The basic paradigm for engineering polymeric structures for cellular transplant in the lung and liver is to utilize a hydrogel or sponge structure with either adult differentiated cells or progenitor cells. In both the lung and the liver, these polymer-cell structures have shown great success in vitro. Additionally, in vivo animal studies show promising results of cellular survival and functional differentiation.

Although the liver has some regenerative ability, there is a large need for liver transplantation due to cirrhosis, drug toxicity, or congenital defect. Tissue engineering functional liver units has been a focus for many years as a potential source for liver replacement. The first examples of functional transplantable liver-like structures used prevascularized, non-degradable polyvinyl alcohol sponges to accommodate transplanted hep-atocytes with limited success [131, 132]. Modified gelatin sponges with lactobionic acid (MGLA) have also been used to support mouse hepatocyte growth [133]. Microhydrogels made of fibrinogen attached to poly(ethylene glycol) (PEG)-diacrylate side chains were used as a cell carrier for intravascular transplantation of hepatocytes in a rat model [134]. Another naturally derived material, silk fibroin, has been utilized in several tissue engineering approaches. Silk fibroin is a naturally-derived material which has high molecular weight organic polymers characterized by repetitive hydrophobic and hydrophilic peptide sequences. Silk fibroin assembles into regular structures during materials formation and can be considered as nature’s equivalent to synthetic block copolymers [135]. Due to its polymeric nature, silk fibroin has been utilized in a cryogel to support hepatocyte growth with aims for future transplantation [136]. A cryogel is a porous hydrogel formed from polymerization under low temperature conditions. The ice crystals that form allow for highly interconnected pores within the hydrogel. Preparation and assessment of a gelatin hydrogel with methacrylated crosslinking is shown in Fig. 7.1.

In the lung, similar approaches have been employed focusing on naturally-derived polymers alone or in combination with synthetic scaffolds. Gelfoam sponges are commercially available surgical devices that are derived from porcine skin gelatin (Pfizer.com). Gelfoam sponges are water-insoluble and capable of absorbing up to 45 times their weight (Pfizer). Gelfoam sponges have been utilized to grow lung organotypic structures in vitro [137, 138]. Additionally, Gelfoam sponges seeded with fetal rat lung cells were viable in the adult animal lung for up to 35 days with neovascularization apparent [139]. Other hydrogels with natural and synthetic components have been used to engineer pulmonary cell structures. The natural components of these are typically either gelatin or Matrigel. Gelatin is a low-cost, non-immunogenic natural material derived from collagen. Matrigel is a product from Corning Life Sciences that is a gelatinous extracellular matrix mixture made from mouse sarcoma cell secreted matrix. These two naturally-derived materials have been used in vitro and in vivo animal models extensively. Additionally, gelatin has been used in numerous human applications, which makes it an appealing polymer for new tissue engineered products. Gelatin in the form of a three-dimensional microbubble scaffold was used to provide the proper cellular microenvironment for differentiation of mouse pulmonary stem cells into alveolar pneumocytes [140]. Similarly, hydroxyethyl methacrylate-alginate-gelatin (HAG) hydrogels have been employed as three-dimensional structures for lung epithelial cell growth [141]. Additionally, three-dimensional structures for alveolar cell growth using matrigel hydrogel and synthetic polymer scaffolds of poly-lactic-co-glycolic acid (PLGA) and poly-l-lactic-acid (PLLA) fabricated into porous foams and nanofibrous matrices have been used successfully in vitro [142]. Engineering the cellular microenvironment for progenitor cell differentiation will likely continue to be the focus of new polymer formulations in lung and liver cell delivery.

While hydrogels or sponges have been examined for cell transplant for the parenchyma of the liver or lung, more traditional tissue engineering approaches have been utilized for some of the larger structures such as bile ducts, trachea, or bronchioles. These traditional tissue engineering approaches utilize scaffold materials made from naturally-derived and/or synthetic polymers and have been implanted with some success in vivo in both humans and in animal models.

In the liver, there have been few attempts to re-create functional liver structures using polymeric biomaterials in the context of engineering the bile duct or liver capsule. Three dimensional stacked polycarbonate membranes with hepatocytes showed functional liver units in vitro [143]. Functional liver units were also observed using micropatterned Poly(ethylene glycol)-poly(dl-lactide) and lactosylated poly(dl-lactide) electrospun fibrous mats [144]. For larger structures, like the bile duct or vascular structures, polymers may be formed into sheets and then tube structures. The bile duct was successfully replaced in an animal study using polycaprolactone and polylactic acid reinforced with polyglycolic acid fibers in a tube structure [145]. Continuing advances in synthetic polymer processing, such as three-dimensional printing and micropatterning, will improve the field of engineering larger liver structures such as the bile duct.

In the lung, tissue engineering strategies have been employed to engineer larger airway structures such as the trachea, main stem bronchi, and bronchiole. In 2008, the first transplantation of a tissue-engineered trachea in a human being was done to replace an end-staged left main bronchus with malacia in a 30-year-old woman. The implanted trachea was engineered from a decellularized cadaveric trachea seeded with autologous epithelial cells and mesenchymal stem cell-derived chondrocytes [146]. Five years post implantation, the tissue engineered trachea remained viable and patent with stenting needed in the native trachea near the implant [147]. This initial success advanced the field of clinical applications of tissue engineered airways (Fig. 7.2).

