrices may also be processed to form particulates that can be used either alone or in combination with other materials to promote tissue repair (26).

In other cases, synthetic polymeric scaffolds have been fabricated and modified through covalent immobilization of ECM-derived moieties to control presentation of growth factors, promote cell attachment, and enhance directed differentiation of progenitor cell populations (27). Additional methods to introduce an ECM-mimetic coating on scaffolds have been explored, including coating of synthetic polymeric scaffolds with naturally derived polymers (collagen or gelatin) and ceramics (calcium phosphate) for bone tissue engineering applications (28, 29). Loai et al. (30) combined particles of acellular tissue matrices (urinary bladder submucosa) with polymeric materials in the fabrication of scaffolds with biological activity and tunable properties for generation of a vascularized bladder in murine and porcine preclinical models. Alternative approaches have seeded cell populations onto scaffolds and leveraged culture conditions to drive the differentiation of cells and the concomitant production of ECM. Recently, this was demonstrated in the production of bone-like ECM (31, 32).

To support tissue formation, a tissue engineering scaffold must present an interconnected porosity or be capable of resorbing as a function of time to create space for new tissues. Many fabrication techniques have been developed to enable the fabrication of 3D scaffolds with an interconnected porosity, ranging from particulate leaching techniques to electrospinning methods (33, 34). Although traditional methods for scaffold fabrication can enable introduction of interconnected pores with a tunable pore size, control of pore architecture has been a challenge (35). Three-dimensional printing methods have emerged to enable the fabrication of scaffolds with precise control of the architecture throughout the structure (36, 37). Printing techniques have even been used to produce scaffolds with controlled gradients in mechanical properties and gradients of biologically active factors (38). With this technology, scaffolds with spatially controlled properties have been created for the regeneration of complex tissue structures, such as bone and cartilage (3941). However, in some cases, printing of complex scaffolds of dimensions of clinical relevance, such as whole kidneys or livers, may be too time-consuming for widespread application.

Tissue engineering scaffolds should support the attachment and proliferation of cells and the subsequent formation of the tissue of interest. However, scaffold materials alone often lack the biological cues to induce tissue formation. Accordingly, scaffolds are commonly used for the presentation or controlled delivery of biologically active factors to induce tissue regeneration. Growth factors, ranging from angiogenic factors, such as vascular endothelial growth factor, to osteogenic factors, such as bone morphogenetic protein-2, have been incorporated into scaffolds to promote tissue formation (42, 43). Key challenges associated with growth factor delivery in tissue engineering include not only selection of the appropriate factor or combination of factors necessary to induce the desired response but also the dose and spatiotemporal delivery needed for proper tissue development (4446). Another challenge has been the maintenance of the biological activity of the factor, especially once released from the scaffold.

Cells in Engineered Complex Tissues

Scaffolds used in tissue engineering approaches are commonly divided into two general categories, namely, acellular scaffolds, which depend on cells in the recipient to effect tissue formation, and cellular scaffolds, which serve as cell transplantation vehicles. In both cases, the success of a scaffold technology toward achieving tissue growth depends largely on the action of the cells. Accordingly, many current efforts in tissue engineering seek to identify and optimize cell populations that can be leveraged for delivery with a scaffold to promote tissue repair where it otherwise might not occur.

Autologous cell populations have been of great interest for application in tissue engineering approaches because they have minimal risk of rejection. Some early efforts in the field focused on isolating primary cells from a biopsy of the tissue or organ of interest and growing the cells ex vivo for subsequent introduction back into the patient in a tissue engineering therapy. However, a major limitation encountered in this area has been the difficulty in expanding cells to sufficient numbers for clinical application. As an alternative, precursor cells and their necessary culture conditions for tissue engineering have been identified for several tissues and organs. For instance, urothelial cells have been grown and expanded in vivo, but traditionally, the expansion has been limited (47, 48). Methods have been developed in recent years to identify undifferentiated cells within the urothelial cell population and to maintain the undifferentiated state even through the growth phase to obtain sufficient numbers of cells for seeding of scaffolds (47, 49). These methods have enabled the isolation of urothelial cells from a single specimen with dimensions of 1 cm2 and the expansion of these cells over a period of 8 weeks to sufficient numbers to cover the equivalent of a football field (4202 m2) (47).

