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1 Laboratory of Regenerative and Nano-Medicine; 2 Departments of Internal Medicine and Infectious Diseases; and 3 Department of Anesthesiology, University of Texas Medical Branch, Galveston, Texas
Correspondence and requests for reprints should be addressed to Joan E. Nichols, Ph.D., Division of Infectious Diseases and the Departments of Anesthesiology and Microbiology and Immunology, Laboratory of Regenerative and Nano-Medicine, AIDS Clinical Trials Support Laboratory, Infectious Diseases Biosafety Level 3 Laboratory, 301 University Boulevard, UTMB at Galveston, Galveston, TX 77555-0435. E-mail jnichols{at}utmb.edu
ABSTRACT
Although there has been slow progress in the engineering of the lung, recent advances in the use of stem or progenitor cells leading to the reliable production of component parts of the lung show promise for the future development of engineered lung tissue. Progress toward the goal of developing an engineered lung will only be accomplished through the parallel development of effective and functional tissue-engineered components that include both upper and lower respiratory tract as well as scaffold material suitable for use in the lung. The knowledge acquired from developing each individual component of lung will, over time, be integrated to allow for the development of larger complex organ structures. To accomplish the goal of developing engineered lung for regenerative medicine, many advances will be required in scaffold design and production, including improved biocompatibility, improved elasticity, and better control of scaffold ultrastructure and porosity. Development of new materials designed to meet the anatomic and physiologic needs of the lung must occur before we can begin to realize the goal of engineering functional lung tissue. Better understanding of factors promoting cell adhesion, migration, differentiation, and vascularization of grafts and lung regeneration as a whole is also needed. Advances in the development of mathematical models to examine the conditions that promote lung morphogenesis and tissue growth for computational investigations of tissue development will also be necessary if we are to realistically evaluate the production of lung tissue strictly from the engineering perspective. It is obvious that engineering of lung tissue will require a multidisciplinary approach if we are to eventually succeed in our attempts to produce tissues worthy of clinical application in the future.
Key Words: engineered lung • lung stem cells • matrices for lung engineering • tissue engineered lung
Lung disease remains a significant cause of morbidity and mortality in the world and can be either obstructive or restrictive in nature. Chronic obstructive pulmonary disorder (COPD) is collectively the fourth leading cause of death in the world (1). Individuals with obstructive disease have obstructions caused by anatomic narrowing of the airways or blocking of airways with mucus that interferes with normal breathing. Interstitial lung disease, also known as pulmonary fibrosis, is a restrictive disease that includes a variety of chronic lung disorders. Restrictive disease is also a serious problem and there are approximately 5 million people worldwide who are affected by pulmonary fibrosis (2). Development of fibrosis eventually leads to scarring of the lungs and causes a restrictive disease that limits normal respirations and oxygenation. Management of lung disease whether obstructive or restrictive often includes drug therapy, oxygen therapy, surgery, and pulmonary rehabilitation (chest physiotherapy). For individuals 65 years or younger, lung transplantation may also be considered an appropriate therapy. Unfortunately, many patients currently on transplantation lists will succumb to their disease before transplantation due to lack of donated lungs for replacement of the damaged tissues. The generation of lung tissue through tissue engineering raises the possibility for treatment of lung diseases/disorders and provides an answer for current organ shortages by the production of new functional replacement tissues.
Tissue engineering for regenerative medicine purposes is the reconstruction of tissue equivalents to replace the physiologic functions of tissues lost due to disease or injury. Tissue engineering requires the use of a cell source to allow for the generation and maintenance of tissue-specific biological functions as well as the use of synthetic or natural matrix materials to support and guide tissue development. Clearly, engineering a complex organ such as lung, liver, kidney, heart, or small intestine presents so many scientific challenges that development of clinically applicable replacement tissues has not yet been realized. Problems to be faced in the development of any complex tissue, including lung, depend on the development of better systems to promote angiogenesis, the selection of appropriate cell sources, the reproducible differentiation of the selected cell type or types along organ-specific lineages, and the development of appropriate scaffolds or matrices to enhance and support three-dimensional (3D) production of tissues. Because complex organs such as the lung involve more than one cell type, an understanding of the factors involved in the differentiation potential of the selected cell source is invaluable. A major problem in engineering of any tissues for clinical application is selecting human cell sources with the potential to provide sufficient numbers of cells for development of tissue used to repair critical defects caused by disease or injury beyond the repair capabilities of the human body.