New advances and improvements on tissue engineered trachea involve synthetic and natural combination polymer strategies. Polypropylene meshes with collagen and/or poly(l-lactic-acid-co-e-caprolactone) coating have been used in an animal model for replacement of the left main stem bronchi [148, 149]. Greater success was achieved utilizing the poly (l-lactic-acid-co-e-caprolactone) coating for epithelial regeneration. Other success has been reported in a rabbit model where tissue engineered tracheas were formed from articular cartilage matrix and chondrocytes [150]. New advances in 3D printing polymers have been harnessed for applications to the airway. A half-pipe polycaprolactone 3D printed trachea seeded with mesenchymal stromal cells in a fibrin matrix was implanted into a rabbit model with initial success of regenerated epithelium [151]. Tissue engineering polymer approaches for the large airways have gained momentum due to the ease of fabrication and implantation. However, smaller airways, termed bronchioles, have been attempted for in vitro understanding of small airway diseases such as asthma with the eventual goal of a functional tissue engineered lung. For the bronchiole, a type I collagen gel seeded with fibroblasts, epithelial cells, and airway smooth muscle cells was engineered utilizing a pulsatile flow bioreactor [152]. Similarly, another study cites tubular bronchiole structures engineered using airway smooth muscle tubular structure collagen pulsatile flow [153]. The engineering of these smaller airway structures highlights the need for advances in bioreactor technology to go hand in hand with the advances in polymeric biomaterials. A true tissue engineered lung will likely utilize a combination approach of polymeric materials, the proper cell choices, and proper in vitro conditioning.

Vascularization is often the limiting factor in tissue engineering complex three dimensional tissues for lung and liver replacement. Both hepatocytes and pulmonary epithelium have high oxygen needs, so the design of polymeric structures must allow for a highly vascularized system. Furthermore, the blood supply is crucial in both of these systems to maintain functionality. The liver’s main function is to filter the blood coming from the digestive tract, and the lung’s main function is to allow for gas exchange in the blood to occur. This functionality, while seemingly obvious to include, is non-trivial when engineering three dimensional structures for replacement of liver and lung. Several strategies have been employed to improve oxygenation of cells grown in vitro including the addition of perfluorcarbons and the bubbling of oxygen through the culture media.

In addition to the flows from the vasculature, the mechanical environment of the three-dimensional construct must be taken into account. When engineering a portal triad for the liver, the shear forces in the bile duct must also be considered in the design. Furthermore in the lung, a very complex mechanical environment exists as the lung is constantly distended and relaxed. The polymeric structures used to engineer the lung must have large elastic recoil in order to withstand this repeated deformation. Due to the complex mechanical environment and vascularization structure of the lung and liver, decellularized organs may provide the best natural polymeric scaffolding material.

Decellularized organs from human cadavers or animal sources provide the structural architecture necessary for complex three dimensional tissues. Key natural polymers, proteins, and structural components remaining in decellularized organs are collagen types I–IV, elastin, fibronectin, and laminin. Depending on the mode of decellularization (as described in Sect. “Biologically Generated Polymers”), many active growth factors and cytokines may also be present in the matrix.

Lung decellularization has had successes in animal implantation [154]. Several other studies have examined effects of decellularization processes on matrix architecture and strength [155, 156]. Lung decellularization leaves the complex architecture of the lung airways and alveoli with the vascular network available for repopulation with endothelial cells (Fig. 7.3). Repopulation of decellularized lung matrix with progenitor cells and subsequent differentiation into lung phenotypes shows that the decellularized lung matrix provides a hospitable environment for regeneration [157]. Although there has been much progress in the field of decellularized lung scaffolds, it remains unclear what the proper cellular choices are for recapitulating proper lung function and air-tight gas exchange.

Liver structure/function relationship is paramount in engineering hepatic units. To retain the complex structure/function relationship, decellularized livers have also been utilized for repopulation and transplantation into in vivo models [158, 159]. Using the decellularized livers as scaffolds causes the proper cellular spatial distribution that is so difficult to achieve using other polymeric scaffold fabrication techniques. As with the lung, similar issues remain to be worked out in whole organ liver tissue engineering. These issues include clotting, cell choice, cell fate, and long term functionality of the organ.

Due to the complexity and remaining issues in engineering replacement organs for lungs and livers, “lab on a chip” approaches have been utilized to engineer assist devices. The assist devices can provide essential replacement of function that is missing even in decellularized organs. These assist devices are meant to perform the function of the lung or liver as either a bridge to transplant or as a permanent assist technology. The majority of assist devices are fabricated using soft photolithography or three dimensional printing. In the former, polydimethysiloxane (PDMS) is typically the polymer of choice because it can be poured into a photolithographic mold. The PDMS is often coated with a naturally-derived extracellular matrix protein such as collagen or fibronecin. Cells are then placed within the channels to perform function [160, 161]. With 3D printing, the polymer choice is broadened such that any polymer that can be put into solution at a viscosity matching the printer nozzle may be utilized. For the liver, polymers have been used as the structural components of assist devices as well as to perform function by capturing toxins (Fig. 7.4). In the lung, endothelial cells and ECM proteins have been added to polymeric tubes for enhanced gas exchange in extracorporeal membrane oxygenation devices [162]. There will continue to be a future push toward incorporating tissue engineered external devices to avoid coagulation and better prognosis for assist or hybrid technologies.