Although advances in cell culture protocols have allowed for expansion of autologous cells to sufficient numbers for clinical application, expansion of primary cells from some tissues and organs, such as the pancreas, remains a challenge. Additionally, in some cases, tissue engineering strategies rely on autologous cells derived from diseased tissues or organs, which may not yield a sufficient number of normal cells for clinical application. As a result, tissue engineers seek to leverage autologous stem and progenitor cell populations, such as bone marrow–derived mesenchymal stem cells (MSCs) and adipose-derived stem cells (50). Although MSCs have received a great deal of attention in the tissue engineering literature, advances with other adult-derived stem cell populations have generally progressed slowly, owing in part to difficulties associated with maintaining the stem cells in culture or achieving attachment of the cells to scaffolds (51). Nevertheless, some clinical strategies have involved seeding of patient-derived stem and progenitor cells on biomaterial scaffolds and then leveraging the body as a bioreactor for tissue growth. For example, a ceramic scaffold within a titanium mesh was seeded with bone marrow as a source of stem cells and implanted in the latissimus dorsi of a patient to grow a mandibular replacement ectopically (52).

Other types of stem cells have been included in biomaterial scaffolds for tissue engineering applications. These cell-based therapies are beyond the scope of this review, but the reader is referred to (5355). It should be noted that a vascular network is generally needed to support the viability of cells throughout a larger, more complex tissue-engineered construct. Accordingly, a variety of methods have been developed to promote vascularization of tissue-engineered constructs, ranging from functionalization of scaffolds with bioactive factors to development of bioreactor systems to promote vessel formation ex vivo (5658). A detailed discussion of vascularization strategies is provided in (59, 60).

Creating Complex Organs

An expansive toolbox of biomaterial- and cell-based technologies stands ready to contribute to the production of tissue engineering solutions to meet clinical needs. However, immense complexity can be found in the various tissues and organs targeted for replacement. Moreover, the injury or disease driving the need for tissue repair or replacement can add levels of complexity. A common challenge encountered in the development of tissue engineering technologies is the need to repair tissue defects or to regenerate organs that have intricate 3D structures. Furthermore, it is challenging to integrate the regenerating tissue with surrounding tissues and to maintain cell viability in large constructs.

To better understand the structural design of human tissues and organs that regenerative medicine attempts to replicate, it may be helpful to categorize them into four levels according to their increasing complexity: flat tissue structures; tubular structures; hollow, nontubular, viscus structures; and complex solid organs (Fig. 1). Within these levels of complexity, there are several strategies used to achieve restoration of function. We also consider the unmet clinical needs in these areas and the barriers to translation in existing demonstrations.

Fig. 1

Four structural levels of complex tissues and organs. Human tissues and organs can be categorized generally into four levels of structural complexity: flat tissue structures, such as the cornea; tubular structures, such as the trachea; hollow, viscus structures, such as the bladder; and solid organs, such as the kidney. The complexity of a tissue engineering approach generally increases with the structure and metabolic functions of the tissue or organ targeted for repair.


Flat structures

Sheets of cells consisting of multiple layers of predominantly one cell type represent the simplest architectural subtype in the body. This level of tissue complexity is exemplified by the integument system, which represents one of the earliest attempts at culturing autologous cells in vitro for repair purposes (61). The effects of substantial loss of skin surface area are detrimental, as can be seen in burn patients. Traditional treatments, such as skin grafts harvested from unburned portions of the body or allogeneic grafts that provide temporary protection, are the current clinical “gold standard.” However, skin autografts require harvesting healthy tissue, which may not be available in adequate supply in some clinical cases, such as severe burns affecting large surface areas. Likewise, skin allografts present a risk of immunologic rejection and disease transmission.