Some progress has been made in the engineering of less complex tissues such as skin (epidermis) (3) and urinary bladder (4). Engineering of these tissues has already been accomplished and these materials are either in current use clinically or have at least been evaluated in human trials. For engineered skin, the primary clinical application is development of a surface cover to reduce fluid loss and prevent bacterial colonization. This has been accomplished using autologous human keratinocytes or nonautologous cell sources such as neonatal fibroblasts on a variety of natural and synthetic matrices (3). Patients with end-stage bladder disease have been treated using autologous bladder constructs for reconstruction with an omental wrap to provide for blood supply. Although the trial was small, with seven patients aged 4–19 years, urodynamic studies showed that bladder compliance was improved in all patients treated, a very promising result (4).
Engineering of more complex tissues is obviously still in the early stages of development and has not progressed beyond preliminary animal trials. Some progress has been made in the development of engineered gastrointestinal tract (5, 6) and in development of engineered vascular grafts and heart valves (7–9). Recently, myocardium was developed on de-cellularized cadaveric hearts, which could be shown to generate pump function, a central component in the development of a functional engineered heart (10).
A great deal of progress has been made in development of many of these complex tissues, but before considering the design process to engineer lung tissue replacements, it is important to understand aspects of the anatomy and physiology of the normal lung as well as that of the diseased or damaged lung. The lung itself is composed of both upper and lower respiratory tract, and the anatomy and physiology are different in these regions as are the cell types found there. There are over 40 distinct cell types in the respiratory tract, but the structure of an individual distal lung alveolus is very simple. The alveolar wall consists of a narrow connective tissue core that contains fibroblasts, myofibroblasts, capillary endothelial cells, and extracellular matrix (ECM) components, such as elastin. The alveolar epithelium itself is made up of two cell types: type I and type II pneumocytes. Alveolar type I cells are flattened cells and, together with the capillary endothelium, form the actual gas–blood exchange barrier. Alveolar type I cells comprise about 90% of the surface epithelial layer, with the rest of the surface (10%) made up of the more cuboidal type II cells. The definition of what is engineered lung can be as simple as replication of the distal lung epithelia or as complex as the development of fully functional replacement tissues that include both distal lung and branching airways.
This article presents the current progress toward development of some of the tissue components that will eventually lead to the engineering of functional replacement lung tissue.
CELL SOURCES FOR DEVELOPMENT OF ENGINEERED LUNG TISSUE
Tissue engineering of the lung has not progressed rapidly, with only a few published reports describing the development of in vitro–grown tissues or implantation of engineered tissue constructs in vivo. Although these studies have generated initial enthusiasm about the potential for lung therapies, the actual engineering of all of the component parts of lung tissue has been limited. The slow progress in this area may be due to the complexity of the tissue and the variety of cell types present in functional lung, including ciliated epithelial cells, smooth muscle cells, endothelial cells, Clara cells, and specialized pneumocytes, to name some of the most important mature lung cell lineages. There are two approaches using progenitor cell populations to promote the growth of functional complex tissue that could be used to provide for this cellular diversity: (1) the use of mixtures of unipotent somatic progenitor cells, each giving rise to an array of lung-specific single-cell lineages, and/or (2) the use of multipotent cells such embryonic stem cells or fetal lung tissue capable of differentiating into progeny with multiple differentiation phenotypes.
Several recent works have described potential sources of progenitor cells capable of generating some of the cellular components of lung tissue. These include adult-derived as well as embryonic cell sources. An unresolved issue is that growth factor requirements that drive lung progenitor and embryonic stem cells to differentiate and develop into functional tissue with high efficiency are not well understood and there has been limited success in growth of engineered lung tissues for this reason. There is an urgent need for a systematic approach leading to development of better methods to increase efficiency of differentiation of stem cells along lung lineages for the production of tissues for use in regenerative therapies.