Accordingly, several technologies are currently being used to engineer adequate skin for human replacement. Normal skin cells are being harvested from the patient in the operating room and are then sprayed over the burn area (62, 63). Patient-derived skin cells are also being expanded and layered ex vivo, with subsequent implantation over the burn area, thereby reducing the donor site morbidity required for burn coverage (64, 65). Although coverage of a burn with a tissue-engineered skin construct can facilitate repair, the size and severity of the burn play an important role in determining the ultimate outcome. Large, full-thickness burns present a greater clinical challenge for repair than small, superficial, partial-thickness burns because blood vessels and regenerative epithelial elements of the dermis are destroyed in full-thickness wounds. Nevertheless, clinical and commercial success has been realized with tissue engineering approaches for functional repair of skin in several applications (66, 67). The functional and cosmetic outcomes, however, may be improved through ongoing efforts to recapitulate more fully with the tissue-engineered constructs the complex strata; vascular, lymphatic, and neural elements; pigment; hair follicles; and secretory glands of natural skin.

The cornea represents another flat tissue structure that has been the target of biomaterial-based tissue engineering approaches for repair in the clinic. It performs a fundamental function in the refraction of light for vision and depends on maintenance of its characteristic transparency for efficacy. The cornea retains its transparency in vivo through maintenance of the shape and organization of a highly aligned collagen matrix and active pumping for continuous removal of aqueous humor from the tissue. A variety of disorders can disrupt proper corneal function, and surgical transplantation of donor corneal tissue has long served as a clinical standard of treatment for such conditions. However, transplantation requires procurement of donor tissue matched to the specific requirements of the recipient. To address this challenge, biomaterial-based tissue engineering approaches have been developed and translated clinically to enable corneal repair without the need for human donor tissue (68).

Tubular structures

Regenerative medicine has been able to successfully replicate many types of tubular structures, including the urethra, trachea, and esophagus, in both animals and humans. In general, these structures consist of two different cell types arranged as sheets of cells. These sheets form into circular, bilayered tissues, which usually serve as means of transporting fluid or air throughout the body. The tubular tissue structures are composed of an inner layer of epithelial or endothelial cells that provide a functional barrier and an outer layer of smooth muscle and connective tissue to provide support. Whereas the single cell–layered skin constructs do not require a complex foundation, tubular structures need to incorporate a matrix of synthetic or naturally derived scaffolding for support.

Tubular, tissue-engineered urethral constructs comprising synthetic, biodegradable poly(glycolic acid) (PGA)/PLGA scaffolds seeded with autologous urethral muscle and epithelial cells were implanted into five patients needing complex urethral reconstruction, and the engineered urethras remained functional over the clinical follow-up period of up to 6 years (69). For blood vessels, autologously derived cells cultured from peripheral vein biopsies have been grown in both biodegradable collagen and synthetic scaffolds and successfully used as pulmonary artery transplants (70). With a different method, vascular access grafts for patients with end-stage renal disease requiring hemodialysis have been engineered and implanted in humans (71). To accomplish this, fibroblasts and endothelial cells were harvested from patients, expanded ex vivo as sheets of cells, and then wrapped around a stainless steel cylinder to allow for fusion. In both situations (70, 71), clinical trials have yielded functional implants. Synthetic materials have also been seeded with cells to create new blood vessels for implantation. For example, human and canine smooth muscle cells were cultured on PGA tubular scaffolds, then treated with detergents to produce acellular vascular grafts capable of long-term storage. These vessels demonstrated patency in both baboon and canine models (72).

Decellularized scaffolds have been used to create tracheas. In animal models, autologous chondrocytes cultured from cartilage biopsies were seeded in biodegradable collagen scaffolds and successfully implanted in the pulmonary tree (73). Autologously derived chondrocytes have been differentiated from bone marrow MSCs, and epithelial cells were isolated from a bronchial mucosa biopsy. The cells were seeded in the decellularized donor trachea and cultured in a bioreactor (74).

Hollow, viscus structures

Like tubular structures, hollow, viscus organs, such as the bladder and vagina, generally consist of an inner layer of epithelial-type cells surrounded by an outer layer of smooth muscle and/or connective tissue to provide functional capacity and to anchor the structure in place. Whereas tubular structures generally serve as conduits for air or fluid, viscus, nontubular organs have wider functional parameters, higher metabolic requirements, and more complex intracellular and inter-organ interactions. Similar to tubular organs, the biofabrication process depends on a scaffold seeded with at least two different cell types. However, the scaffold design is more complex in terms of both its architecture and its predetermined anatomical space limitations, which are often patient-specific. In addition, once the engineered construct is completed, there are special considerations for implanting the engineered construct and for connecting it with other tissues and organs. Regeneration of bladder tissue has been accomplished in patients by using autologously derived urothelial and smooth muscle cells (75). A computed tomography scan was performed on patients before tissue biopsy to determine the size of the organ to be constructed. Thus, the scaffold architecture and size were individualized for each patient. The cells were harvested, cultured in vitro, and seeded onto biodegradable collagen-PGA composite scaffolds, and the cell-scaffold constructs were placed in bioreactors to develop the tissues.