The generation of new lung tissue from mature lung lineage cells derived from somatic lung progenitor cells is particularly appealing, because it offers the possibility of autologous therapy or in situ therapy, which minimizes the risks of graft rejection and disease transmission. This is not the only option available, although it is the least problematic one for the purpose of transplantation. There are few published reports regarding lung-derived stem cell populations, and to date the only lung-derived progenitor cell that has been used to generate lung tissue in vitro as well as in vivo is a mixed population of somatic lung progenitor cells (SLPCs) described by Cortiella and colleagues (11). Other endogenous lung cells with potential to function in the future for development of engineered lung tissues are the bronchioalveolar duct junction cells, which include the bronchioalveolar stem cells (12) and variant Clara cells (13). The bronchioalveolar stem cells in mice are a rare Sca-1+ CD45– CD31– CD34+ stem cell population found at the bronchioalveolar duct junction, which coexpress both the alveolar type II cell marker pro–surfactant protein (SP)-C and the Clara cell marker, Clara cell protein 10 (CC10) (12). Variant Clara cells were shown to contribute to distal airway epithelial repair, express CC10, and resist injury by naphthalene treatment (13). There have also been numerous reports supporting the possibility of bone marrow–derived cell engraftment in the lung (14–18) and/or differentiation into lung lineages (18–20), although these cell types have not yet been used to generate lung tissue in vitro or in vivo.
Cells derived from fetal lung tissue have been used consistently by a number of research groups to produce mature lung cell lineages. Other cell sources that have been shown to be able to differentiate into distal lung epithelial lineages include embryonic stem cells (ESCs), which have been differentiated into type II cells (21–24), and upper airway lineages, including ciliated cells and Clara cells (25). Although the use of ESCs or fetal tissue as a source of progenitor cells has produced some promising results, tissues produced from these sources would have to be human leukocyte antigen (HLA) matched or engineered to express the graft recipient's HLA repertoire to be used for human transplantation. This would require the development of fetal tissue banks and, if a perfect HLA or tissue match was not available, immunosuppressive treatment of recipients of the engineered tissues.
Currently, use of stem cell sources to engineer tissues will require considerable cell expansion, manipulation, and culture time in vitro before cell matrix construct formation or tissue formation. Because of this, evaluation of potential contamination and/or damage to tissue compliance requirements for "good manufacturing practice" and "good tissue practice" will have to be followed (26, 27). Similar systems to confirm that cell differentiation only occurs along desired lineages will also be necessary. These procedures will eventually require evaluation of the karyotypic stability of the cells and the possibility of formation of genetic alterations due solely to the manipulation of cells in vitro. The same potential for self-renewal and plasticity that makes somatic and embryonic stem cells as well as germ tissue attractive sources for production of engineered tissues also raises concern about the potential for tumorogenicity of the stem cell source, and regulatory frameworks may develop to support implantation of differentiated tissues rather than growth factor–primed cell matrix constructs. The risk that growth factor–primed and activated somatic or tissue-specific stem/progenitor cells have the capacity to give rise to cancer or can result in tumor formation is low. Reports identifying lung tissue–specific tumorogenic lung cancer stem cell populations (28) or transformed counterparts of lung-derived stem cells that have the potential to give rise to carcinomas (12) have shown that the transforming events required to initiate tumorogenic properties in stem cells comprise more than the normal effects of growth factor activation and priming typically used for the purposes of cell differentiation and induction of tissue formation. It is also important to note that there was no indication of tumor formation from implantation of in vitro–amplified, growth factor–primed SLPCs into immunodeficient mice by Cortiella and coworkers, although development of mature lung lineage cells was shown.
SUPPORT SCAFFOLDING FOR ENGINEERED LUNG
The main requirements for any scaffold for use in regenerative medicine practices are biocompatibility of the material and ability to provide for 3D development of tissues. Of critical importance in scaffold selection for development of lung tissue is the elasticity as well as the adsorption kinetics of the material used. For development of lung tissue, the scaffolding must remain long enough to provide the framework necessary to support cell growth and tissue development without impeding the elasticity or altering the elastic recoil of the engineered tissue or, due to its proximity after implantation, the adjacent normal lung tissue. If the biomaterial used is not as elastic as normal lung tissue, it will contribute to the restrictive condition similar to the disease process caused by the restrictive scar tissue formation seen in patients with idiopathic pulmonary fibrosis or sarcoidosis.