Similarly, a construct for vaginal replacement was synthesized by combining PGA fibers with a coating of PLGA, seeding this scaffold with rabbit epithelial cells, and culturing within a perfusion bioreactor. After implantation as a total vaginal replacement in a rabbit model, the biomaterial constructs yielded organs that were successfully integrated by the host animal and subsequently demonstrated histological characteristics similar to natural tissue after several months of growth (76). The engineered constructs had to be designed with determined specifications that would allow a patent connection with both the uterus superiorly and the introitus opening inferiorly. As a result of these experiments, human clinical trials for vaginal regeneration are under way (COFEPRIS HIM87120BSO).

Solid organs

Solid organs, such as the kidney, heart, pancreas, and liver, have the highest level of tissue complexity. The traditional treatment for end-stage solid organ disease is either temporary supportive treatment with drugs or devices or whole-organ transplantation. Conventional transplantation allows select patients to regain a functional organ, yet it is exceptionally complicated to obtain a histocompatible match that does not require the use of immunosuppressive agents. The ultimate goal of regenerative medicine is to bioengineer and transplant complex, solid organs composed of cells derived from the patient in need. This objective, however, presents an exceedingly difficult and challenging task given the tissue complexity and developmental process of these organs. Complete regeneration of these whole organs requires incorporation of extensive vascular networks to support the viability of cells throughout the organ as well as precise organization of multiple cell types—two challenges traditionally not faced in the biofabrication process of simpler tissues like the skin. Whereas creation of flat, tubular, and hollow viscus organs primarily uses cell-seeded scaffolds, replication of solid organ function must incorporate other methods to be successful because these organs have complex architecture that extends beyond simple layers. To this end, efforts are focused on the development of biomaterial-based approaches that incorporate gradients of growth factors, hybrid composite materials (77), and use 3D printing methods (37).

Patients with end-stage renal disease suffer major medical sequelae secondary to loss of the many physiological duties carried out by the kidneys. With complete renal failure, these patients must undergo mechanical dialysis to replace the waste disposal function of the kidneys and must be closely monitored for electrolyte and acid-base derangements, among many other medical complications. Complete renal functional capacity is not required for survival, and providing sufficient tissue for functional recovery may be an achievable, shorter-term goal. Clonal bovine kidney cells have been used to engineer miniature kidney structures that were implanted in cows and steers and were able to filter blood and secrete dilute urine (78). This was accomplished by expanding renal cells from a cloned bovine metanephros and seeding the cells on collagen-coated cylindrical polycarbonate membranes. These constructs were able to produce a fluid within the physiologic range of bovine pH, urinary glucose, and specific gravity. However, magnesium and calcium were outside normal physiologic ranges. It remains unclear if human physiology could be replicated with this methodology or if magnesium and calcium levels could be corrected to fall within normal ranges. However, if these challenges are met, this approach could be used to treat chronic renal failure as a tissue-engineered, functional kidney substitute in patients.