Porosity of the scaffold is also a consideration. Materials that contain interconnected micropores with geometry supporting movement of nutrients into the tissues and waste removal away from tissues are essential. The scaffold must also provide sufficient cell surface area to support cell seeding and movement of cells within the scaffold with subsequent cell attachment, while at the same time allowing for the interactions of cells in three dimensions that promote appropriate cell-to-cell signaling. The nature of complex organs such as the lung may eventually require the development of hybrid scaffolds formed from more than one material to provide all of the above scaffold requirements. Lung diseases and their subsequent pathology will influence the choice of scaffold material to be used to support replacement tissue. In the case of obstructive diseases, creation of tissues on scaffold materials that limit the size of the airway would add to the existing obstructive problems. A restrictive disease that impedes pulmonary function could be exacerbated by scaffolding that lacks the appropriate elasticity or does not degrade sufficiently.
Both natural and synthetic polymers have been used in lung tissue engineering. Natural materials include collagen (29, 32–37), Matrigel (BD Biosciences, San Jose, CA) (38, 39), Gelfoam (Pfizer, Brussels, Belgium) (40) and Englebreth-Holms tumor basement membrane (29, 41). Collagen (generally type I) is commercially available and has been used as a scaffold for engineering a variety of tissues, including lung, and a number of scaffolds based on it are available for clinical use. Matrigel, a scaffold composed of basement membrane proteins is commercially available and has also been used to culture a wide variety of cell types. Gelfoam is a compressed sponge of porcine skin gelatin and was originally developed as a hemostatic device to arrest bleeding and promote clotting. There are limitations in the use of natural scaffolds such as these due to their mechanical properties and variation in degradation rates (42). There is also the possibility that natural scaffolds may be immunogenic and invoke a severe immunologic reaction leading to inflammation or that natural materials may harbor bacteria or viruses if adequate steps have not been taken to ensure the cleanliness of the materials produced.
Synthetic polymers can be produced with a much wider range of mechanical and chemical properties than that of natural materials. To maintain the elastic nature of the environment, it is generally accepted that a degradable or an extensively biomodifiable material would be the best overall choice for engineering lung tissue. Degradable synthetic matrices that have been used to engineer lung tissue include polyglycolic acid (PGA) (43) in the form of a felt sheet, PGA combined with pluronic F-127 (PF-127) (11), poly(lactic-co-glycolic acid) (39), or poly-L-lactic-acid (39). PGA degrades by acid hydrolysis to lactic acid and glycolic acid (42). The degradation rate of PGA is controlled by the molecular weight of the polymers as well as the ratio of glycolic acid to lactic acid subunits. PGA is also a widely used polymer and is better known as the suture material Dexon (Tyco, Mansfield, MA). Polymers often referred to by the trade name Pluronic (Prill; BASF, Mount Olive, NJ) are liquids at low temperatures (below 15°C) and gel at higher temperatures such as at body temperature. The firmness or density of the gel increases as the concentration of the hydrogel is increased. Poloxamer hydrogels such as Pluronic, when used in cell culture systems, maintain cells in a 3D structure that enables them to secrete ECM and engage in cell signaling (11).
Currently, there have been few instances of in vivo use of scaffolding material in engineering of distal lung tissue. The first was done in both a large-animal model (sheep) and a small-animal model (nude mouse) by Cortiella and coworkers (11) using ovine SLPC scaffold constructs produced using a combination of PGA and/or PF-127. These constructs were implanted onto either the backs of nude mice or in sheep; autologous constructs were implanted directly into the right upper lobe of the lung into a pocket created by a wedge resection. Implanted constructs were well tolerated and tissue assembly was facilitated in vivo by the use of the synthetic polymer scaffolds. In this same study, an autologous SLPC/PGA construct was implanted into the thoracic cavity of three adult sheep with attachment of the construct to the right main stem bronchus site after a full pneumonectomy. When harvested after 3 months, the implants were not shown to support lung epithelia development but did form soft tissue fragments. Another in vivo use of scaffolds in lung engineering involved the use of Gelfoam delivered into the lung by injection of the constructs (sponge with fetal rat lung–derived cells) directly into lung parenchyma (40). The sponge was shown to degrade over several months and was also well tolerated. Most of the newly formed alveolar-like structures, however, were found close to the border between the sponge and the surrounding normal tissue with few found within the sponge itself. It is unclear why this occurred because the porosity of the Gelfoam scaffold should have provided cells with an adequate environment to support cell movement and tissue development. Mondrinos and colleagues also developed constructs using fetal pulmonary cells (FPCs) using a Matrigel scaffold (38). The FPC/Matrigel construct was injected subcutaneously into the anterior abdominal wall of a mouse. The Matrigel was shown to support both development of lung epithelia and vascularization of the construct.
Although each of these materials was shown to be adequate for the development of tissue, the degradation of the scaffold material is also an important consideration. Andrade and colleagues (40) demonstrated that the scaffold material Gelfoam was an excellent supporting material for lung cell attachment, and the timely degradation of the material left the newly formed "alveoli" in place once degradation was complete. This is also true of PGA/PF-127 in both in vitro and in vivo use of this combination matrix, which also degraded as the development of epithelial tissues progressed (11). FGF-2–loaded Matrigel produced highly vascularized tissue with few structures reminiscent of alveolar forming units (38), possibly due to either the structure of the scaffold or slow degradation of the scaffold material.
To specifically meet the needs for future production of engineered lung, we believe that development of a biodegradable, highly elastic material with shape and pore size similar to that found in the alveolar structure itself is essential. A recently identified scaffold with the potential to meet these requirements features a novel inverted colloidal crystal (ICC) geometry (44–46). Primary colloidal crystals are hexagonally packed lattices of spheres, with a wide range of diameters from nanometers to micrometers. ICCs are similarly organized structures in which the spheres are replaced with cavities, whereas the interstitial spaces are filled. When ICC cavities exceed the diameter of cells, they can be used as 3D cell scaffolds (44–46). The open geometry of the ICC lattice, high porosity (74% of free space), and large surface area make ICCs an attractive structure to support the development of alveolus formation (45, 46). ICC topology affords a simple method for biomaterial design that can allow for some degree of control over cellular interactions or cell migration by varying the sphere diameter, which is an important consideration for development of engineered lung.
EARLY IN VITRO LUNG MODELS
Early in vitro models of the lung, although simplistic, mimic some aspects of lung function. Information gained from these simple one- or two-cell–type model systems paved the way for the design and execution of later complex multicell-engineered tissue equivalents. Some of the earliest models of lung function involving organ culture of intact fetal lung showed that there were striking differences in the morphologic development of submersion-cultured versus air-interface–cultured explants of fetal lung (33). Submersion cultures of the lungs expanded and retained a discrete lobular structure, whereas fetal lung grown in an air interface tended to flatten during culture with the parenchyma taking on a more glandular appearance. The importance of culture condition (submersion or air interface) on developing lung morphology suggests that use of tissue construct submersion bioreactor culture may provide for better development of engineered distal lung tissues in the future.
Human tissue culture-derived monolayers have provided basic information related to functions of lung epithelium and endothelium. A model developed by Birkness and colleagues (47) to examine early events in Mycobacterium tuberculosis infection used a human endothelial cell line, HULEC (human lung microvascular endothelial cell line), and an epithelial cell line, A549, in a two-layer system that allowed cell-to-cell contact and provided a great deal of information regarding the influences of cell contact on cell differentiation, cell morphology, and cell orientation, which are critical issues for the development of 3D tissues. Addition of peripheral blood–derived leukocytes allowed for examination of the influence of leukocyte movement and production of inflammatory cytokines on epithelial and endothelial cell growth and survival.
Development of a human composite respiratory mucosa using normal human bronchial cells cocultured with bone marrow mesenchymal stem cells under air interface conditions allowed for the formation and maintenance of epithelial cells as well as mucus-producing cells in vitro (48). The development of this model of respiratory epithelium exhibited structural and functional features of tracheal mucosa, such as cytokeratin production and airway repair mechanisms, which was an important step in the process of developing more complex lung tissues.
Two excellent 3D models of airway mucosa used IMR-90 (human lung fibroblast) and normal human bronchial epithelial cells in a collagen scaffold to create an in vitro model of airway mucosa (30, 49, 50). Collagen II and fibronectin were shown to be produced by the model as was mucus. These models allowed for the evaluation of the role of mechanical stress on epithelial and endothelial cell development, response, and survival, and more importantly provided information related to interactions of epithelial cells and fibroblast-rich ECM. Understanding the basic interactions of epithelial cells with ECM provided a better understanding of scaffold–ECM–epithelial cell interactions necessary for the design of biomaterials to support attachment of engineered lung tissues.