Various solid organs have been decellularized, followed by attempts at recellularization in vivo in animal models. Decellularization of renal organs can be accomplished, and initial attempts at repopulating the remaining ECM architecture with cells have revealed some organizational capacity of the perfused and seeded cells, as reflected by maintenance of renal ultrastructure (79). In a rabbit model, donor phallus tissue has been decellularized, the ECM structure preserved, and the scaffold seeded ex vivo with both corpora cavernosa penile muscle and endothelial cells (80). The erectile organ was replaced, and the rabbits were able to show successful erection, penetration, copulation, and ejaculation. After mating, sperm was noted in the vaginal vault of the female partners, and reproduction was feasible with viable offspring. In a rodent model, the preserved ECM architecture of a heart, including walls, valves, and blood vessels, was perfused with harvested endothelial cells and then injected with neonatal cardiac cells (81). Macroscopic contractile function was observed. Similarly, livers from animal models, including mice, rabbits, ferrets, and pigs, were decellularized and repopulated with human fetal hepatocytes and human umbilical vein endothelial cells (Fig. 2), resulting in histologically viable cells that secreted albumin and urea—two products of normally functioning liver parenchyma (82). The decellularization process involves using mild detergents, usually infused through the vasculature, which remove the cellular elements but allow the ECM to remain intact (Fig. 2, A to C). Attempts are made to preserve the vascular matrix architecture, from the large vessels to the capillary tree (Fig. 2D). The cells are then reseeded both in the vasculature and in the parenchyma (Fig. 2E), leading to adequate tissue architecture (Fig. 2F) and function (82). Lastly, harvested rat pancreatic islet cells have been seeded on a decellularized pancreas matrix and observed to secrete insulin in vivo in a rat model (83).

Fig. 2

Engineering liver tissue with decellularized donor organs. (A to C) Macroscopic views of a ferret liver before decellularization (A), 20 min after vascular infusion of mild detergents (shows blanching) (B), and 120 min later, demonstrating clear parenchyma, vasculature, and a defined liver capsule (C). (D) Liver visualized by fluorescence microscopy shows the decellularized native vascular tree. (E) A decellularized ferret liver scaffold, with a cell-seeded right lobe 7 days after vascular tree infusion with human endothelial cells and parenchymal infiltration with liver progenitor cells. (F) Fluorescently labeled endothelial cells seeded from the portal vein distributed predominantly in the periportal areas, delineating the typical hexagonal pattern seen in hepatic tissue. All panels are reproduced from (82) with permission.

Looking forward to human studies, the lack of available homologous scaffolding for solid organs represents a limitation. Semi-xenotransplantation, in which the decellularized donor scaffold is an animal source but the cells used to populate it are obtained from the patient, may be a solution. Studies evaluating transmission of animal-based infectious agents and rejection potential have shown promising preliminary results in a vascular graft application of an acellular porcine matrix in a sheep model (84). Furthermore, although still in the preclinical stage, bioprinting may ultimately be able to fabricate solid organs, including the vascular network and functional parenchyma components (37). Continued efforts are focused on reducing the time required for printing scaffolds while retaining sufficient spatial resolution to maintain the scaffold architecture. Despite advances, not enough progress has yet been made to translate tissue-engineered solid organs into the clinical realm.

Clinical challenges

Developing complex, tissue-engineered constructs that suitably recapitulate a tissue or organ, that integrate well with surrounding tissues, and that are appropriately vascularized is the holy grail of tissue engineering. However, when considering the clinic, several aspects complicate translation. Among these considerations is the inevitable variability in health status and intrinsic healing/remodeling potential between patients. For instance, a tissue-engineered heart valve may be designed such that it will resorb and be remodeled in a patient over time, yet variations in remodeling potential and rates of remodeling between patients could lead to inadequate valve function with time. Preclinical animal studies typically are not designed—and, in some cases, are not suitable—to investigate the variability that patient-specific factors such as health status, age, and ethnicity might introduce with respect to remodeling potential and rates between patients in the clinic. Consequently, designing a tissue-engineering approach that can accommodate patient variation may present immense challenges, and biomarkers may be needed to predict their clinical success. Similarly, techniques for minimally invasive monitoring of a tissue engineering technology after application, such as vascularization of a construct or integration with surrounding tissue, may be needed to track performance over time in each recipient.

As highlighted in previous sections, many tissue engineering approaches involve the use of cell populations. However, complexities and variation can be found with respect to the cell populations used in a construct. For instance, the use of autologous cells is generally preferred, yet primary cells harvested from a patient might be obtained from a diseased tissue or organ or yield insufficient cell numbers for the therapy. Furthermore, variation between the density of stem and progenitor cells as well as the regenerative capacity of the cells themselves could lead to considerable differences in outcomes between patients for a given therapy, as has been observed in clinical applications of MSCs for bone repair (85). Allogeneic cells may be harvested from donors and pooled to potentially decrease variability between patients. Nevertheless, the recipient contributes to the potential for the ultimate success of the therapy by virtue of their health status and intrinsic capacity for tissue repair, among other considerations. Moreover, allogeneic cells present the potential for rejection and disease transfer.

The clinical context defining the need for a tissue engineering therapy in a given patient can add complexity to the approach required to address the need. For example, many technologies have been developed with a goal of promoting bone regeneration in bony defects. The context in which the bony defect is present affects the potential efficacy of the tissue engineering approach for repair. A bone defect arising from a traumatic injury may present with concomitant loss of surrounding soft tissue or with bacterial contamination. The regeneration of bone in the defect will be complicated by the absence of surrounding soft tissue and associated vasculature or by the possibility of infection and potentially osteonecrosis (86).

To address such clinical complexities, the field of tissue engineering is investing more in the development of technologies within the context of a more focused clinical application—that is, shifting from the development of technologies to meet general clinical challenges (for example, scaffolds for bone regeneration) to development of technologies to meet the demands of a specific clinical application (for example, scaffolds for alveolar bone regeneration in tooth extraction sockets). Such focused investigation needs to consider at the outset the clinical need for the technology and the commercial and regulatory pathways that may be appropriate to allow for the clinical translation and market viability of the technology.

Regulatory Considerations

Various regulatory considerations can affect the development of tissue engineering technologies, and a clear vision of the appropriate regulatory pathway for approval or clearance of an envisioned product should be established at the onset of the technology development (77). In general, the regulatory complexity associated with a particular tissue engineering product increases with the complexity of the product itself; for example, a synthetic, polymeric scaffold might be regulated as a device, whereas the same scaffold that delivers growth factors and contains cells might be regulated as a combination product. Consequently, the regulatory complexity and associated burden can be minimized through identification and pursuit of the simplest product sufficient to meet the desired clinical need. Additional considerations include the manufacturing consistency associated with a tissue-engineered product (87). As expected, manufacturing is more challenging as the complexity of the product increases, and the cost of the product will be expected to increase accordingly. Moreover, tissue-engineered products may be produced in small lot sizes, which creates a challenge in establishing and maintaining inter-lot consistency (88).

Safety associated with materials used in a tissue-engineered product must be assessed, and a variety of standards are available from organizations such as the American Society for Testing and Materials International and the International Organization for Standardization to facilitate such testing. However, the evaluation of the safety of a material can be convoluted by complex compositions, such as decellularized tissues or ECM constructs created in vitro (88). Further complexity can be associated with the regulatory pathway associated with biodegradable materials and materials presenting or releasing biologically active moieties. Indeed, interactions with the relevant regulatory body should be initiated at the initial stages of technology development to facilitate identification of the appropriate regulatory pathway for the envisioned technology, as well as to guide selection of suitable methods and study designs for preclinical and clinical investigations to support regulatory consideration of the technology.

There are several additional regulatory concerns in the development of biomaterial products that use cells. Inconsistency in cell populations used in a tissue engineering product can complicate the regulatory pathway. To this end, the cell populations should be thoroughly characterized, and reproducible standard methods for cell isolation and culture, if applicable, should be established (8991). Methods to assess and ensure the purity of cell populations, especially with respect to foreign contaminant, infectious agents, and toxins, may also be needed (88). Given these and other considerations regarding the development of products incorporating cells in tissue engineering products, one must carefully weigh the potential benefits of inclusion of cells in the product versus the potential regulatory and logistical manufacturing burdens that may be associated with the use of cells in the technology.

The preclinical evaluation of a tissue-engineered product must be considered on the pathway to the market and clinical application. Various preclinical animal models are available for the evaluation of tissue engineering technologies within a complex clinical context. For example, small-animal models have been developed recently that have controlled composite defects involving multiple tissues, such as bone, muscle, and nerve (4). Other models introduce a contaminating bacterial species to allow for evaluation of tissue engineering technologies within the context of an infected wound bed (92). Nevertheless, some challenges remain regarding preclinical models, including selection of appropriate controls, establishment of suitable end points, and identification of appropriate measures and success criteria. Considerable effort has also been invested in the development of standardized models, which may facilitate broad evaluation and comparison of technologies (93, 94). The participation of representatives from governmental bodies, such as the U.S. National Institutes of Health and the U.S. Food and Drug Administration, in discussions regarding translational models for regenerative medicine underscores the complexity and pressing urgency of the topic (94). Considering the importance of data arising from preclinical and clinical development of a tissue-engineered product in supporting its evaluation by regulatory oversight bodies, consultation with the appropriate regulatory bodies at the onset of product development may facilitate identification of appropriate models and study parameters to mitigate wasted efforts and resources.