Although a great deal of information regarding cell contact and rudimentary tissue formation was gathered from the study of these transformed and clonally related cell lines grown in either 2D or 3D systems, better and more accurate models will need to be developed from primary cell types or from adult or embryonic stem cell sources in the future.
DEVELOPMENT OF IN VITRO–ENGINEERED LUNG
The above primitive models provided a great deal of information related to culture of cell types found in the upper airway and eventually led to the production of more complex engineered tissues by a number of research groups. Early work to construct alveolar-like structures from primary type II cells in a 3D collagen gel showed production of surfactant protein (32). This was followed by development of engineered tissue models of pulmonary–alveolar–capillary barriers, which is the major anatomic component necessary for the development of the alveolar–capillary junctions required for support of gas exchange (35).
Early reports by Blau and coworkers showed production of type II pneumocytes in vitro using rabbit fetal lung tissue cultured on Englebreth-Holm-Swarm tumor membrane (29). The cells were shown to be cuboidal and contained lamellar bodies, a characteristic marker of type II pneumocytes, depending on the in vitro culture conditions (29). In vitro development of engineered tissues containing type II pneumocytes as well as endothelial cells from fetal lung has been accomplished (39), as has in vivo development of these cell types after implantation into animal models directly into the lung (40) or into the abdominal wall (38).
The capacity of murine ESCs to generate airway epithelia, containing both ciliated epithelial cells and Clara cells as determined by electron microscopy and immunostaining for CC10 and SP-D with validation by reverse transcriptase–polymerase chain reaction of CC10, was reported in 2005 (25). This was the first report of the ability of ESCs to produce differentiated airway epithelial tissue and was a major milestone in the development of engineered lung from a renewable stem cell source with the potential to produce all of the cell types found in the normal lung. The importance of environmental influences on cell differentiation was shown by coculture of murine ESCs with pulmonary mesenchyme isolated from distal lung, which produced a microenvironment that promoted the differentiation of pulmonary epithelium (22). Epithelium-lined channels in this study expressed cytokeratin and thyroid transcription factor-1, an early developmental marker of pulmonary epithelium. Differentiation of type II cells was demonstrated by immunostaining for pro–SP-C and the presence of mRNA for SP-C in some of the epithelial cells using reverse transcriptase–polymerase chain reaction (22).
In vitro differentiation of a mixture of ovine somatic lung progenitor cells into pulmonary epithelium on PGA or PF-127 scaffolds with expression of CC10, cytokeratin, and SP-C with validation by Western blot for CC10 and SPC was shown by Cortiella and colleagues in 2006 (11). Scanning electron microscopy of the engineered tissue demonstrated organization of the cells into identifiable pulmonary structures morphologically similar to alveoli. The most recent report of development of in vitro–engineered distal lung by Mondrinos and coworkers (36) described the differential effects of FGF-2, FGF-7, and FGF-10 on distal lung morphogenesis from primary isolates of FPCs in collagen gel. Results showed that FGF-2/7/10 induced robust budding of the epithelial structures and the formation of a uniform endothelial network parallel to these structures (Figures 1A–1L). Production of cytokeratin (Figure 1A), CD31 (endothelial cells), SP-B, and pro–SP-C (Figure 1B) as well as tropoelastin was used to confirm the development of components of distal lung. This in vitro model of the lung presented critical information related to the specific role of these growth factors in the generation and maturation of engineered lung. Proliferation was also shown to be enhanced by addition of FGF-2/7/10 (Figure 1M).
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There are few reports related to implantation of engineered tissues in vivo. One of the earliest references using implantation of cell constructs was done to examine branching morphogenesis and differentiation of pulmonary mesenchymal cells (51). In these studies, pulmonary mesenchymal cells were removed from the fetal lung and grafted onto a denuded section of the trachea to show that pulmonary mesenchyme induced the production of type II cells by the tracheal epithelium in response to the mesenchymal cells (51). This suggested that the placement of cells or cell constructs is a critical factor to consider and that, without the presence of specific growth factors to promote differentiation along a selected lineage, tissues will develop according to the microenvironmental cues surrounding them.