Apart from regulatory considerations, commercial considerations are also important in the potential success of a tissue-engineered product. A sufficient market must be available to drive the viability of the product for the clinical application (95). Indeed, the intended use must be clearly defined for a product to enable regulatory evaluation as well as to allow for a thorough market and competition analysis. Furthermore, the clinical ease of use should be considered, especially with respect to competing and established technologies. Additional commercial considerations include manufacturing in compliance with good manufacturing practice, sterility and purity issues, quality control, and scale-up, each of which can be a barrier and an added expense, especially for complex technologies incorporating multiple components or relying on complex manufacturing procedures. Because tissue engineering products may incorporate cells or biologically active factors, logistical issues regarding distribution and shipping of the product can present considerable challenges (96). Clinicians must also adopt the product, and reimbursement strategies should be established to ensure payment. These issues should be considered at the earliest stages of product development to establish that a product has suitable potential to be commercially viable.

Future Directions

The complexity of many tissues and organs targeted for tissue engineering therapies, coupled with confounding factors associated with the clinical context, adds up to many barriers to product development and translation. Indeed, the field of tissue engineering will continue to focus on repairing complex tissues because clinical defects often involve more than one type of tissue. Characterization and consideration of the contributions of other tissue components that are often overlooked in tissue engineering, such as the lymphatics, will present a unique challenge in the development of strategies for composite tissue repair. Additional challenges include developing a deeper understanding of the role of the health status of the patient and the host response in determining the ultimate outcome of a tissue engineering therapy. Similarly, we will need to better understand the in vivo fate of various components of tissue-engineered products, including transplanted cells and the biomaterial-based scaffold, before widespread translation is possible.

Every day in the literature, a new tissue-engineered construct that combines biomaterials, bioactive factors, and cells is described. However, before translation is possible, the complexity of these technologies must be considered carefully with respect to the regulatory pathway. Tissue engineers should remain mindful that pursuit of a complex solution could eclipse a suitable simple alternative and seek to adopt the simplest approach possible to achieve the desired results in people. Additionally, technologies presenting an insufficient market base or a complex regulatory pathway may not generate sufficient funding to drive them from preclinical development to clinical implementation.

Recently, increasing effort in the field has been invested in the collaborative development of approaches for tissue regeneration with a focus on translation to the clinic as rapidly and as safely as possible, as demonstrated by the Armed Forces Institute of Regenerative Medicine ( Some approaches seek to leverage existing clinical products or components thereof to potentially mitigate the regulatory burden for clinical translation. In other cases, staged approaches for tissue repair or regeneration are being developed (Fig. 3), which will use the body as a bioreactor to facilitate production of complex, vascularized tissues to fill defects (97, 98). At the same time, clinical development continues for technologies that are already positively affecting the lives of patients, examples of which have been highlighted throughout this review. Indeed, the capability of the field of complex tissue engineering to develop technologies to advance patient care has already been demonstrated through a variety of products and clinical successes (77, 99, 100). The impact of the field will continue to grow with the collaborative development of tissue-engineered products that present simple solutions to complex problems.

Fig. 3

Staged approach for craniofacial bone regeneration. Illustration of a staged approach being developed for tissue engineering repair of large craniofacial bone defects. (A and B) The first stage involves implantation of an alloplastic implant to maintain the bony defect space and prime the wound site for repair (97) (A), and a chamber is implanted in approximation to the periosteum of bone at a distal site to promote formation of an autologous bone flap (98) (B). (C and D) In the second stage, the space maintainer is removed (C), and the autologous bone flap is harvested and transferred to fill the defect, with microsurgical anastomosis of the pedicle to vasculature at the recipient site to enable perfusion and support the viability of the transferred bone (D).


References and Notes

  1. Funding: This work was supported by the Armed Forces Institute of Regenerative Medicine (W81XWH-08-2-0032) for research toward the development and clinical translation of tissue engineering technologies. Competing interests: The authors declare that they have no competing financial interests.
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