Implantation of a mixture of ovine SLPCs on a PGA/PF-127 scaffold into both a small-animal model (on the backs of nude mice) or of autologous SLPCs on the same scaffold into the lung of a large-animal model (sheep) was done by Cortiella and colleagues (11). Implantation onto the backs of nude mice was done to determine if the growth factors alone could be used to promote differentiation of SLPCs and formation of lung tissue at a site distant from the natural environment of the lung. Results from these experiments showed production of SP-C (Figures 2A–2F), CC10, collagen, and smooth muscle, and development of structures similar to alveoli. Validation of the immunostaining for these products was validated by Western blot (for CC10 and SP-C) as well as competition by an SP-C blocking peptide (Figures 2E–2F) with comparison to normal lung. Implantation of fluorescently labeled autologous SLPC/PGA–PF-127 constructs into a wedge resection site in the right middle lobe of an adult sheep lung developed into structures resembling alveolar tissue, but no histologic examination of the tissue was done to define the cell types present in the engineered tissue. Implantation of a similar SLPC/scaffold construct into three sheep at the site of a full pneumonectomy produced fleshy tissue pieces the largest of which was 5 x 13 cm after 3 months. Examination of these pieces of tissue showed that there was little lung epithelial cell development and no obvious development of lung morphology, but the tissue fragments were highly vascularized. This suggests that the severity of the inflammatory response due to the pneumonectomy may have influenced development of lung tissue by the cell construct.
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FUTURE OF ENGINEERED LUNG
To produce engineered lung tissue in the future for regenerative therapeutics, we will have to (1) select the cell source with the highest potential for use in transplantation (HLA matched), which may require tissue typing and HLA matching of banked cells; (2) select the best biocompatible and degradable scaffold suitable for lung development, designed to meet the needs for implantation into the lung with consideration of the type of lung disease to be treated (restrictive or obstructive); (3) determine the best combinations of growth factors and culture conditions that promote cellular differentiation and increase the efficiency of differentiation of ESCs, fetal lung–derived cells, or autologous stem cells in a manner suitable for in vitro production of 3D tissue-engineered culture; and (4) select the best conditions overall that promote 3D production of lung tissue.
The progress toward development of individual components of engineered lung to date has been encouraging, particularly in the development of numerous cell types into tissues of the upper respiratory tract with production of type II pneumocytes as well as endothelial cells. Although significant advances must be made before we will achieve engineered tissues worthy of clinical application, we have progressed to the point of in vivo implantation of engineered tissues. Only one study has shown improvement in pulmonary function tests after in vivo implantation of an adipose tissue–derived stromal cell/PGA sheet scaffold construct after lung reduction surgery in an emphysema model (44). Implantation of this construct increased both alveolar and vascular regeneration in this study and restored gas exchange and exercise tolerance. This suggests that, if provided the correct support system and necessary growth factors, autologous cells have the potential to invade the scaffold, proliferate, and aid in the reparative process. One can argue that in the future we may be able to target self-directed healing and in vivo engineering of lung using similar techniques.
Many advances will be required in scaffold design and production, including improved biocompatibility, improved elasticity, better control of scaffold ultrastructure and porosity, development of new materials to promote cell adhesion, and development of smart matrices that control release of growth factors, cell migration, and tissue development as needed during tissue assembly. Better understanding of factors promoting vascularization of grafts and lung regeneration as a whole is also needed. Finally, advances in the development of mathematical models will allow examination of mechanical strength, diffusion rates of specific growth factors, and structural and physical constraints of engineered tissues (52, 53). This suggests that development of engineered lung will take the combined efforts of multiple disciplines, including, but not limited to, engineers, stem cell specialists, cell biologists, physicians, biomaterials specialists, and mathematicians, to allow for the discoveries that will lead to the development of clinically useful tissue-engineered lung.
FOOTNOTES
Conflict of Interest Statement: J.E.N. was reimbursed by Boehringer Ingelheim for expenses after participation in the TransAtlantic Airways Conference, 2008, in Lucerne, Switzerland. She was given a 
1,500 honorarium. She was reimbursed by MedImmune for expenses in 2007 to give a lecture at the company. J.C. does not have a financial relationship with a commercial entity that has an interest in the subject of this manuscript.
(Received in original form February 29, 2008; accepted in final form March 27, 2008)
REFERENCES
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